Electrospinning for Tissue Regeneration (Woodhead Publishing in Materials)

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Electrospinning for tissue regeneration

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-Pre.indd 1

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Related titles: 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. Biomaterials and tissue engineering in urology (ISBN 978-1-84569-402-9) Patients with urological disorders should greatly benefit from the recent improvements in the biomaterials used in devices and the prospect of effective regenerative medicine. Biomaterials and tissue engineering in urology provides a comprehensive review of this important area. The first part of the book explores the fundamentals of biomaterials and urology. Chapters in part two discuss design of devices, catheters and stents. The final group of chapters provide a thorough analysis of urological tissue engineering and regeneration in areas such as the bladder, kidney and reproductive organs. This book is a valuable resource for all those concerned with urological research. Regenerative medicine and biomaterials for the repair of connective tissues (ISBN 978-1-84569-417-3) Regenerative medicine for the repair of connective tissues is a fast moving field which generates a lot of interest. The biomaterials and biomechanics for soft tissue repair have been under-represented in the past. This book addresses this gap in the market by bringing together the natural association of cartilage, tendons and ligaments to provide a review of the different structures, biomechanics and, more importantly, provide a clear discussion of practical techniques and biomaterials which may be used to repair the connective tissues. 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 832819; tel.: +44 (0) 1223 499140 ext. 130; address: Woodhead Publishing Limited, 80 High Street, Sawston, Cambridge CB22 3HJ, UK) If you would like to receive information on forthcoming titles, please send your address details to: Francis Dodds (address, tel. and fax as above; e-mail: francis.dodds@ woodheadpublishing.com). Please confirm which subject areas you are interested in.



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Electrospinning for tissue regeneration Edited by Lucy A. Bosworth and Sandra Downes

Oxford

Cambridge

Philadelphia

New Delhi



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Published by Woodhead Publishing Limited, 80 High Street, Sawston, Cambridge CB22 3HJ, UK www.woodheadpublishing.com Woodhead Publishing, 1518 Walnut Street, Suite 1100, Philadelphia, PA 19102-3406, USA Woodhead Publishing India Private Limited, G-2, Vardaan House, 7/28 Ansari Road, Daryaganj, New Delhi – 110002, India www.woodheadpublishingindia.com First published 2011, Woodhead Publishing Limited © Woodhead Publishing Limited, 2011 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. ISBN 978-1-84569-741-9 (print) ISBN 978-0-85709-291-5 (online) The publisher’s 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 acidfree and elemental chlorine-free practices. Furthermore, the publisher ensures that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by Replika Press Pvt Ltd, India Printed by TJI Digital, Padstow, Cornwall, UK



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Contents

Contributor contact details

xi

Part I Fundamentals of electrospinning

1 3

1

Introduction to electrospinning

L. Wang and A. J. Ryan, The University of Sheffield, UK

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9

Introduction Basic concepts Morphology and structural formation Parameters Apparatus Materials Applications Future trends References

3 4 5 7 10 11 15 21 21

2

Polymer chemistry P. Christian, The University of Manchester, UK

34

2.1 2.2 2.3 2.4 2.5

Introduction Natural polymers Synthetic degradable polymers Conclusions References

34 36 41 48 49

3

The electrospinning process, conditions and control B. Robb and B. Lennox, The University of Manchester, UK

51

3.1 3.2 3.3 3.4 3.5

Introduction Solution parameters Processing parameters Ambient parameters Conclusions

51 52 56 61 64



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vi

Contents

3.6

References

65

4

Regulatory issues relating to electrospinning A. Wilson, CellData Services, UK

67

4.1 4.2 4.3 4.4 4.5

Introduction Regulation of materials in regenerative medicine Future trends Sources of further information and advice References

67 69 84 87 89

Part II Electrospinning for tissue regeneration

91

5

Bone tissue regeneration A. Bassi, J. Gough, M. Zakikhani and S. Downes, The University of Manchester, UK

93

5.1 5.2 5.3 5.4 5.5 5.6 5.7 5.8 5.9

Introduction Principles of bone biology Strategies for bone regeneration Fabrication of scaffolds for bone tissue engineering Potential materials for scaffolds Osteoporosis: a growing problem Strategies for the treatment of bone defects Conclusions and future trends References

93 94 97 97 99 101 102 105 106

6

Cartilage tissue regeneration T. Hardingham, The University of Manchester, UK

111

6.1 6.2 6.3 6.4 6.5

Introduction Culture of chondrogenic cells for implantation Electrospun nanofibre scaffolds Future trends References

111 113 119 124 124

7

Muscle tissue regeneration K. D. McKeon-Fischer and J. W. Freeman, Virginia Polytechnic Institute and State University, USA

127

7.1 7.2 7.3 7.4 7.5 7.6

Introduction to skeletal muscle Skeletal muscle injuries Mechanical properties of skeletal muscle Tissue engineering Contractile force Conductive elements

127 128 129 130 131 133



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Contents

7.7 7.8

vii

Conclusion and future trends References

140 143 148

8

Tendon tissue regeneration

L. A. Bosworth, The University of Manchester, UK

8.1 8.2 8.3 8.4 8.5 8.6 8.7 8.8 8.9 8.10

Introduction: tendon tissue Tendon structure and composition Tendon pathology Clinical need Tissue engineering Cell response to electrospun bundles Mechanical properties of electrospun bundles Conclusions and future trends Acknowledgements References

148 148 151 152 153 155 158 160 164 164

9

Nerve tissue regeneration C. Wang, H. Koh and S. Ramakrishna, National University of Singapore, Singapore and S. Liao, Nanyang Technological University, Singapore

168

9.1 9.2 9.3 9.4

Introduction Clinical problems in nerve tissue therapy Nerve tissue engineering Biomimetic nanoscaffolds for peripheral nerve regeneration Stem cell therapy with nanofibre for nerve regeneration Conclusion and perspectives References

168 169 171

9.5 9.6 9.7

181 187 192 193

10

Heart valve tissue regeneration M. Simonet, A. Driessen-Mol, F.P.T. Baaijens and C.V.C. Bouten, Eindhoven University of Technology, The Netherlands

202

10.1 10.2 10.3

Introduction Tissue to be replaced: heart valves Specific tissue requirements as a blueprint for scaffold properties Selection of scaffold material Scaffold properties to meet tissue requirements Future trends Acknowledgment References

202 203

10.4 10.5 10.6 10.7 10.8

205 211 212 217 218 218



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viii

Contents

11

S. C. Baker and J. Southgate, The University of York, UK

11.1 11.2 11.3 11.4 11.5 11.6 11.7 11.8

Bladder tissue regeneration Structural/functional properties of the bladder Bladder disease and the need for bladder substitution Electrospun and other scaffolds for bladder tissue engineering Electrospinning fit for purpose Future trends Conclusions Acknowledgement References

225 227

242

12

Tracheal tissue regeneration

F. Acocella and S. Brizzola, Università degli Studi di Milano, Italy

12.1

Anatomy of the trachea and main pathologies of surgical concern Tissue engineered trachea (TET) Electrospun biodegradable tubular tracheal scaffold Scaffold fulfilment In vitro and in vivo evaluation of the cell and tissue response Conclusions Acknowledgements References

12.2 12.3 12.4 12.5 12.6 12.7 12.8

225

228 234 237 237 238 238

242 245 252 258 263 275 276 276

13

Dental regeneration

I. U. Rehman, The University of Sheffield, UK and A. S. Khan, COMSATS Institute of Information Technology, Pakistan

13.1 13.2 13.3 13.4 13.5

Introduction Periodontal regeneration Reinforcement of dental restorations Conclusions and future trends References

280 281 284 292 293 298

14

Skin tissue regeneration

A. Subramanian, U. M. Krishnan and S. Sethuraman, SASTRA University, India

14.1 14.2 14.3

Introduction Biology of skin and wound healing Challenging problems in existing therapies

280

298 299 301



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Contents

14.4 14.5 14.6 14.7 14.8 14.9 14.10

ix

Restoring functional skin tissue Nanofibers as extracellular matrix analogue Ideal properties of scaffold Choice of biomaterial Cellular interactions on skin substitute Conclusions and future trends References

302 303 304 308 310 311 312 317

15

Wound dressings

T. R. Hayes and B. Su, The University of Bristol, UK

15.1 15.2 15.3 15.4 15.5

Introduction: wound healing Nanofibres Antimicrobial nanofibrous wound dressings Conclusions References

317 323 328 334 335

Part III Electrospinning for in vitro applications

341

16

Cell culture systems for kidney research

343

L. A. Bosworth, S. Schuler and R. Lennon, The University of Manchester, UK

16.1 16.2 16.3 16.4

Introduction Current work Electrospun materials Scanning electron microscopy of cells on electrospun scaffolds 16.5 Immunostaining of extracellular matrix proteins on electrospun scaffolds 16.6 Immunostaining of cells on electrospun scaffolds 16.7 Comparison of culture methods 16.8 Discussion and future trends 16.9 Acknowledgements 16.10 References

343 345 348

359

17

Cell culture systems for pancreatic research

J. D. D. Wan, S. Downes, M. Dunne and K. Cosgrove, The University of Manchester, UK

17.1 17.2 17.3 17.4 17.5

Introduction Min6 cell line Nes2y cells Novel scaffolds and production methods Methods

350 351 352 353 355 357 357

359 361 362 362 363



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x

Contents

17.6 17.7 17.8 17.9 17.10

Results Discussion Future trends Conclusion References

364 368 369 370 370 372

18

Cell culture systems for stem cell research

K. Meade, R. J. Holley and C. L. R. Merry, The University of Manchester, UK

18.1 18.2 18.3 18.4 18.5 18.6 18.7

Introduction Embryonic stem cells Current culture techniques 3D scaffolds Combining ES cells with electrospun scaffolds Future trends References

372 373 375 384 385 390 390

Index

397



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(* = main contact)

Editors

Chapter 2

Dr Lucy A. Bosworth and Professor Sandra Downes Materials Science Centre The University of Manchester Grosvenor Street Manchester M1 7HS UK

Dr Paul Christian 48 Winnington Lane Northwich Cheshire CW8 4DE UK

E-mail: Lucy.Bosworth@manchester. ac.uk; sandra.downes@ manchester.ac.uk

Chapter 3

Chapter 1 Dr Linge Wang and Professor Anthony J. Ryan* Department of Chemistry The University of Sheffield Dainton Building Brook Hill Sheffield S3 7HF South Yorkshire UK

E-mail: [email protected]

Mr Brendan Robb and Professor Barry Lennox* School of Electrical & Electronic Engineering The University of Manchester Sackville Street Manchester M13 9PL UK E-mail: Barry.Lennox@manchester. ac.uk




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xii


Chapter 4

Chapter 7

Mrs Alison Wilson Principal Consultant CellData Services 3 Burgate Court North Newbald York YO43 4TZ UK

Mrs Kristin D. McKeon-Fischer and Dr Joseph W. Freeman* Virginia Tech-Wake Forrest School of Biomedical Engineering and Sciences Virginia Polytechnic Institute and State University Blacksburg Virginia 24061 USA



Chapter 5 Dr Anita Bassi, Dr Julie Gough, Dr Mohsen Zakikhani and Professor Sandra Downes* Materials Science Centre The University of Manchester Grosvenor Street Manchester M1 7HS UK E-mail: sandra.downes@manchester. ac.uk

Chapter 8 Dr Lucy A. Bosworth Materials Science Centre The University of Manchester Grosvenor Street Manchester M1 7HS UK E-mail: Lucy.Bosworth@manchester. ac.uk

Chapter 9

Chapter 6 Professor Tim Hardingham Wellcome Trust Centre for CellMatrix Research Faculty of Life Sciences The University of Manchester Michael Smith Building Oxford Road Manchester M13 9PT UK E-mail: tim.hardingham@manchester. ac.uk

Ms Charlene Wang, Ms Huishan Koh and Professor Seeram Ramakrishna* National University of Singapore Singapore 117576 Singapore E-mail: [email protected]

Dr Susan Liao School of Materials Science and Engineering Nanyang Technological University Singapore 639798 Singapore



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xiii

Chapter 10

Chapter 13

Ir Marc Simonet, Dr Anita Driessen-Mol, Professor Dr Frank P.T. Baaijens*, Professor Dr Carlijn V.C. Bouten Eindhoven University of Technology Department of Biomedical Engineering PO Box 513 5600 MB Eindhoven The Netherlands

Dr Ihtesham Ur Rehman* Department of Materials Science and Engineering The Kroto Research Institute North Campus The University of Sheffield Broad Lane Sheffield S3 7HQ UK



Chapter 11

Dr Abdul Samad Khan Interdisciplinary Research Centre in Biomedical Materials COMSATS Institute of Information Technology Lahore Pakistan

Dr Simon C. Baker and Professor Jennifer Southgate* Jack Birch Unit of Molecular Carcinogenesis Department of Biology The University of York York YO10 5DD UK E-mail: [email protected]

Chapter 12 Dr Fabio Acocella* and Dr Stefano Brizzola Department of Veterinary Clinical Science Faculty of Veterinary Medicine Università degli Studi di Milano Via Celoria 10 20133 Milano Italy

Chapter 14 Mrs Anuradha Subramanian, Dr Uma Maheswari Krishnan and Dr Swaminathan Sethuraman* Director, Center of Nanotechnology & Advanced Biomaterials School of Chemical and Biotechnology SASTRA University Tamil Nadu India E-mail: [email protected]




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xiv


Chapter 15

Chapter 17

Dr Thomas R. Hayes and Dr Bo Su* Biomaterials Engineering Group School of Oral and Dental Science University of Bristol Lower Maudlin Street Bristol BS1 2LY UK

Mr Jimmy D. D. Wan and Professor Sandra Downes* Materials Science Centre The University of Manchester Grosvenor Street Manchester M1 7HS UK


Chapter 16 Dr Lucy A. Bosworth Materials Science Centre The University of Manchester Grosvenor Street Manchester M1 7HS UK Miss Stephanie Schuler and Dr Rachel Lennon* Wellcome Trust Centre for CellMatrix Research The University of Manchester Manchester M13 9PT UK E-mail: Rachel.Lennon@manchester. ac.uk

E-mail: sandra.downes@manchester. ac.uk

Professor Mark Dunne and Dr Karen Cosgrove Faculty of Life Sciences Core Technology Facility 46 Grafton Street Manchester M13 9NT UK

Chapter 18 Dr Kate Meade, Dr Rebecca J. Holley and Dr Catherine L. R. Merry* Materials Science Centre The University of Manchester Grosvenor Street Manchester M1 7HS UK E-mail: Catherine.Merry@manchester. ac.uk



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1


L. W a n g and A. J. R y a n, The University of Sheffield, UK

Abstract: Electrospinning is a simple and highly versatile method for generating ultrathin fibres (mainly polymers) with diameters ranging from a few micrometres to tens of nanometres. This technique has attracted tremendous recent interest in both academia and industry, owing to its unique ability to produce ultrafine fibres of different materials in various fibrous assemblies. In this chapter, a brief introduction to the process (principles, setup, parameters and apparatus) and associated morphology of electrospun fibres is provided. Thereafter, the materials used (synthetic and natural polymers, polymer blends, etc.) and the applications of electrospun ultrafibres (particularly in the field of tissue engineering), are reviewed and discussed. Key words: biomedical applications, electrospinning, nonwovens, polymers, tissue engineering, ultrafine fibres.

1.1

Introduction

Electrospinning, which may be considered to be a variant of the electrostatic spinning (or spraying) process, is currently the only technique that is able to produce continuous ultrafine fibres from submicrometre to nanometre diameters. The original idea of using high electric potentials to induce the formation of liquid drops can be traced back more than 100 years (Bose, 1745; Cooley, 1902; Lord Rayleigh, 1882; Morton, 1902). The first patent that described the operation of electrospinning appeared in 1934, when Formhals disclosed an apparatus for producing polymer filaments by taking advantage of the electrostatic repulsions between surface charges (Formhals, 1934). Despite these early discoveries, the procedure was not utilised commercially with any great success. In the early 1990s, several research groups (in particular, Reneker and coworkers) revived interest in this technique by demonstrating the fabrication of thin fibres from a broad range of organic polymers (Doshi and Reneker, 1995; Reneker and Chun, 1996). At this time the term ‘electrospinning’ was coined and is now widely used in the literature. In recent years, the number of publications (Fig. 1.1) in this field has grown exponentially owing to a number of factors: improvements in imaging techniques, the ability to fabricate complex scaffolds and the convergence of nanotechnology and biotechnology for the application of tissue engineering. 3 © Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-01.indd 3

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4

Electrospinning for tissue regeneration 900 All documents Articles

800

Number of publications

700 600 500 400 300 200 100 0

1999 2000 2001 2002 2003 2004 2005 2006 2007 2008 2009 Year published

1.1 Annual number of publications on electrospinning (source: ISI Web of Sciences®). Collector Liquid jet Pump

Polymer solution

High voltage power supply

V

1.2 Schematic illustration of the setup used for electrospinning ultrafine fibres.

1.2

Basic concepts

In a typical electrospinning process a high voltage is used to create an electrically charged jet of polymer solution or melt, which dries or solidifies on extrusion to leave a polymer fibre (Doshi and Reneker, 1995). Three major components are needed to complete the process (Fig. 1.2): a high voltage power supply, a capillary tube with a spinneret and a collector which is normally earthed (Huang et al., 2003). Most often the spinneret is connected to a syringe which supplies the polymer solution and the solution can be fed through the spinneret at a constant rate using a syringe pump. When a high voltage is applied, the pendant drop of polymer solution at the nozzle of the spinneret becomes statically charged and the induced charges are evenly distributed over the surface (Li and Xia, 2004b). The surface tension of the



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5

droplet would normally result in a sphere at equilibrium but it is distorted in the electric field, because charges within the droplet migrate to the surface that faces the collector. The accumulation of charge causes a protrusion to appear on the end of the droplet, distorting the droplet into a conical shape known as the Taylor cone (Reznik et al., 2004; Taylor, 1969). With increasing field strength, the repulsive electrostatic force overcomes the surface tension and a charged jet of fluid is ejected from the tip of the Taylor cone when a critical value is attained. The polymer solution is discharged as a jet which then undergoes a stretching and whipping process (a series of connected loops) (Reneker et al., 2000; Shin et al., 2001), leading to the formation of a long thin thread. As the solvent evaporates, solid polymer fibres with diameters ranging from micrometres to nanometres are formed and lay themselves on a grounded collecting metal screen or drum. Theoretical understanding of the electrospinning process has advanced greatly in the last few years and has been discussed in detail (Agarwal et al., 2009a; Greiner and Wendorff, 2007; Huang et al., 2003; Reneker et al., 2007). The process of electrostatic spraying low viscosity fluids to form droplets is well established, for example, in the processing of paints and emulsions, formation of dispersions and aerosols and in producing lacquers from dilute solutions of film-forming polymers. Because the forces generated by surface tension in a solvent are larger than those required for viscous flow, the Taylor cone ejects individual droplets or a jet that breaks up into a regular stream of droplets through Rayleigh instability. As the viscosity of the fluid increases, the balance between surface tension and viscous flow favours the formation of a cylindrical jet rather than breaking up into droplets. Given sufficient time a fluid will always break up, but polymer solutions provide two mechanisms to stabilise the jet. A stretched polymer solution will strain harden because of the entanglements, increasing the viscosity further, and solvent evaporation will also increase viscosity by evaporation, with the structure being eventually frozen either by vitrification or crystallisation.

1.3

Morphology and structural formation

Electrospun fibres can be assembled into three dimensional (3D) porous, random nonwoven mats as a result of a bending instability in the spinning jet. In the work that caused a resurgence of interest in electrospinning a small number of different polymers were electrospun in the laboratory (Reneker and Chun, 1996). These electrospun fibres normally showed a cylinder shape with a smooth surface. However, as the technique developed (by adjusting apparatus and choosing different parameters, which will be discussed in following paragraphs), ultrafibres made of natural polymers, polymer blends, nanoparticles- or drug-impregnated polymers, and so on,



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6


have been successfully demonstrated. Different fibre morphologies have also been produced (Fig. 1.3), such as ribbons, beaded, aligned, porous, core-shell (Ramakrishna et al., 2006) and in most cases the underlying mechanism for the morphology can be rationalised. Technically, almost any soluble polymer with a sufficiently high molecular weight can be electrospun, even highly branched copolymers (L Wang et al., 2006). During electrospinning from solution, structure formation within ultrafibres is controlled by the simultaneous processes of evaporation of the solvent and extreme elongation of the solidifying fibres (Reneker et al., 2000). Since the solvent is removed from the polymer in sub-second timescales (similar to a quench process), the molecular chains are highly ordered but generally have insufficient time to form a well-defined, microscale equilibrium structure. As a result, for crystalline or semi-crystalline polymers, there is a retardation of the crystallisation process (Dersch et al., 2003; Y Li et al., 2006; Zong et al., 2002). However, using thermal (Kalar et al.,

(a)

(b)

10 µm

10 µm

(c)

(d)

2 µm

10 µm

20 µm

(e)

(f) 100 nm

5 µm

500 nm

1.3 SEM (Scanning electron microscope) images of different electrospun fibre morphologies. (a) random (smooth), (b) ribbon (L Wang et al., 2006) (reproduced with permission from Wiley-VCH), (c) aligned (porous), (d) beaded, (e) core-shell, (Zussman et al., 2006) (reproduced with permission from Wiley-VCH) and (f) hollow (Li and Xia, 2004a) (reproduced with permission from American Chemical Society).



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7

2006; Ma et al., 2006) or solvent–vapour (L Wang et al., 2007) annealing treatments, highly orientated polymer chains can be relaxed and an ordered structure of microdomains are obtained within the electrospun fibres (Fig. 1.4).

1.4

Parameters

The morphology and diameter of electrospun fibres are dependent on a number of parameters that include: (a) substance parameters (Bognitzki et al., 2001; Demir et al., 2002; Doshi and Reneker 1995; Theron et al., 2004), including the properties of polymer itself and the intrinsic solution properties such as viscosity (or concentration), conductivity and surface tension; and (b) process parameters (Deitzel et al., 2001a; Theron et al., 2004), including operational conditions (such as the applied electrical potential and the distance between tip and collector) and ambient parameters (such as temperature and humidity) (Huang et al., 2003; Li and Xia, 2004b).

1.4.1 Substance parameters Generally the polymer solution must have a concentration high enough to have a sufficient number of polymer entanglements, yet not so high that the viscosity prevents sufficient polymer flow being induced by the (lowpressure) pump and sufficient stretching being induced by the electrical field (Martins et al., 2008). The solution must also have a low enough surface tension, a high enough charge density and be sufficiently viscous to prevent the jet from coalescing into droplets (via the Rayleigh instability) before the solvent has evaporated (Doshi and Reneker, 1995). It has been recognised that the intrinsic properties of the polymer such as molecular weight (Fong et al., 1999), molecular weight distribution and architecture (branched, linear, etc.) (L Wang et al., 2006) of the polymer could affect the concentration range that is suitable for electrospinning fibres. For instance, in electrospinning of poly(ethylene oxide) (PEO) in aqueous solution at low concentration, beaded fibres are often found (Fong et al., 1999) owing to the low solution viscosity and the relative high surface tension (as well as there being insufficient force to stretch the polymer jet). With increasing viscosity (proportional to the concentration), the beaded fibres disappear and are replaced by smooth cylindrical fibres, and an even higher viscosity results in a larger fibre diameter. Moreover, considerably increasing the conductivity, by addition of salts (Demir et al., 2002; Fong et al., 1999; Theron et al., 2004) or drugs (Jing et al., 2003; Taepaiboon et al., 2006; M Wang et al., 2007), or increasing the polarity of a solvent mixture (Reneker et al., 2000; Shin et al., 2001; L Wang et al., 2007) (Fig. 1.5) increases the net charge density and results in a marked increase of Coulombic repulsion



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0.01

0.02

0.03 q(Å–1)

0.04

70

Time (min)

Intensity (A.U)

60 min

30 min

12 min

6 min

2 min

Before

(c)

(b)

Pmma-pdeapmma

n An

ea

l

Annealed fibre or gel

Fibre

30 µm Annealed 12 min


Electrospinning (quench)

F an urth ne er ali ng

Solution


30 µm Before anneal

1.4 SAXS (Small angle X-ray scattering) data (a) and SEM micrograph images (b) of the tetrahydrofuran (THF) annealing process (at 20°C) of electrospun fibres processed from a poly(methyl methacrylate)-block-poly[2-(diethylamino)ethyl methacrylate]-block-poly(methyl methacrylate) (PMMA-b-PDEA-b-PMMA)/THF solution (35 wt.%). The time-resolved SAXS experiments were performed at the ESRF and the intensity axis is presented in logarithmic scale by arbitrary units. (c) Schematic diagram of the block copolymer macromolecular arrangement after electrospinning and annealing (solution, electrospun fibre and gel) (L Wang et al., 2007) (reproduced with permission from Wiley-VCH).

30

10

0

(a)

Fibre Gel Solution


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Introduction to electrospinning (b)

(a)

10 µm

10 µm (d)

(c)

10 µm

10 µm (f)

10 µm

10 µm 160

Average fibre diameter Solution conductivity

140

5.0 4.5 Average fibre diameter (µm)

(e)

Solution conductivity (µs cm–1)

9

4.0

120

3.5

100

3.0

80

2.5

60

2.0

40

1.5

20

1.0

0

0.5 0

10 20 30 40 50 60 70 80 90 100 % THF (g)

1.5 SEM micrographs of PMMA-b-PDEA-b-PMMA fibres, electrospun from 35 wt.% copolymer solutions with solvent mixtures of THF/DMF (dimethylformamide) at (a) 100/0, (b) 90/10, (c) 70/30, (d) 40/60, (e) 10/90, (f) 0/100 and (g) the average fibre diameter and the solution conductivity vs. the % THF for the 35 wt.% copolymer solution. A 0% THF solution indicates a 100% DMF polymer solution. (L Wang et al., 2007) (reproduced with permission from Wiley-VCH)



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10


and electrostatic forces, which cause the charged jet to be more stretched or extended and to form thinner fibres.

1.4.2 Process parameters The structure and morphology of electrospun fibres is also affected by the applied electrical potential and the distance between tip and collector because these parameters directly affect the deposition time, evaporation rate and whipping or instability regions. In general, for a given solution viscosity and polarity, a higher electrical potential ejects more fluid in a jet, resulting in a larger fibre diameter (Demir et al., 2002) whereas a shorter tip–collector distance tends to produce wetter fibres and thus beaded structures. Evaporation rate affects the fibre formation process because the loss of solvent increases the viscosity (exponentially as the glass transition temperature, Tg, is approached) during spinning reducing the tendency to form beads via a Rayleigh instability. Aqueous polymer solutions require longer distances to form dry cylindrical fibres than systems that use highly volatile organic solvents (Martins et al., 2008). Ambient temperature and humidity are also very important process parameters, owing to their influence on the solvent evaporation process and the resultant fibre morphology. For instance, the use of volatile solvents such as dichloromethane for the electrospinning of poly- l lactide (PLLA) yielded polymer fibres with a regular pore structure on the fibre surface (Fig. 1.3(c)), generated by rapid phase separation into polymer-rich and polymer-poor regions (Bognitzki et al., 2001). Moreover, electrospinning in a very humid environment is another route to porous fibres (Casper et al., 2004; Megelski et al., 2002) where tiny droplets of water precipitate onto the jet and generate phase separation. These droplets then form pores in the solidified fibre after solvent evaporation. The extent of pore formation and the pore size can be tuned by varying the humidity.

1.5

Apparatus

Great variation in fibre assemblies and morphologies can be effected through the design and construction of the electrospinning apparatus (or setups) and, obviously, the key to making reproducible fibres and assemblies is to control the spinning environment. In many cases the purpose of modifications to the process is to improve control or tailor the process to suit the needs of specific materials and applications. There are three main categories of modification to the apparatus for electrospinning: the spinneret, manipulation of the electric field (controlling



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fibre deposition) and the collector (Fig. 1.6). Recent developments in electrospinning apparatus have been discussed in several good reviews (Greiner and Wendorff, 2007; Huang et al., 2003; Teo and Ramakrishna, 2006). By modifying the spinneret design, different properties can be introduced into the ultrafine fibre. Coaxial (or dual-capillary) spinneret design (Larsen et al., 2003; Li and Xia, 2004a; Li et al., 2005; Sun et al., 2003; M Wang et al., 2006; Yu et al., 2004; Zhang et al., 2004) has been utilised by various researchers either to protect or to exhibit functionalising agents or to electrospin material that cannot otherwise be electrospun perhaps owing to high surface tension or low molecular weight. For commercial activities in the fibre production and utilisation industries, one important consideration is the rate of production of fibre assemblies. Much work has been done on the use of multiple spinnerets (Ding et al., 2004; Gupta and Wilkes, 2003; G Kim et al., 2006) for fabricating fibres. Another method to obtain a highrate production of fibres is using a needleless spinning setup (Dosunmu et al., 2006; Jirsak 2005; Yarin and Zussman, 2004), which ensures constant renewal of the solution surface without the problem of needle clogging or droplet setting and enables spinning over a larger area. During the electrospinning process, the force that stretches the solution into a fine strand is the electrostatic charge applied to it using a high voltage power supply. Since the electrostatic charges are distributed along the electrospinning jet, an external electric field can be used to control the jet. To manipulate the external electric field so as to exert some control on the electrospinning jet, the shape, position and polarity of the charges applied to the auxiliary electrodes (such as electrostatic lenses) have to be considered (Buttafoco et al., 2006; Deitzel et al., 2001b; Kim, 2006; Stankus et al., 2004). Since the profile of the electric field between the tip of the spinneret and the collector has an influence on the electrospinning jet, a number of approaches (each with a different way of controlling the distribution of electric field) have been demonstrated to create aligned or patterned fibres. One very simple method uses a rotating wheel, drum or frame as a collector (Katta et al., 2004; Matthews et al., 2002; Teo and Ramakrishna, 2005; Theron et al., 2001). Another common technique uses a pair (or an array) of electrodes in parallel (Dalton et al., 2005; Li et al., 2003b, 2004c). The arrangement and the patterns formed by the fibres, however, are different throughout the fibre mesh and depend on their location in the gap.

1.6

Materials

As mentioned above, technically, almost any soluble polymer with a sufficiently high molecular weight can be electrospun. The method can be



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Plastic syringe

Power supply (–)

Power supply (+)

Power supply (+)

Coaxial jet

Motor

(e)

Syringe

Computer

(b)

Motor

Axis

Rotating disk collector

Sharpened edge

Inverted envelope cone

Envelope cone

Pendant drop

(f)

Conductive substrate

Power supply

H.V. Power source

(c)

Fibre

Needle

Polymer solution

Hot air

Earthed substrate

Si

2 µm

100 µm

Si

High voltage power supply

Collection mandrel

1.6 Schematic illustration of different electrospinning setups (a) coaxial or dual-capillary spinneret, (Li and Xia, 2004a) (reproduced with permission from American Chemical Society), (b) multiple spinnerets (Ding et al., 2004) (reproduced with permission from Elsevier), (c) needleless spinning setup (Jirsak, 2005), (d) electrostatic lens (focus rings) (Deitzel et al., 2001b) (reproduced with permission from Elsevier), (e) wheel-like rotating collector (Theron et al., 2001) (reproduced with permission from Institute of Physics) and (f) parallel strip collector and parallel aligned fibres. (Li et al., 2003b) (reproduced with permission from American Chemical Society).

Target

Biased ring

Jet

Cone

Orifice

(d)

Collector

Oil

Silica capillary

PVT/Ti(OiPr)4

Metal needle

(a)


13

applied to synthetic and natural polymers, polymer blends and polymers loaded with chromophores, nanoparticles, or active agents, as well as to metals and ceramics (Greiner and Wendorff, 2007). In the early studies of electrospinning, the technique was mainly focused on synthetic polymers such as polyethylene oxide (PEO), nylon, polyimide, polyaramid and polyaniline (Doshi and Reneker, 1995; Reneker and Chun, 1996; Srinivasan and Reneker, 1995). However, it was very quickly realised that this technique could also be used to produce ultrafine fibres of some naturally occurring polymers. Reneker and co-workers first attempted fabricating DNA fibres in the laboratory (Fang and Reneker, 1997; Reneker and Chun, 1996). In less than a ten-year period, hundreds of publications have been published in this field and more than 100 different polymers have been successfully electrospun into ultrafine fibres using this technique (Huang et al., 2003). With an increased emphasis on potential applications in the biomedical field (drug delivery, tissue engineering and cell biology), a large number of biopolymers (biodegradable and biocompatible polymers) and modified biopolymers are being electrospun into ultrafine fibres, including PEO (Aluigi et al., 2008; Huang et al., 2001; Moroni et al., 2006; Subramanian et al., 2005; Uyar and Besenbacher, 2009), poly(lactic acid) (PLA), which includes PLLA and poly(d-lactic acid) (PDLA) (Blackwood et al., 2008; Chen et al., 2007; Park et al., 2007; Xu et al., 2009), poly(glycolic acid) (PGA) (Dong et al., 2008; Park et al., 2006; You et al., 2006), poly(lactide-co-glycolide) (PLGA) (Almeria et al., 2010; Carletti et al., 2008; Lee et al., 2009; Nie et al., 2009; Nie and Wang, 2007), poly(vinyl alcohol) (PVA) (Duan et al., 2006; Koski et al., 2004; Lee et al., 2004; Sakai et al., 2010) and poly(ecaprolactone) PCL (Duling et al., 2008; Kim, 2008; Kim et al., 2010; Yang et al., 2010; Zhang et al., 2005b). Figure 1.7 shows a successful example of using PCL/collagen electrospun scaffolds in vascular reconstruction. In addition to these synthetic biopolymers, natural biopolymers, such as DNA (Bellan et al., 2006; Fang and Reneker, 1997; Luu et al., 2003; Sakai et al., 2009), collagen (Asran et al., 2010; Buttafoco et al., 2006; Matthews et al., 2002; Powell et al., 2008; Tillman et al., 2009; Venugopal et al., 2006), silk fibroin (Alessandrino et al., 2008; Gui-Bo et al., 2010; Jin et al., 2004; Lee et al., 2005; Min et al., 2004a; X H Zhang et al., 2008), chitosan (Feng et al., 2009; Geng et al., 2005; Jayakumar et al., 2010; Min et al., 2004b; Y Zhang et al., 2008), cellulose (Du and Hsieh, 2009; Ma and Ramakrishna, 2008; Son et al., 2004; Suwantong et al., 2007; Tungprapa et al., 2007; Vidinha et al., 2008; Zhao et al., 2004) and fibrinogen (McManus et al., 2007; Sindelar et al., 2006; Wnek et al., 2003) have also been successfully processed by electrospinning. Compared to synthetic polymers, naturally occurring polymers can have better biocompatibility; conversely, their process-ability is in general poor. A number of functional biopolymers are not suitable for direct electrospinning into fibres, because of their limited



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(a)

(b)

(c)

(d)

1.7 Evaluation of endothelial lined PCL/collagen scaffolds exposed to blood. Scaffolds without ((a), (b)) and with ((c), (d)) endothelial cells were bioreactor conditioned. Washed grafts were examined by haematoxylin and eosin (H&E) staining ((a), (c)) and scanning electron microscopy ((b), (d)). Adherence of particulate blood elements ((a) (arrow) and (b)) is noted on the unseeded scaffolds. The endothelial cell covered scaffolds showed minimal adherence ((c), (d)) (Tillman et al., 2009) (reproduced with permission from Elsevier).

molecular weights and/or solubilities. One of the most effective strategies for solving this problem is to make mixtures with polymers that are well suited to electrospinning, or by using special processes, for example coaxial electrospinning (D Li et al., 2004a). This is a readily feasible approach that may not only reduce potential cytotoxicity problems, by removing the need for a chemical cross-linking reagent, but also provides a well designed solution to bypass the shortcomings of synthetic and natural polymers. Indeed, the production of new biomaterials with good biocompatibility and improved mechanical and physical/chemical properties has been achieved (Jiang et al., 2004; Min et al., 2004c; Zhang et al., 2004, 2005a). Besides biopolymers, a large number of inorganic salts, inorganic and organic particles and carbon nanotubes (CNTs), can also be immobilised in polymer fibres. These composite fibres allow the fabrication of assemblies with special functionalities or of precursor fibres (Fig. 1.8). Precursor fibres are usually converted into inorganic fibres by pyrolysis (Greiner and Wendorff, 2007). Various metal oxides (such as TiO 2, ZnO, CuO,



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Introduction to electrospinning (a)

1 µm

15

(b)

50 nm

200 nm (c)

(d)

440 511 311 220 400

20 nm

100 nm

1.8 Different morphologies of polymer composite fibres: (a) transmission electron microscope (TEM) image of CNTs/ polyacrylonitrile (PAN) composite fibres (Hou and Reneker, 2004) (reproduced with permission from Wiley-VCH), (b) SEM image of TiO2/HAuCl4/poly(vinyl phenol) (PVP) composite nanofibres (Li et al., 2004b) (reproduced with permission from Elsevier), (c) TEM image of NiFe2O4/PVP composite nanofibres (Li et al., 2003a) (reproduced with permission from American Institute of Physics) and (d) SEM image of PAN/Fe(Acc)3 composite nanofibres (Hou and Reneker 2004) (reproduced with permission from Wiley-VCH).

NiO, Mn3O4, Fe3O4, MoO3, NiFe2O4, BaTiO3, Gd2O3, PbS, Ag2S, CdS), inorganic compounds (such as SiO2, TiO2, SiO2/ZrO2, TiO2/SiO2, and Al2O3) and CNTs (both single-wall CNTs and multi-wall CNTs) have been produced by electrospinning with a wide range of polymers, including polyacrylonitrile (PAN), PEO, PVA, PLA, polycarbonate (PC), polyurethane (PU), polystyrene (PS) and poly(methyl methacrylate) (PMMA) (Greiner and Wendorff, 2007; Iskandar, 2009; Liu et al., 2009; Mieszawska et al., 2007).

1.7

Applications

Electrospinning has the unique ability to produce ultrafine fibres of different materials in various fibrous assemblies. Owing to their submicrometre size,



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electrospun fibres are able to form a highly porous mesh and their large surface area to volume ratio improves performance for many applications.

1.7.1 Filters and textiles High porosity, interconnectivity, microscale interstitial space and a large surface area to volume ratio mean that nonwoven electrospun fibre meshes are an excellent material for membrane preparations and uses such as filters (Ahn et al., 2006; Aussawasathien et al., 2008; Gibson et al., 2001; Leung et al., 2010; Rangarajan et al., 1999; Tovmash et al., 2005; Vetcher et al., 2008) and textiles (Deitzel et al., 2001a; Han et al., 2006; Ko and Yang, 2008; Lee, 2009; Schreuder-Gibson et al., 2002; Stegmaier et al., 2007). Electrospun nanofibres can form an effective size exclusion membrane for particulate removal from wastewater. The filtration efficiency of a nanofibre membrane has been studied (Gibson et al., 2001) finding that the nanofibrous membrane was extremely effective for the removal (~100% rejection) of airborne particles with diameters between 1 mm and 5 mm by both physical trapping and adsorption. Lee (2009) investigated the application of zinc oxide nanoparticles to polypropylene nonwoven fabrics via electrospinning in the development of UV-protective materials. Layered fabric systems with electrospun zinc oxide composite fibre webs were developed, at various concentrations of zinc oxide, in a range of web area densities. It was found that a very thin layer of electrospun zinc oxide composite fibres significantly increased the UV blocking for both UV-A and UV-B ranges, and exhibited an ultraviolet protection factor (UPF) of greater than 40, indicating excellent UV protection. Increasing the electrospun web area density of the zinc oxide nanocomposite fibre web also enhanced the UV-protective properties of layered fabric systems. Even though electrospun fibrous membranes can exhibit great advantages over conventional media in environmental applications, such as air and water filtration, there still remain many challenges in their applications as filters and textiles, such as the low mass production rates for high quality ultrafibres and ultrafibre-based composites and the selection of suitable materials and appropriate chemistry to introduce the desired functionality to meet specific needs. None of these challenges are trivial but they are also not insurmountable. For example, a large scale electrospinning setup is shown in Fig. 1.9. New approaches in structure manipulation (such as core-shell nanofibres and bi-/multi-component nanofibres) and modification of nanofibre surfaces (chemical grafting and plasma treatment) have been rapidly demonstrated. These new processes and new chemistries may all be incorporated in the fabrication of better (e.g. higher flux, more efficient and stronger) filtration membranes and texture (Yoon et al., 2008).



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(a)

17

(b)

1.9 Large-scale electrospinning setup (a) demonstrating technical textile sheets being coated by continuous electrospinning (b) (Agarwal et al., 2008) (reproduced with permission from Elsevier).

1.7.2 Medical applications (drug delivery and tissue engineering) In the field of medical usage, drug delivery and tissue engineering are the two major areas of established work and are the focus of on-going research into electrospinning biopolymers fibres. Kenawy and coworkers (Kenawy et al., 2002) explored electrospun mats of PLA, poly(ethylene-co-vinyl acetate) (PEVA) and their blends for use as drug delivery systems. It was found that the drug-loaded electrospun mats gave relatively smooth release of tetracycline hydrochloride, which was used as a model drug. The total percentage of drug released from the electrospun fibre mats was higher than that from the as-cast films, owing to their much higher surface area. Since tetracycline hydrochloride is a water-soluble drug, a burst release of drug was observed in the first several hours. Similar findings (in vitro based experiments) were reported from a number of works on electrospun polymer/ hydrophilic or water-soluble drug composite fibre mats, such as the composite fibre mats of PDLA and PLLA/Mefoxin® (an antibiotic drug) (Zong et al., 2002), PLGA and poly(d,l-lactide)-poly(ethylene glycol) (PDLA-PEG) block copolymers/DNA (Luu et al., 2003), PLLA/doxorubicin hydrochloride (Zeng et al., 2005b), PLGA/Mefoxin® (Kim et al., 2004) and PVA/sodium salicylate (Taepaiboon et al., 2006). In contrast, the burst release of drug was rarely observed in drug delivery systems comprising electrospun polymer/ hydrophobic drugs or in composite fibre mats comprising drugs with poor water solubility, for example PLLA/rifampicin (an antibiotic drug) (Zeng et al., 2003), PLLA/paclitaxel (a promising anti-tumour agent) (Zeng et al., 2005b) and PVA/diclofenac sodium, naproxen and indomethacin (Taepaiboon et al., 2006). Although different types (hydrophilic and hydrophobic) of drugs were used in the above electrospun polymer/drug composite fibre mats, only



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water-insoluble polymers (Kenawy et al., 2002; Kim et al., 2004; Luu et al., 2003; Zeng et al., 2003, 2005b; Zong et al., 2002) were used in those early studies with the release mechanism being that of erosion following polymer degradation. In addition to low molecular weight drugs, macromolecules such as proteins (Chew et al., 2005; Jiang et al., 2005, 2006; Maretschek et al., 2008; Yang et al., 2008), enzymes (Chen et al., 2009b; Dror et al., 2008; Lee et al., 2005; Wang et al., 2009; Zeng et al., 2005a) and DNA (Bellan et al., 2006; Luu et al., 2003; Nie and Wang 2007; Sakai et al., 2009) have also been encapsulated or immobilised in ultrafine fibres to fabricate release systems. For instance, a study of the incorporation of DNA plasmids into PLA-b-PEG-b-PLA block copolymers, and their subsequent release, found that the released DNA was still fully functional (Luu et al., 2003). Some specific applications in drug release, such as tumour therapy (Jing et al., 2003; Xu et al., 2005; Zeng et al., 2005b) and inhalation therapy (Melaiye et al., 2005) have also been explored. Tissue engineering (TE) uses scaffolds to provide temporary support for cells to regenerate new extracellular matrix (ECM) that has been destroyed by disease, injury, or congenital defects (Agarwal et al., 2009b). Electrospinning provides a loosely connected 3D porous structure with a high surface area which can mimic the ECM structure and therefore makes itself an excellent candidate for use in tissue engineering (Agarwal et al., 2008). The use of electrospun fibres and fibre mats in TE applications often involves the optimisation over several considerations, including choice of material, fibre orientation, porosity, surface modification and tissue application. The requirements for a material to be used for tissue engineering purposes are biocompatibility and biodegradability, as the scaffold should degrade with time and be replaced by newly regenerated tissues. After fabrication the surface of the scaffold can be modified with a high density of bioactive molecules owing to the relatively high scaffold surface area. The scaffold architecture is also very important and affects cell binding (Fig. 1.10). Cells binding to scaffolds with microscale architectures flatten and spread as if cultured on flat surfaces, whereas scaffolds with nanoscale architecture have a greater surface area for absorbing proteins and present more binding sites to cell membrane receptors (Stevens and George, 2005). The adsorbed proteins can further change their conformation, exposing additional binding sites, and these are expected to provide an edge over microscale architectures for tissue regeneration applications (Stevens and George 2005). A proper choice of biomaterials is required in terms of mechanical properties and degradation time, which depends upon the type of scaffold required, the type of the tissue to be regenerated and their regeneration time (Bhattacharyya et al., 2009; Boland et al., 2001; He et al., 2005; Kidoaki et al., 2005; Liao et al., 2006; Min et al., 2004a; Prabhakaran et al., 2009; Riboldi et al., 2005; Xie et al., 2010; Yoshimoto et al., 2003). Biocompatible and biodegradable



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Cell binding

Scaffold architechture

Introduction to electrospinning Microfibre scaffold

19

Nanofibre scaffold

Cell Cell

1.10 Schematic illustration of scaffold architecture and how it affects cell binding and spreading.

1 µm

10 µm

1.11 PLA/collagen electrospun fibres and growth of mesenchymal stem cells with a subsequent differentiation towards an osteo-lineage after 22 days in culture (Agarwal et al., 2009b) (reproduced with permission from Wiley-VCH).

synthetic polymers (such as PLA, PGA, PCL) and their copolymers in various compositions and segmented PU and poly(phosphazenes) as well as natural polymers (such as collagens, gelatin, chitosans, silks and alginates) are used as the carrier materials. Furthermore synthetic polymers such as PLA or PCL mixed with, coated with, or grafted with natural polymers such as gelatin or chitosan have been used to make scaffolds (Ma et al., 2005; Zhang et al., 2005a). The aim of such modifications is to provide surfaces with enhanced adhesion and proliferation capability for cells (Fig. 1.11). Owing



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to the flexibility of material selection, as well as the ability to control the scaffold properties, electrospun scaffolds have been employed in a number of different human tissue or organ applications including skin (Kumbar et al., 2008; Pawlowski et al., 2003; Powell et al., 2008; Powell and Boyce, 2009; Sun et al., 2005; Yang et al., 2009b; Zhu et al., 2008), vascular grafts (Drilling et al., 2009; He et al., 2009; Lee et al., 2007; Stitzel et al., 2006; Xu et al., 2004; Y Zhang et al., 2008; Zhu et al., 2010), bone (Jin et al., 2004; C M Li et al., 2006; Mohammadi et al., 2007; Ngiam et al., 2009; Sombatmankhong et al., 2007; Sui et al., 2007; Yu et al., 2009), neural fibre (Lee et al., 2009; Xie et al., 2010; F Yang et al., 2004, 2005) and tendons or ligaments (Chen et al., 2009a; Sahoo et al., 2006).

1.7.3 Composites and templates One of the most important applications of traditional (micro-size) fibres, especially engineering fibres, is their use as reinforcements in composites. With these reinforcements, the composite materials can provide superior structural properties such as high modulus and strength to weight ratios, which generally cannot be achieved by other monolithic engineered materials alone (Huang et al., 2003). This methodology can also be applied to electrospun nanofibres for medical applications. Fibrous polymer structures can be applied as the supporting matrix or template, and hence superior structural properties in nanocomposites can be anticipated. A variety of functional components (e.g. nanoparticles, nanowires (in particular, CNTs), or molecular species) (Chew et al., 2007; Fong et al., 2002; Ge et al., 2004; Guan et al., 2003; Kim et al., 2005; Madhugiri et al., 2003; Saquing et al., 2009; Y Z Wang et al., 2007; X H Yang et al., 2004) can be directly added to the solution for electrospinning to obtain nanofibres with a diversified range of compositions and well-defined functionalities (Li and Xia, 2004b). Meanwhile, electrospun polymer fibres can be used as templates in the preparation of hollow fibres to produce nanofibres (Czaplewski et al., 2004; Dong et al., 2003; Li and Xia, 2004a; Liu, 2004; Pantojas et al., 2008; Wan, 2008; Zhou et al., 2010).

1.7.4 Others Electrospun fibres are also being explored for use in many other functional applications, such as micro/nano electronic devices, sensors and catalysis, all of which could have useful biomedical applications. Nanofibres from polymers with piezoelectric effects will make the resultant nanofibrous devices piezoelectric (Chang et al., 2010; Chen et al., 2010; Xu et al., 2006). Electrospun polymer nanofibres could also be used in developing functional sensors, with the high surface area of nanofibres facilitating increased



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sensitivity (Aussawasathien et al., 2005; Ding et al., 2009; Liu et al., 2004; Patel et al., 2006; Song et al., 2009; Wang et al., 2002, 2010). Furthermore, catalytically active agents (both heterogeneous and homogeneous catalysis) can be dispersed in polymer nanofibres (Dersch et al., 2005; J Kim et al., 2006; Lin et al., 2009; Shan et al., 2008; Yang et al., 2009a). The highly porous structure of electrospun nonwoven fabrics, with a total porosity of up to 95%, allows an undisturbed flow of gases or fluids through the catalytic system. The nanoscale dimension of the nanofibres results in high surface areas and thus intimate contact between the components of the reaction mixture and the fibres and a short diffusion path for the reaction compounds to the catalytic centres and back into the environment surrounding the fibres (Ramakrishna et al., 2006).

1.8

Future trends

Given the versatility of electrospinning for generating highly porous nanofibre mats made from different materials, it is no surprise that it has found possible uses in different fields ranging from filters, textiles, tissue engineering, drug delivery, catalysis, sensors, and so on. It is still expected that research into electrospinning will become more interdisciplinary in the near future. In particular regard to its application in tissue engineering, further research is required to elucidate the influence of nanofibres on the biochemical pathways and cellular signalling mechanisms that regulate cell morphology, growth, proliferation, differentiation, motility and genotype. Insight into how the natural ECM components secreted by cells replace the biodegradable polymeric scaffolds is also needed. This complete understanding of cell–scaffold interactions will pave the way for successful engineering of various tissues and organs, such as vascular grafts, nerve, skin and bone regeneration, corneal transplants, skeletal and cardiac muscle engineering, gastrointestinal and renal/urinary replacement therapy, and even stem cell expansion and differentiation to specific cells types and organ regeneration (Ramakrishna et al., 2006). The subsequent chapters in this book are designed to explore current achievements and future opportunities in this burgeoning area.

1.9

References

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2

Polymer chemistry P. C h r i s t i a n, The University of Manchester, UK

Abstract: In this chapter we will look at the nature, preparation and properties of degradable polymers. We will discuss both natural and synthetic polymers and look at how they are either isolated or prepared and also at how they degrade via both biotic and abiotic pathways. Key words: biotic, abiotic, degradation mechanism, natural polymers, synthetic polymers.

2.1

Introduction

2.1.1 Polymers Polymers or so-called macromolecules are large molecular structures containing one or more units which are repeated within the molecule. The term ‘polymer’ has its origin in Greek and derived from the term ‘poly’ meaning many and ‘mer’ meaning unit. Polymers as molecules have pervaded life much longer than the synthetic materials we often associate with the word and biology has made prolific use of their diverse and unique properties. Their emergence as synthetic materials dates back to 1907 when Dr. Leo Baekeland prepared Bakelite from formaldehyde and phenol, perhaps the first fully synthetic polymer. Later the emergence of polyesters and polyethylene revolutionised life as a range of new, hardy, mouldable polymers allowed a rapid explosion in art, communication and fashion. The ability to form complex shapes while retaining their properties with little degradation made them excellent materials for the new industrial age. The term ‘plastic’ has become ubiquitous with the term ‘polymer’ although ‘plastic’ refers to the mechanical properties of some polymers, especially in the bulk form. However, the term polymer is far more encompassing when considering molecular materials in general. Polymers are usually prepared by a reaction between identical or similar molecules in order to produce a theoretically infinite chain (Fig. 2.1). In reality the reactions usually result in a range of long chains as the reactions themselves are hindered by molecular motions and side reactions. There are several excellent books on polymer chemistry and reactions and we will not discuss this in detail here.

34 © Woodhead Publishing Limited, 2011


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35

n

Free radical initiator (AIBN)

(a) O

O Cl

Cl

N H

O H 2N

n

n HCl

NH2 (b)

2.1 (a) Formation of polystyrene by free radical polymerisation. (b) Formation of Nylon 6,6 by condensation polymerisation. n denotes that the chains are long but polymerisation results in a range of chain lengths and molecular weight cannot be defined as a single number but is usually represented as a weight and number average molecular weight.

2.1.2 Degradable polymers The preparation and properties of degradable polymers in general is a massive field which encompasses biomaterials, drug delivery, sustainable development and environmental impact. As a result there are a vast range of polymers, polymer blends and composite materials which could be included under this title. In order to keep this chapter to a sensible length and make it as applicable as possible to the general theme of the book we will take time to look at the main degradable polymers. The concept of a degradable material is, however, a diverse one and the definition may depend on the perspective of the application and of one’s understanding of degradation. Two types of degradation often discussed in the literature are biotic and abiotic degradation. Biotic degradation relies specifically on the action of enzymes and organisms such as bacteria to break down the material. The processes can be complex and depend on a complex mix of extracellular processes, the diversity of the microbe community and, of course, the health of the microbial community. Abiotic degradation relies on processes which are not related to enzymic or bacterial processes. These may rely on a range of processes including photolysis or thermal ageing but is generally centred on hydrolysis reactions. In addition to this, the concept of degradation may be purely related to the loss of a specific property of a polymer without any resulting mass change, a mass loss of a polymer as it degrades into smaller fragments of polymer chain, or complete degradation to water and carbon



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dioxide or methane. Loss of mechanical properties may be associated with photo or thermal ageing of a material, whereas break up of the polymer chains is a chemical breakdown of the polymer molecules themselves. Degradation of a polymer into carbon dioxide/methane and water is usually associated with biotic degradation, the by-products being formed as a result of the respiration process as small fragments of the polymer chain act as a hydrocarbon source for the organism. There is also an issue over the use of the word ‘degradable’. Essentially all polymers are degradable, as few will survive high temperature treatments and so on. However, usually the term is reserved for ambient conditions. Even restricting the parameters to standard temperature and pressure, this still leaves the question of the environment in which the material is left and the time over which the measurement is conducted. The latter is of particular interest as this will be strongly dictated by the application to which the polymer is to be put and allows us to put limits on what is meant by degradable. This is important as it has been shown that low density polyethene (LDPE) shows significant degradation in soil, but the process takes 32 years (Otake et al., 1995). In a similar manner the degradation of a polymer for drug delivery should be over within hours and occur at the required site of delivery, whereas a degradable tissue scaffold may need to last months before it completely degrades. Therefore a degradable polymer can be said to be degradable if it is completely broken down within the lifetime of the device or article into which it has been fashioned. In this chapter we will particularly look at the degradation of two polymer types. Firstly those that are considered to be natural polymers and therefore derived from a biological source, though in most cases are altered later. Secondly we will consider synthetic polymers. In each case we will look at preparation methods for the polymers, their molecular structure, applications and their degradation. We do not have space to consider all possible materials here and therefore will focus on the most common examples of homopolymer systems (i.e. those which consist of one polymer type). In many cases copolymers and polymer blends have also been prepared with a range of degradation properties, but these will not be discussed in detail here.

2.2

Natural polymers

Biology has been making polymers for as long as life has existed. They form the basis of genetic encoding, cell structures and extracellular matrices. Their biodegradability is evident in the cycle of life. Of these polymers the two main classes in use today are proteins and polysaccharides. Silk and cotton are excellent examples of biodegradable natural polymers in everyday use. Silk is a protein-based material whereas cotton is a polysaccharide-based



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material. Both have been used by humans for millennia to great effect. Particular interest in these materials arises from their relative stability to abiotic hydrolysis compared to relatively rapid biotic degradation. This degradation process is important for a plant where life generally involves water-based chemistry.

2.2.1 Sources and structure of natural polymers Natural polymers are generally extracted from organisms. These are often plants or bacteria. The challenge of many of these is to obtain a commercially significant amount of material from the organism to make its breeding worthwhile. This has often given rise to an interest in bacterial synthesis where the organism may be grown in a broth under easily controlled conditions. The second challenge is purification of the material from the other components of the system. Clearly this final issue is of particular concern if a natural polymer is to be used as a biomaterial. Both the isolation and purification of these polymers varies greatly and will not be discussed in detail here. Proteins All proteins may be classed as biodegradable. Life relies on the ability to break down and reuse the amino acids in proteins in order to construct cells, enzymes and many other building blocks of complex organisms. Their structures are complex, not just in their chemical make up but also in their exact structure. The folding and coiling of proteins into specific geometries affects their properties and their function. The thermal disruption of these complex structures usually spells the end of life at elevated temperatures for most organisms. The degradation of proteins is left to specific enzymes and simple chemical degradation is generally slow without enzymic intervention. With this in mind we will look briefly at the chemistry and properties of silk as an example protein but would refer the reader to more in depth discussions on the more common protein-based materials such as collagen (Friess, 1998) and fibrin (Gaffney, 2001). Silk is a protein, based to a large extent on the fragment Gly-Ala-GlyAla-Gly-[Ser-Gly-(Ala-Gly)n]8-Ser-Gly-Ala-Ala-Gly-Try (Lucas et al., 1962). The material has a high tensile strength and excellent mechanical properties which make it especially interesting. The high performance of silk is thought to be a result of the folding of the proteins into large numbers of ordered sheets. Silk itself is obtained from the cocoons of the mulberry silkworm (Bombyx mori) and is generally obtained by unravelling the fibres which enwrap the cocoon. The larvae themselves are now farmed. The ease of obtaining silk with little processing explains its prevalence since ancient times.



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Polysaccharides Polysaccharides form a massive class of natural polymers, often modified for use. They form some of the first polymeric materials used by humans in manufacturing, together with nitrocellulose which was once used to make film and gun cotton. The polysaccharide class of natural polymers is based on polymers with repeat units based on sugars. Figure 2.2 shows the ring structures of the hexose range of sugars which form the basic structure of many polysaccharides. The sugar repeat units on the backbone of polysaccharides have many chiral centres (five in the case of the hexoses shown). These centres determine the exact saccharide. For example all the saccharides in Fig. 2.2 have the same nominal formula, C6H12O6, however each one is distinctly different owing to their chiral nature. In a polysaccharide the saccharide repeat units are connected via the oxygen on carbon 1 which forms a glycosidic bond to carbon 4 on another molecule with the subsequent elimination of water. An example of this is given in Fig. 2.3 where the structure of sucrose is shown. There are a large number of polysaccharides being employed as degradable polymers. The most common of these is probably cellulose (Fig. 2.4) which is the most common component of cotton. Cellulose itself is usually derived from plant matter and is often modified to provide improved properties such as solubility for processing and so on. Alteration of the polymer usually involves modification of the OH groups on the backbone of the polymer using an acid anhydride. This results in various degrees of substitution to give ester functionalised celluloses (Edgar et al., 2001). Similar modifications have also been attempted with other polysaccharides (Campoccia et al., 1998). Other polysaccharides which form part of this group of degradable polymers are starch, larch gum, alginic acid, agar, carrageenan, chitin, hyaluronic acid, dextran, gellan gum and pullulan. The structures of many of these polysaccharides are complex and a full discussion is given in the literature (Mano et al., 2007). Of these, starch, cellulose and larch gum are plant derived (although cellulose may also be obtained from microbial sources), alginic acid, agar and carrageenan are obtained from algae, chitin and hyaluronic acid are obtained from animal sources or microbial sources and the others are obtained from microbes. Polyesters The structure and degradation of the naturally produced polyesters polyhydroxy butyrate (PHB) and polyhydroxyvalerate (PHV) are discussed in the next section on synthetic degradable polymers as these materials have much in common with their synthetic counterparts. These polymers were first



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Polymer chemistry H OH

H OH 6

4

5

HO

H

OH

O 2

3

H

HO 1

H

OH

H

H

H

OH

OH

O H

H OH

OH

d-Allose

H

39

d-Altrose

OH

H OH H

OH O

O

HO

HO HO

H

H

OH

HO

H

OH

H

H

d-Glucose

OH OH H

H

OH O

O

H

H H

OH

OH

H

H

OH H

H

OH d-Idose

OH OH

H

H

H

OH

OH d-Gulose

HO

OH d-Mannose

OH OH

H

H

H

OH

O

H H

OH

H

OH

HO

OH d-Galactose

H

O H

H OH

H d-Talose

2.2 Basic hexose sugar rings. The common numbering of carbon atoms is shown in the first example (d-allose).

prepared by feeding Alcaligenes eutrophus with propionic or valeric acid along with glucose (Holmes et al., 1985). Since then there have been several other species found which either produce these polymers naturally when



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Electrospinning for tissue regeneration OH O

HO

Glycosidic bond OH

HO

OH O OH

O

OH

OH Glucose unit

Fructose unit

2.3 Structure of sucrose showing the glycosidic bond.

H

H

OH

H

HO HO

H OH

H

H

O

O HO

H HO

H H

H

H

O O

HO

n

OH

OH

H O

OH H

H

HHO

2.4 Structure of cellulose.

exposed to the correct substrate or may be genetically modified to do so (Suriyamongkol et al., 2007).

2.2.2 Uses of natural polymers The use of natural polymers in the modern world is prevalent. We have a society which mixes natural and synthetic materials in a plethora of devices and artefacts. Apart from our personal biosynthesis of polymers, we commonly use cotton, silk, wood and wool in many everyday articles. Rather than discuss here these many and often interesting and diverse applications we will restrict examples to situations where biodegradability is of foremost importance. Uses of proteins Proteins such as silk have been used for some time as a suture material (Postlethwait et al., 1962). They are often highly resistant to hydrolysis owing to the stability of the amide bond but may be degraded by common proteases. Collagen has found numerous uses as a biomaterial in a wide



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range of applications and is of particular interest as a drug delivery carrier (Friess, 1998) and in materials for reconstructive surgery of soft tissue (Sajjadian et al., 2010). Uses of polysaccharides Whilst there is a large body of continuing work on blending polysaccharides with other polymers to enhance or control degradability, there is still great interest in using these materials, in a modified form, as hydrogels (Campoccia et al., 1998; Schmidt et al., 2008). These hydrogels are biocompatible and generally have excellent absorption/release properties which make them interesting materials as drug carriers.

2.2.3 Degradation of natural polymers As we have already mentioned, the degradation of biologically derived polymers usually proceeds via an enzyme-catalysed pathway. This allows the polymer to remain stable and function for long periods of time until final degradation is required. Degradation of proteins Hydrolytic degradation of proteins is unfavourable owing to the resonance of the lone pair on the nitrogen in the amide bond which results in increased stability to hydrolysis. However, proteases are able to catalyse the process efficiently and thereby allow the degradation of organic matter as well as digestion of foodstuffs and removal of dead tissue by biota. Degradation initially produces amino acids and then these will also be degraded to carbon dioxide/methane and water depending on the nature of the protein, the organism acting on it and the culture’s environment. Degradation of polysaccharides In a similar manner to proteins, the degradation of the glycosidic bond in polysaccharides is resistant to hydrolysis under normal conditions. Digestion of polysaccharides is again an enzyme-catalysed process and usually results in the organism completely digesting the sugars in order to provide energy.

2.3

Synthetic degradable polymers

Degradable synthetic polymers, unlike natural polymers, may be readily mass produced on a multi-tonne scale and have properties similar to nondegradable polymers. There are a wide range of degradable polymers and



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more are being discovered each day, however they may be divided into several classes, each with specific properties. Key to the properties of a degradable polymer is the ability to manufacture it and its components without significant degradation, as well as it having a suitable lifetime for the final material to fulfil its role. The challenge of balancing these factors drives continued research into this area. In addition to this there are a wide range of applications of degradable polymers and this often will dictate the types of polymer used. For example a degradable polymer for human implant as a slow release mechanism must not degrade into toxic or irritating components which may cause an inflammatory response.

2.3.1 Preparation of synthetic degradable polymers There are three main classes of degradable synthetic polymer which we will consider here: polyesters, polycarbonates and polyanhydrides (Fig. 2.5). In addition to these are a range of polymers which include other degradable groups including amino acids and phosphates. These are reviewed in detail in the primary literature and will not be discussed in detail here (Kamath and Park, 1993; Falco et al., 2008). Polyesters Polyesters contain the ester linkage between monomer units. This type of bond is usually, but not exclusively, formed via either a transesterification reaction or a ring opening reaction (Fig. 2.6). There are five main polyesters which are considered to be biodegradable: polyglycolic acid (PGA), polylactic acid (PLA), poly-e-caprolactone (PCL), polyhydroxybutyrate (PHB) and polyhydroxyvalerate (PHV). We have already discussed the preparation of the latter two (PHB and PHV) as both of these are biologically derived. PGA, PLA and PCL are all readily prepared via ring-opening polymerisations (Deasy et al., 1989; Gilding and Reed, 1979; Labet and Thielemans, 2009). Typically the polymers are prepared in the molten monomer using a catalyst to promote the ring opening process (Fig. 2.7). A traditional catalyst for this is tin octanoate. Owing to the hazards of the use of tin and its associated toxicity, there have been several other catalysts developed over recent years. Of particular interest in this respect is the development of catalysts to control O

O

O Ester

n

O Carbonate

O

O

n

O

O Anhydride

n

2.5 Common functional groups in degradable polymers.



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Polymer chemistry O

43

O OMe

MeO

O

O

O

O

MeOH OH

HO

n

(a) O O

O

Catalyst

O

n

(b)

2.6 (a) Preparation of polyethylene terephthalate (non-degradable) by transesterification. (b) Preparation of polycaprolactone (PCL) via ring opening polymerisation. L

L M R

L L

O

M

O

O

R

O

O L

O

O

M

O

O

R

L

2.7 Ring opening polymerisation of caprolactone where R is any alkyl group or polymer chain and L is a ligand from the metal centre. O

O

*

O

O

*

O

* O

*

O

O

O (R, R) lactide

(S, S) lactide meso-lactide

*

O

O

*

O (S, R) lactide rac-lactide

2.8 Lactide monomers with stero centres marked (*)

the chiral centres in PLA (Thomas, 2010). The lactide monomer has two stereo centres (Fig. 2.8) which may be either SS, RR or SR. The former are termed the meso-form and the latter the racemic form. Polymerisation of these monomers and their mixtures can result in a range of structural differences in the final polymer. This affects the packing of the polymer chains and hence the final properties of the polymer.



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Polycarbonates Polycarbonates are a similar class of polymer to polyesters and may be prepared by a similar ring opening polymerisation to that for the polyesters above (Fig. 2.9) (Rokicki, 2000). Examples of polycarbonate-based degradable polymers are restricted in the main to poly(trimethylene) carbonate. Polyanhydrides The polyanhydrides are perhaps some of the least hydrolytically stable degradable polymers. There are many examples which all contain the anhydride group (Fig. 2.10) as part of the repeat unit. Poly(1,3-bis-pcarboxypheloxypropane anhydride) is probably one of the best examples of a degradable polyanhydride currently in use (Kumara et al., 2002). They are generally prepared by condensation reactions (Fig. 2.11). O O O

O

O

O

2.9 Synthesis of polytrimethylene carbonate via ring opening polymerisation (ROP). O

O

O

n

2.10 An example of a polyanhydride showing the anhydride group. O

R

HO O

Cl

O

OH +

R¢

O Cl2CH2

O

O

O

O

Et3N 0°C R

O

R¢

O

n

Cl

2.11 General method for the preparation of polyanhydrides by condensation polymerisation of a dicarboxylic acid with a diacid chloride.



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2.3.2 Uses of synthetic polymers The use of synthetic degradable polymers is generally limited owing to their ability to degrade in aqueous media. Slow degrading polyesters have been used in a range of applications including packaging (Sinclair, 1996), bone replacement materials (Nelson et al., 1977), adhesives and films and coatings as well as implantable soft tissue scaffolds (Chena and Wua, 2005). The application of these polymers depends on their stiffness (modulus) and strength as well as their ability to degrade. These factors are influenced in a large part by the degree of crystallinity in the polymer, the orientation of the crystallites to the line of applied force and the temperature at which the polymer undergoes a transition from the glassy to the plastic state (the glass transition temperature or Tg). The specific use of these types of polymer as biomaterials generally focuses on polyesters and polycarbonates. These materials can offer mechanical support for the growth of new tissue for long enough periods of time for the generation of new tissue. From a practical perspective, the properties of biomaterials such as these often need to be modified further to improve biocompatibility, elicit the correct mechanical properties and modify the degradation profile of the materials. This is achieved by a range of approaches including inclusion of bioactive molecules in the polymer (Niu et al., 2009) or modification of the polymer surface (Kiss et al., 2010) in order to promote tissue growth or improve biocompatibility, blending of two or more polymers or preparation of copolymers to improve the mechanical properties and modify the degradation profile (Wang et al., 2010). The rest of this book will deal with specific applications of many of these polymers.

2.3.3 Degradation of synthetic polymers A major difference between synthetic and natural biodegradable polymers is the ability of synthetic polymers to degrade via hydrolysis rather than solely via a biotic mechanism. Clearly life could not exist if the polysaccharides, which are commonly used by biota, degraded solely on contact with water. This however does not preclude the action of biota on synthetic polymers. As we have already seen, naturally derived polymers such as PBH are polyesters which have specific roles in energy storage in some bacteria. This therefore implies that biota have developed the enzymes necessary to degrade these polymers in order to release the small molecules which are then metabolised by the organism to provide energy. We will therefore consider both processes of degradation here.



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Abiotic degradation As already mentioned the abiotic degradation of biodegradable polymers relies, for the most part, on hydrolysis of groups in the polymer chain. Degradation of these groups results in the formation of small molecules which may then be absorbed by cells and degraded further into water and carbon dioxide through normal metabolic pathways or excreted from the organism. Hydrolytic degradation of ester (Hakkarainen, 2002), anhydride (Gopferich and Tessmar , 2002) or carbonate (Zhu et al., 1991) groups is promoted by the polarity of the carbonyl bond (Fig. 2.12). The d+ polarity at the carbon of the carbonyl bond allows attack at the carbonyl by the oxygen on a water molecule which has a complimentary d– polarity. The result of the reaction is either a carboxylic acid and an alcohol, in the case of an ester, or two carboxylic acids in the case of an anhydride. The degradation of the carbonates is more complex as the hydrolysis results in the formation of an alkyl hydrogencarbonate which is not stable and readily decomposes further to an alcohol and carbon dioxide. The degradation of the polymers proceeded generally via random chain scission. This means that the hydrolysis may occur, in principle, at any point along the chain. This is turn means that the degradation of polymers by this method results in a general broadening and decreasing of the O

O

O

O H

O

O

H

H

O

O

H

H

O O

O

O OH OH

O

2

OH

O

O

OH

O H

2 Ester

H

Anhyride

OH

CO2

Carbonate

2.12 Hydrolytic degradation of ester, anhydride and carbonate linkages.



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average molecular weight of the polymer. The process of degradation may be accelerated by changes in pH. Both an increase in pH or a decrease in pH results in catalytic acceleration of the hydrolysis process, particularly in the case of polyesters. The formation of an acid as the by-product of the degradation can result in auto acceleration of the degradation process where the acid produced during degradation becomes trapped in the polymer and further accelerates the degradation. Often degradation of polymers is observed by looking at the mass loss of the polymer and it is worth pointing out that, although a polymer may lose mass during degradation, it does not necessarily follow that the associated mass loss is the result of direct conversion of polymer to monomer/smallest units. It is well known that degradation of polymers via hydrolysis can result in the elimination of many small molecules based on the polymer backbone (Vidil et al., 1995). These may take the form of more complex molecules other than linear short chains of the polymer, for example cyclic oligomers (short polymer chains) are often observed. With time these too will degrade to the monomer and be either excreted or metabolised. The rate of hydrolysis depends on two important factors, the sensitivity of the group to be hydrolysed and the ease with which water can penetrate the polymer. With regard to the sensitivity of the group to hydrolysis, anhydrides hydrolyse far more rapidly than esters or carbonates and it is for this reason they are often used as carriers of drugs for release over less than one week (Gopferich and Tessmar, 2002). However there is a considerable range in the degradation of even one class of polymers. Consider the two polyesters, polylactic acid (PLA) and polyethylene terephthalate (PET). The degradation of PLA may be counted in months in water (Lunt, 1998) and that of PET in tens of years (Kint and Munoz-Guerra, 1999). Three factors play an important role in controlling the permeability of water in the polymers, the hydrophobicity of the polymer, crystallinity and polymer chain mobility. The rate of degradation of the polymer will be restricted initially to its surface as polyesters are generally insoluble in water. The magnitude of the interaction of water with the surface of the polymer will be dictated by the structure of the polymer backbone. A polyester containing a greater proportion of hydrophobic groups will be more hydrophobic than a similar polyester containing a greater proportion of hydrophilic groups. This effect can be seen in polylactic acid which has a water–polymer contact angle (Kiss and Vargha-Butler, 1999) of about 55º, whereas PET has a water–polymer contact angle (Merrill and Pocius, 1991) of 77º (this means that PET is more hydrophobic). It should also be noted that the initial contact angle at a water–polymer interface is likely to change with time as the polymer degrades. In addition to the effect of hydrophobicity, polymer crystallites also have a role to play. Many polymers form semi-ordered crystalline regions. This will offer little free volume for



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water ingression and, therefore, degradation and will also impede diffusion of water through the structure as a whole, dependent on the size, shape and orientation of the crystallites within the polymer structure. A final restriction to water ingression is the free volume within the amorphous regions of the polymer. The free volume in a polymer changes markedly as the material moves from a plastic to a glassy state. In the glassy state the polymer chains are fairly immobile and there is little free volume. In the plastic state the polymer chains are more mobile and there is sufficient free volume to allow ingression of small molecules such as water. The point at which this change occurs is called the glass transition temperature. The general approach we have taken here is greatly simplified but provides a good overview of the area and factors affecting abiotic degradation. We would refer the reader to the referenced articles for more detailed discussions of the chemistry involved. Biotic degradation Degradation of polymers by biota is particularly important for polymers with slow degradation rates. The action of bacteria, fungi and other biota on degradable polyesters can be marked. Many enzymes have the ability to accelerate the degradation process and in the case of biologically derived PHB, it is the main route of degradation in the environment. The by-products of degradation can depend strongly on the polymer concerned. For example, the enzymatic degradation of PLA generally results in the formation of lactic acid, however the degradation of PCL results in the formation of a range of components such as succinic acid, butyric acid, valeric acid and hexanoic acid. There are fewer reports of enzymatic degradation of polycarbonates although it has been shown that lipases are able to degrade polytrimethylene carbonate into its cyclic monomer as well as linear and cyclic oligomers and 1,3-propanediol (Matsumura et al., 2001).

2.4

Conclusions

The area of biodegradable polymers is very broad and includes a massive range of materials, both as simple polymers and as blends or composites. In this chapter we have tried to capture the bulk of this work and highlight key points in understanding polymer degradation. Other polymer types which may be of interest to the reader but have fallen outside of the scope of this chapter are reviewed in the literature including; polyperoxides (Sato and Matsumoto, 2009), polyethylenimines (Jiang et al., 2008) and polycyclic acetals (Falco et al., 2008).



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References

Campoccia D, Doherty P, Radice M, Brun P, Abatangelo G and Williams DF (1998) ‘Semisynthetic resorbable materials from hyaluronan esterification’, Biomaterials, 19, 2101–27. Chena GQ and Wua Q (2005) ‘The application of polyhydroxyalkanoates as tissue engineering materials’, Biomaterials, 26, 6565–78. Deasy PB, Finan MP and Meegan MJ (1989) ‘Preparation and characterization of lactic/ glycolic acid polymers and copolymers’, J. Microencapsulation, 6, 369–78. Edgar KJ, Buchanan CM, Debenham JS, Rundquist PA, Seiler BD, Shelton MC and Tindall D (2001) ‘Advances in cellulose ester performance and application’, Progress in polymer science, 26, 1605–88. Falco EE, Patel M and Fisher JP (2008) ‘Recent developments in cyclic acetal biomaterials for tissue engineering’, Applications Pharmaceutical Research, 25, 2348–56. Friess W (1998) ‘Collagen – biomaterial for drug delivery’, European Journal of Pharmaceutics and Biopharmaceutics, 45, 113–36. Gaffney PJ (2001) ‘Fibrin degradation products – a review of structures found in vitro and in vivo’, Annals of the New York Academy of Sciences, 936, 594–610. Gilding DK and Reed AM (1979) ‘Biodegradable polymers for use in surgery polyglycolic/ poly(actic acid) homo- and copolymers: 1’, Polymer, 20, 1459–64. Gopferich A and Tessmar J (2002) ‘Polyanhydride degradation and erosion’, Advanced Drug Delivery Reviews, 54, 911–31. Hakkarainen M (2002) ‘Aliphatic polyesters: abiotic and biotic degradation and degradation products’, Advances in Polymer Science, 157, 113–38. Holmes PA, Wright LF and Collins SH (1985) Beta-hydroxybutyrate polymers European Patent Specification 0 052 459. Jiang H-L, Arote R, Jere D, Kim Y-K, Cho M-H and Cho C-S (2008) ‘Degradable polyethylenimines as gene carriers’, Materials Science and Technology, 24, 1118– 26. Kamath KR and Park K (1993) ‘Biodegradable hydrogels in drug delivery’, Advanced Drug Delivery Reviews, 11, 59–84. Kint D and Munoz-Guerra S (1999) ‘A review on the potential biodegradability of poly(ethylene terephthalate)’, Polymer International, 48, 346–52. Kiss E and Vargha-Butler EI (1999) ‘Novel method to characterize the hydrolytic decomposition of biopolymer surfaces’, Colloids and Surfaces B: Biointerfaces, 15, 181–93. Kiss E, Kutnyanszky E and Bertoti I (2010) ‘Modification of poly(lactic/glycolic acid) surface by chemical attachment of poly(ethylene glycol)’, Langmuir, 26, 1440–4. Kumara N, Langerb RS and Domba AJ (2002) ‘Polyanhydrides: an overview’, Advanced Drug Delivery Reviews, 54, 889–910. Labet M and Thielemans W (2009) ‘Synthesis of polycaprolactone: a review’, Chemical Society Reviews, 38, 3484–504. Lucas F, Shaw JTB and Smith SG (1962) ‘Some amino acid sequences in the amorphous fraction of the fibroin of Bombyx mori.’, Biochemical Journal, 83, 164–71. Lunt J (1998) ‘Large-scale production, properties and commercial applications of polylactic acid polymers’, Polymer Degradation and Stability, 59, 145–52. Mano JF, Silva GA, Azevedo HS, Malafaya PB, Sousa RA, Silva SS, Boesel LF, Oliveira JM, Santos TC, Marques AP, Neves NM and Reis RL (2007) ‘Natural origin biodegradable systems in tissue engineering and regenerative medicine: present status and some moving trends’, Journal of the Royal Society Interface, 4, 999–1030. © Woodhead Publishing Limited, 2011


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Matsumura S, Harai S and Toshima K (2001) ‘Lipase-catalyzed transformation of poly(trimethylene carbonate) into cyclic monomer, trimethylene carbonate: a new strategy for sustainable polymer recycling using an enzyme’, Macromolecular Rapid Communications, 22, 215–18. Merrill WW and Pocius AV (1991) ‘Direct measurement of molecular level adhesion forces between biaxially oriented solid polymer films’, Langmuir, 7, 1975–80. Nelson JF, Hilton GS and Cutright DE (1977) ‘Evaluation and comparisons of biodegradable substances as osteogenic agents’, Oral Surgery, Oral Medicine, Oral Pathology, 43, 836–43. Niu X, Feng Q, Wang M, Guo X and Zheng Q (2009) ‘In vitro degradation and release behavior of porous poly(lactic acid) scaffolds containing chitosan microspheres as a carrier for BMP-2-derived synthetic peptide’, Polymer Degradation and Stability, 94, 176–82. Otake Y, Kobayashi T, Asabe H, Murakan N and On K (1995) ‘Biodegradation of low-density polyethylene, polystyrene, polyvinyl chloride, and urea formaldehyde resin buried under soil for over 32 years’, Journal of Applied Polymer Science, 56, 1789–96. Postlethwait RW, Dillon ML and Reeves JW (1962) ‘Experimental study of silk suture’, AMA Archives of Surgery, 84, 698–702. Rokicki G (2000) ‘Aliphatic cyclic carbonates and spiroorthocarbonates as monomers’ Progress in Polymer Science, 25, 259–342. Sajjadian A, Naghshineh N and Rubinstein R (2010) ‘Current status of grafts and implants in rhinoplasty: Part II. Homologous grafts and allogenic implants’, Plastic and Reconstructive Surgery, 125, 99E–109E. Sato E and Matsumoto A (2009) ‘Facile synthesis of functional polyperoxides by radical alternating copolymerization of 1,3-dienes with oxygen’, The Chemical Record, 9, 247–57. Schmidt JJ, Rowley J and Kong HJ (2008) ‘Hydrogels used for cell-based drug delivery’, Journal of Biomedical Materials Research Part A, 87A, 1113–22. Sinclair RG (1996) ‘The case for polylactic acid as a commodity packaging plastic’, Journal of Macromolecular Science A, 33, 585–97. Suriyamongkol P, Weselake R, Narine S, Moloney M and Shah S (2007) ‘Biotechnological approaches for the production of polyhydroxyalkanoates in microorganisms and plants – A review’, Biotechnology Advances, 25, 148–75. Thomas CM, (2010) ‘Stereocontrolled ring-opening polymerization of cyclic esters: synthesis of new polyester microstructures’, Chemical Society Reviews, 39, 165–73. Vidil C, Braud C, Garreau H and Vert M (1995) ‘Monitoring of the poly( d,l-lactic acid) degradation by-products by capillary zone electrophoresis’, Journal of Chromatography A, 711, 323–9 Wang L, Zhang Z, Chen H, Zhang S and Xiong C (2010) ‘Preparation and characterization of biodegradable thermoplastic Elastomers (PLCA/PLGA blends)’, Journal Polymer Research, 17, 77–82. Zhu KJ, Hendren RW, Jensen K and Pitt CG (1991) ‘Synthesis, properties, and biodegradation of poly(1,3-trimethylene carbonate)’, Macromolecules, 24, 1736–40.



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3

The electrospinning process, conditions and control

B. R o bb and B. L e n n o x, The University of Manchester, UK

Abstract: Electrospinning is a process that uses an electric potential to overcome the surface tension of a solution to produce an ultra-fine jet, which elongates, thins and solidifies as it travels through the electric field to a collector. Despite being a relatively simple procedure to undertake in a laboratory as it requires minimal equipment, the physics behind the process is complex. To gain an understanding of all the variables and interactions involved in electrospinning, consideration must be given to polymer chemistry, electric field interactions, fluid mechanics, environmental conditions and kinetics. This chapter will discuss the various parameters relevant to electrospinning and the effect they have on the morphology of the fibres that are produced. Key words: environment, fibre morphology, process conditions, solution.

3.1

Introduction

Process control in the electrospinning process is typically limited to identifying the operating conditions that produce fibres with acceptable properties. However, within a laboratory setting, even with these conditions identified, it is reported that there still remains significant variation in the quality of the produced materials. These variations are a result of an incomplete understanding or consideration of all the process variables. There are many factors influencing the morphology of the fibres or fibrous constructs produced and these can be divided into solution parameters, process parameters and ambient parameters which are listed in Table 3.1. Table 3.1 Variables of the electrospinning process divided into classifications Solution parameters

Process parameters

Ambient parameters

Material selection Solvent selection Concentration Viscosity Dielectric constant Conductivity Surface tension Elasticity

Electromagnetic fields (strength and orientation) Spinning distance Solution flow rate Spinneret morphology Collector morphology

Humidity Temperature Atmosphere Air movement



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These factors form a complex web of interaction and changes in any one of them can affect the impact of the others. Each variable will be discussed in detail later in this chapter. However, the following general comments may be made: • •

•

The electric potential must be large enough to overcome the surface tension of the solution. The longer the polymer jet takes to solidify and the further it travels before solidification, the thinner the fibre will be. This is because the droplet has to stretch over the distance the fibre is travelling (elongation of a fixed volume) and the material is only able to flow while it possesses some fluidity. To produce discernible fibres rather than an interlocking web, the jet must have solidified by the time it reaches the collector. This means that a combination of the time-in-flight, which is related to the spinneret–collector separation, and the evaporation rate of the solvent must, in combination, allow the polymer to solidify before hitting the collector.

Most research studies relating to controlling the electrospinning process have focused on maintaining consistent solution and process parameters, although there are several recent studies that have focused on ensuring that environmental conditions remain constant and can be optimised (de Vrieze et al., 2009).

3.2

Solution parameters

The solution parameters are very well documented in the literature and many commonly used polymer–solvent combinations have been investigated (Li et al., 2006). The choice of solution is often dictated by the necessity to form fibres from a given material and, as such, the solution must be made so that the following parameters are conducive to electrospinning: • • • •

viscosity conductivity surface tension elasticity.

The methods that can be employed to alter the solution properties are numerous and include altering the concentration, the choice of solvent, the molecular weight of the polymer and using additives.

3.2.1 Viscosity The solution viscosity is dependent on the concentration (percentage of polymer in solution), the molecular weight of the polymer, ambient temperature and



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presence of impurities in the solution (Sperling, 2006). The viscosity can be altered by adjusting any of these parameters, or if any of these factors must be kept constant, the viscosity can be adjusted by manipulating the operational conditions, such as temperature. The relationship between intrinsic viscosity, [h] and molecular weight (M) is given by the well-known Mark–Houwink–Sakurada equation as follows:

[h] = KMa

[3.1]

where K and a are the Mark–Houwink parameters for a particular polymer– solvent pair at a given temperature (Sperling, 2006). This equation can be used as a guide when choosing the solvent or molecular weight of your polymer. A reduced viscosity has been reported to be a major factor in the formation of beads along the fibre length, along with several other parameters (Fong et al., 1999). As mentioned in the introduction, the ability of the solution to flow is related to how fast and how far individual polymer chains can move in relation to each other and can influence the fibre diameter. It is believed that the creation of beads at low viscosity may be related to inconsistencies in the distribution of polymer when held within the solvent (Lee et al., 2003a). This work also found that the bead diameter and spacing were related to the fibre diameter: the thinner the fibre, the closer the beads were to each other, and the smaller the diameter of the beads. This dependence on viscosity is further supported by Fong et al. (1999), who found that higher polymer concentrations resulted in a higher viscosity of solution and led to the formation of fewer beads. To produce consistent fibres, a constant solution viscosity should be maintained. If the same polymer and solvent combination are used to make a solution of the same concentration, the viscosity will only be comparable between batches at a specified temperature. In situations where the required temperature needs to be either higher or lower than what is required to ensure the desired viscosity is achieved, or the material choice is restricted by availability, regulation (e.g. FDA) or cost, the viscosity can be manipulated by using additives such as salts. Du and Zhang (2008) demonstrated the effect that salt inclusion had on the viscosity of a polymer solution, as well as subsequent changes to fibre morphology.

3.2.2 Conductivity The solution must be ionic in nature so that the electric field can flow between the needle and collector through the solution. If a solution is not sufficiently conductive, fibres may not form. Conductivity may be altered by changing the polymer/solvent, the concentration of the solution, or by the addition of



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additives such as salts. In addition to the polarity and charge of the molecules in solution, conductivity is also affected by the ion mobility, which in turn can be affected by the solution viscosity (Du and Zhang, 2008). The net charge density (measure of the amount of charge carried from the spinneret to the collector, per volume of material deposited) has been found to be related to both bead formation and fibre diameter. Having a high current flow seems to favour the formation of thin, beadless fibres (Fong et al., 1999; Hsu and Shivkumar, 2004). Typically the conductivity of the solution will be difficult to modify without altering other characteristics of the solution. In these situations it may be easier to alter the electric field strength, solution flow rate or other properties of the material to achieve the desired quality. For solutions with a very low conductivity, salt additives can be used. These salts dissociate into ions in solution, increasing the carrier count inside the liquid jet. The use of large quantities of salts can cause an increase in the solution viscosity and can also affect the surface tension of the solution (Choi et al., 2004).

3.2.3 Dielectric constant It has been reported that the dielectric constant of the solution can have an effect on both the speed of deposition and the diameter of the fibres (Wannatong et al., 2004). This is believed to be a result of the interaction of the internal electrical field created in a dielectric material and the external applied electric field (Lee et al., 2003b).

3.2.4 Surface tension As a stand alone factor, surface tension can be very important, as the electric field needs to overcome surface tension in terms of energy to produce the solution jet. Surface tension can be manipulated by changing the material used or by the addition of surfactants to the solution (Du and Zhang, 2008). It is also a very important factor in beading along the fibre as a solution with a high surface tension will favour the formation of beads (Fong et al., 1999). Beads have a higher volume to surface area ratio than a continuous fibre and, by forming beads, the surface energy of the material is lowered as less surface is created. This is counter-productive to the common aim of electrospinning, which is to create very low volume to surface area ratio fibres. If the surface tension is too high or if the viscosity is too low, fibres will not form at all and droplets will be created instead in a process known as electrospraying (Jaworek et al., 2009).



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3.2.5 Solvent choice The choice of solvent used in the production of fibres can have a major impact on their properties. More volatile solvents will evaporate quicker, which will reduce the required spinning distance. Different solvents will result in polymer-containing solutions that have varying viscosities, which can influence the final morphology of fibres. It has also been shown (Fig. 3.1 and 3.2) that for the same process conditions, changing the solvent

20 µm

3.1 PCL fibres spun in a 10% w/v solution with HFIP over 20 cm with a 25 kV potential.

5 µm

3.2 PCL showing beading spun from a 10% w/v solution with acetone over 20 cm using a 25 kV potential.



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from acetone to hexafluoroisopropanol (HFIP) can eradicate the presence of beading in poly-e-caprolactone (PCL), as well as alter the mechanical properties of the materials produced.

3.2.6 Polymer choice Often the choice of polymer is determined by the specific application of the fibres that are to be produced, such as for biodegradable scaffolds for tissue engineering or chemically resistant filters. However, altering the polymer molecular weight can vary the fibre properties significantly. Higher molecular weight polymers will increase the solution viscosity even if the solution concentration is constant. The polymer molecular weight may also affect its level of dissolution in the chosen solvent.

3.2.7 Summary The solution parameters can be manipulated by changing the composition of the solution. Although there are a great deal of data reported on successful solution compositions (Ramakrishna et al., 2005), they are not necessarily definitive in providing a repeatable, uniform fibre as they still depend on both processing and ambient parameters and may not necessarily be transferable from laboratory to laboratory, or from day to day in an uncontrolled environment. The manipulation of solution variables should ideally be restricted to counteracting any unwanted secondary changes in solution properties as a result of changing one or more of the processing and ambient conditions.

3.3

Processing parameters

The processing parameters when producing fibres include: • • • • •

driving electric field strength collector distance/motion in relation to the spinneret any additional magnetic or electric fields hydrostatic pressure on the solution and flow rate spinneret choice.

3.3.1 Electric field strength The electric voltage required to produce a jet from a polymer solution must have a higher potential than the surface energy of the solution at the solution/ atmosphere interface (Doshi and Reneker, 1995). A higher field strength should provide faster acceleration of the emitted jet, when an equal mass of



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solution is being ejected. This increase in jet speed should cause lower fibre thickening or incomplete solidification as the polymer will hit the collector in a reduced time. As mentioned previously, electrospinning will only occur when the electric potential energy supplied is higher than the surface energy of a droplet of solution. When the electric potential energy is approaching the value of the surface energy, the shape of the droplet is distorted to form a Taylor cone. As the critical voltage at which this occurs is dependent on several factors including the surface tension of the solution, the diameter of the droplet and the electric field strength (distance to target), altering the voltage can be compensated for by small changes in other parameters. Maintaining a constant electric field strength is straightforward when using a variable voltage DC supply. Voltages are usually in the range of 5–35 kV.

3.3.2 Collectors The type of collector that is used in the process affects the morphology of the collected fibres. A simple collector plate will form an unwoven mat of fibres in a random orientation, whereas other specifically designed collectors can collect aligned fibres. Specialised collectors include spinning disks, drums and mandrels for which the speed of rotation is an important factor that affects the fibre deposition rate. If the collector is moving faster than the fibres are being deposited, the fibres will stretch or break into small fragments as the fibre sticks to the collector and is accelerated away. Aligned PCL fibres collected using a rotating mandrel are shown in Fig. 3.3 and can be compared to a non-woven mat of fibres in Fig. 3.4. If the rotation of the collector is too slow, the fibres will mat onto the surface similar to those collected on a simple plate. However, this approach is sometimes used to create non-aligned or partially aligned, porous cylinders. These hollow tubes have potential use in many areas including tissue engineering applications such as vascular scaffolds and nerve guide conduits (Bini et al., 2004). The configuration of the collector depends on the type of fibre orientation desired. For aligned fibres, the speed of collector rotation needs to be calibrated to the polymer deposition rate (Thomas et al., 2007) and can be achieved using a variable speed motor, such as an overhead stirrer, or a variable speed drill to drive the collector.

3.3.3 Spinning distance The distance between the spinneret and the collector is a key factor in determining the morphology of fibres that are produced. As mentioned earlier, jet elongation and thinning only happens while the jet is in flight and still a



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10 µm

3.3 PCL fibres spun in a 10% w/v solution with HFIP over 20 cm with a 25 kV potential onto a mandrel rotating at 700 rpm.

10 µm

3.4 PCL fibres spun in a 10% w/v solution with HFIP over 20 cm with a 25 kV potential onto a stationary plate.



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fluid. This elongation occurs owing to charge repulsion between ions in the solution combined with a net pull towards the collector. Bending instabilities in the jet can cause it to whip around, increasing the travel distance. While in flight, the polymer solution solidifies as the solvent is evaporated from the surface, forming polymer fibres. Therefore, increasing the spinning distance will increase the time for thinning to occur and provided the polymer is not yet solid, the average fibre diameter will be reduced. To ensure that elongation occurs over the entire distance, the solvent, temperature, humidity and atmosphere composition should be specified such that the jet solidifies just before hitting the collector. If the jet is not solid when it hits the collector, the fibres may congeal into a web as shown in Fig. 3.5. Furthermore the electric field needs to be sufficient to accelerate the polymer jet over the entire distance, taking into account opposing forces such as air resistance and localised charge concentrations. The spinning distance also has an effect on the field strength. As this distance increases, the field weakens following an inverse square relationship.

3.3.4 Additional electromagnetic fields Secondary fields may be used to control the deposition of the nanofibres. Just as the polymer jet is attracted to a grounded plate, perpendicular fields may be used to ‘steer’ the path of the polymer jet. Bellan and Craighead

10 µm

3.5 PLGA fibres spun in a 20% w/v solution with HFIP over 20 cm with a 25 kV potential showing incomplete solidification.



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(2009) investigated the use of four additional electrodes arranged around the main jet path. Each electrode was switched on individually by a computer to attract the polymer jet towards it. When the electrode was switched off, the jet returned to its central direct path and two straight aligned fibres were deposited between the electrode and the centre of the collector. By repeating this on–off switching, aligned fibres could be produced without the need for a rotating collector. In theory this method could be expanded to create a three dimensional, controlled deposition manufacturing process for irregular shaped structures.

3.3.5 Flow rate The rate at which the polymer solution is ejected from the needle is typically governed by hydrostatic pressure, using a syringe pump. Occasionally electrospinning is performed in a vertical configuration, allowing the effects of gravity to draw out the polymer. Although in this configuration the issue of polymer dripping from the needle tip onto the collector exists, if the flow rate is not ideal, it will consequently spoil the fibres already collected. The force required to deliver the solution at a constant rate will depend on several factors, including the diameter of the spinneret tip and the solution viscosity. If the flow rate is too low the solution may solidify in the spinneret, or not be able to form a Taylor cone owing to insufficient solution at the surface. Even when the flow rate is sufficient to produce nanofibres, complications due to flow rate can occur such as solution dripping or the formation of ‘caterpillars’ from the tip of the spinneret. A ‘caterpillar’ results when the solution solidifies at the atmospheric interface forming a polymer tube, which has the effect, in practice, of decreasing the spinning distance. To produce high quality fibres, the flow rate should be specified depending on the polymer/solvent choice and concentration. It has also been suggested that the flow rate has an effect on the diameter of the fibres produced as well as any beading that may occur along the length of these fibres (Zong et al., 2002). This is due to a larger volume of solution being charged and subsequently ejected from the needle tip for any time period at a higher flow rate. If there is a larger volume of solution in motion over a set distance, the resulting fibres will be thicker. Druesedow (2008) highlights the issues surrounding up-scaling and commercialisation of the electrospinning process resulting from non-uniform flow rates. He suggests that by using an ultrasonic sensor to monitor the hydrostatic pressure in the capillary with an active feedback loop to the syringe pump, the dripping or early solidification can be restricted. He also suggested that the fibre morphology is not affected significantly by improved flow rate control and that the whipping instability of the solution jet has a more direct influence on fibre diameters (Druesedow, 2008).



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3.3.6 Spinneret The diameter of the spinneret tip was extensively reviewed by Macossay et al. (2007), who concluded that there was no correlation between the internal diameter of the spinneret and the average fibre diameter (Macossay et al., 2007). However, the internal diameter does have an effect on the flow rate of the polymer solution, as well as the shape and size of the solution–air interface, which will ultimately influence the critical voltage required for electrospinning to occur (Yarin et al., 2001). The design of the spinneret varies greatly in the literature. In most applications, a simple hypodermic needle is used, with the tip blunted to form a flat opening. Commercially sold electrospinning units come with custom built spinnerets, some of which are designed with an inbuilt heating element so as to be able to spin high temperature solutions or melts. MECC Ltd, for example, make a variety of spinneret tips for their commercial electrospinners, including a heating spinneret and remote (tube fed) and local (syringe in place) spinneret options. Other spinneret configurations include co-axial spinneret systems which have been extensively used in the formation of both core-shell copolymer fibres and hollow fibres. Core-shell and hollow fibres are produced by injecting one polymer solution into the middle of a stream of another solution, using a dual funnel configuration. To produce a hollow fibre, the central polymer is usually dissolved out, leaving the shell behind.

3.3.7 Summary The manipulation of processing parameters is undertaken by making physical alterations to the equipment used. The degree of control is similarly dependent on the equipment and, once again, there is a variety of proven parameters to be found in the literature for many material/solvent combinations. While standardising the spinning distance and flow rate in a particular field strength can ensure similar results for a certain solution, this is only true if the ambient conditions are kept the same.

3.4

Ambient parameters

As previously mentioned, the ambient parameters can change on a day to day and laboratory to laboratory basis. These parameters offer the biggest opportunity for enabling fibres to be produced more consistently. These factors include: • • •

humidity temperature ambient gas composition and movement.



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3.4.1 Humidity The surrounding humidity will affect each solution differently, depending on the solvent used and the hydrophilicity of the polymer solution. Aqueous solutions are obviously most affected as the humidity is a measure of the vapour pressure of the solvent in the atmosphere and it can be expected that the water in solution and in the atmosphere will interact. It has been reported that increased humidity slows the solidification process for aqueous solutions, increasing both beading defects and fibre diameters (Druesedow, 2008). The evaporation rate for a pure liquid from a free surface is proportional to the difference between the saturated vapour pressure (100% humidity in the case of an aqueous solution) and the vapour pressure of the surrounding environment. A higher humidity will slow the evaporation rate of water, thus increasing the drying time, which would produce thinner fibres given enough time for the water to evaporate completely before the jet hits the collector. Non-aqueous solutions can be affected in several ways. If the relative vapour pressure of water is too high, evaporation will be slowed by a saturation effect even if there is no solvent present in the atmosphere. There will be a limit to how much liquid the atmosphere can hold at a certain pressure before either water or solvent start to condense back into liquid form. This factor can have several consequences, thinner fibres can be produced as longer solidification times may result or a congealed mat of fibres may form owing to incomplete solidification. Another impact that humidity will have on non-aqueous solutions is that the solution can absorb water during flight (Jeun et al., 2007). The rate of absorption will be dependent on the material’s affinity for water and the solution’s ability to absorb water. This phenomenon could cause a slowing in evaporation, which may cause an increase in fibre diameter, introduce pores on the fibre surface owing to solvent concentration variations, or produce congealed mats instead of unwoven fibres. According to Jeun et al. (2007), increasing the humidity from 30% to 50% caused PLLA fibres to display porosity when spun from a chloroform solution. Humidity can be controlled using a ventilation system in a confined area. When air is passed through an environmental chamber, its humidity can be regulated by a variety of apparatus including a hot water bath or dry ice to raise and lower humidity, respectively (Simonet, 2007).

3.4.2 Temperature The ambient temperature in which the spinning is performed has a substantial effect on the fibres produced. There are two main effects resulting from an increase in temperature: the solvent evaporation rate increases and the viscosity decreases. © Woodhead Publishing Limited, 2011


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The solvent evaporation rate increases relative to an increase in temperature. In a warmer environment, individual molecules of solvent will have more energy and thus more molecules will have the energy required to jump from liquid to gaseous form, making the solution dry faster. The viscosity decreases with an increase in temperature. This is attributed to several properties of the polymer solution. Softening of polymers occurs at higher temperatures, with two major steps occurring at two particular points: the glass transition (where the polymer changes from brittle to elastic deformation behaviour) and the melting point (where the polymer chains disentangle to allow viscous flow). The higher the temperature, the more flexibility is displayed by each individual polymer molecule. The polymer chains have a greater degree of freedom to move owing to the increase in energy breaking or weakening cross-linkages, Van der Waals forces and hydrogen bonding between the chains. This effect can be modelled using Barr’s approximation of Vogel’s original equation for the viscosity of hydrocarbon-based liquids over a narrow temperature range. This is shown in equation 3.2. b

m = Ke T +q

[3.2]

where m is the absolute viscosity, T is temperature and b, K and q are empirical constants. If the polymer chains have more freedom of movement (i.e. lower viscosity), the solution will be able to flow faster and will require less energy to induce flow, allowing the polymer to elongate more and become thinner during its flight. overall, the temperature has two opposing effects on the electrospinning process relating to a change in solvent evaporation rate and in solution viscosity. At high temperatures the solution dries quickly, leaving little time for elongation and thinning. However, at increased temperature, the decrease in viscosity will cause the solution to flow faster before drying, allowing the fibre to elongate faster. The ambient temperature has been regulated during fibre production in a number of ways, including convection fan heating and infrared lighting (Givens et al., 2007), usually within a closed chamber. As with humidity regulation, the temperature is initially stabilised at a required level. The regulator is then switched off, if using a fan system, so that the fibre deposition is not affected by air movement (De Vrieze et al., 2009). If the fan system is kept on during electrospinning, consideration should be given to the direction of air movement in relation to the direction the polymer jet will travel. Infrared heat sources will cause less air movement, but will not heat a fixed volume of air evenly, so the temperature must be controlled by a probe as close to the polymer jet as possible.



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3.4.3 Ambient gas composition and movement For aqueous solutions, the relative humidity is an important factor, as the solvent evaporation rate is proportional to the amount of water in the atmosphere as mentioned earlier. The oxidisation of electrospun materials has been investigated in the case of oxidised cellulose (Won et al., 2004). Although this oxidation is currently achieved after electrospinning the fibre mat, it is possible to oxidise the cellulose during the production of fibres by controlling the environment. It has also proved favourable to use a vacuum in order to control the electric field, as shown by Rangkupan and Reneker (2003). Molten polypropylene was electrospun in a vacuum so that higher than normal electric fields could be achieved. In these cases, electrospinning under vacuum or within a non-air atmosphere is likely to have a significant impact on the rate of solidification for the polymer. Evaporation of solvents into a vacuum will be very fast as the boiling point of liquids is proportional to atmospheric pressure. Electrospinning in a closed environment can lead to fibre inconsistency over long periods as the build up of solvent in the atmosphere affects the evaporation rate of the solvent. Constantly removing the solvent-saturated gas and replacing it with ‘dry’ gas during the electrospinning process will maintain a constant concentration of solvent in the atmosphere. Applying a sensor to monitor the concentration of solvent in the atmosphere through venting could be a useful way to maintain a constant solvent evaporation rate. The airflow around an electrospinning set-up will have a profound effect on the fibres produced. All fibres produced in a non-vacuum environment will be subject to drag as they pass through the atmosphere like any other object. By controlling the direction and speed of any air flow, the speed of a polymer jet’s flight can be controlled. For example, if the net direction of air flow opposes the direction of polymer jet flow, increased elongation may occur as the jet will be displaced in flight and the time taken to contact with the collector will be increased.

3.5

Conclusions

Electrospinning is a highly complex process, whose end result depends on a great number of factors. This makes finding the best set of parameters to use for a particular application difficult. However, it does also allow the quality and morphology of fibres produced to be regulated by selectively controlling these parameters. Very often it is favourable to spin at room temperature, as this is cheaper than to operate at particularly high or low temperatures. However, if the fibre quality would benefit from a higher operating temperature, it may be favourable to find an alternative method to effect the same change



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to the process, such as using a more volatile solvent or lowering the solution concentration. Owing to the abundance of processing parameters and the level of interaction between them, there are usually several factors that can be altered to achieve the same result and compensating for any one required variable, such as the temperature, is easily achievable. To understand fully all the factors relating to the electrospinning process, and find an effective and quantifiable understanding of the interaction between these various factors, a flexible electrospinning system (with the ability to change all the process related parameters discussed in this chapter) needs to be placed in an environmental isolation chamber, which allows the ambient conditions to be regulated. A full factorial study to isolate and examine each parameter individually in relation to all of the others will then be possible for a particular material choice. This needs to include analysis of the morphological and chemical effects of changing each of these parameters to establish relationships between each variable. Once this has been established for several similar materials, a mathematical model of the process can be drawn up to approximate the required conditions needed to produce fibres with a desired morphology for any material and application.

3.6

References

Bellan L and Craighead H (2009). ‘Nanomanufacturing using electrospinning’, J Manufact Sci Eng, 131(3), 034001. Bini T B, Gao S, Tan T C, Wang S, Lim A, Hai L B and Ramakrishna S (2004). ‘Electrospun poly(l-lactide-co-glyoclide) biodegradable polymer nanofibre tubes for peripheral nerve regeneration’, Nanotechnology, 15, 1459. Choi J S, Lee S W, Jeong L, Bae S H, Min B C, Youk J H and Park W H (2004). ‘Effect of organosoluble salts on the nanofibrous structure of electrospun poly(3hydroxybutyrate-co-3-hydroxyvalerate)’, Int J Biol Macromol, 34(4), 249–56. De Vrieze S, Van Camp T, Nelvig A, Hagström B, Westbroek P and De Clerck K (2009). ‘The effect of temperature and humidity on electrospinning’, J Mater Sci, 44, 1357–62. Doshi J and Reneker D H (1995). ‘Electrospinning process and applications of electrospun fibers,’ J Electrostat, 35(2–3), 151–60. Druesedow C J (2008). Pressure control system for the electrospinning process: noninvasive fluid level detection using infrared and ultrasonic sensors. MSc Thesis, University of Akron. http://etd.ohiolink.edu/view.cgi?acc_num=akron1217275502. Du J and Zhang X (2008). ‘Role of polymer-salt-solvent interactions in the electrospinning of polyacrylonitrile/iron acetylacetonate’, J Appl Polym Sci, 109(5), 2935–41. Fong H, Chun I and Reneker D H (1999). ‘Beaded nanofibers formed during electrospinning’, Polymer, 40(16), 4585–92. Givens S R, Gardner K H, Rabolt J F and Chase D B (2007). ‘High-temperature electrospinning of polyethylene microfibers from solution’, Macromolecules, 40(3), 608–10. Hsu C and Shivkumar S (2004). ‘Nano-sized beads and porous fiber constructs of poly(ecaprolactone) produced by electrospinning’, J Mater Sci, 39(9), 3003–13.



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Jaworek A, Krupa A, Lackowski M, Sobczyk A T, Czech T, Ramakrishna S, Sundarrajan S and Pliszka D (2009). ‘Electrospinning and electrospraying techniques for nanocomposite non-woven fabric production’, Fibers and Textiles in Eastern Europe, 17(4), 77–81. Jeun J P, Kim Y H, Lim Y M, Choi J H, Jung C H, Kang P H and Nho Y C (2007). ‘Electrospinning of poly(l-lactide-co-d, l-lactide)’, J Ind Eng Chem, 13, 592. Lee K H, Kim H Y, Bang H J, Jung Y H and Lee S G (2003a). ‘The change of bead morphology formed on electrospun polystyrene fibers’, Polymer, 44, 4029–34. Lee K H, Kim H Y, Khil M S, Ra Y M and Lee D R (2003b), ‘Characterization of nano-structured poly(e-caprolactone) nonwoven mats via electrospinning’, Polymer, 44(4), 1287–94. Li W-J, Cooper J A, Mauck R L and Tuan R S (2006). ‘Fabrication and characterization of six electrospun poly(alpha-hydroxy ester)-based fibrous scaffolds for tissue engineering applications’, Acta Biomateria, 2(4), 377–85. Macossay J, Marruffo A, Rincon R, Eubanks T and Kuang A (2007). ‘Effect of needle diameter on nanofiber diameter and thermal properties of electrospun poly(methyl methacrylate)’. Polym Adv Technol, 18(3), 180–3. Ramakrishna S, Fujihara K, Teo W E, Lim T C and Ma Z (2005). ‘Electrospinning process’, An Introduction to Electrospinning and Nanofibers, Chapter 3, World Scientific Publishing, Singapore, 117. Rangkupan R and Reneker D H (2003). ‘Electrospinning process of molten polypropylene in a vacuum’, J Metals, Mater Minerals, 12(2), 81–7. Simonet M, Schneider O D, Neuenschwander P and Stark W J (2007). ‘Ultraporous 3D polymer meshes by low-temperature electrospinning: Use of ice crystals as a removable void template’, Polym Engi Sci, 47(12), 2020–6. Sperling L H (2006). Introduction to physical polymer science, 4th edition, Hoboken, NJ, USA. Thomas V, Dean D R, Jose M V, Mathew B, Chowdhury S and Vohra Y K (2007). ‘Aligned bioactive multi-component nanofibrous nanocomposite scaffolds for bone tissue engineering’, Biomacromolecules, 8, 631. Wannatong L, Sirivat A and Supaphol P (2004). ‘Effects of solvents on electrospun polymeric fibers: preliminary study on polystyrene’, Polym Int, 53, 1851–9. Won K S, Ji H Y and Won H P (2004). ‘Preparation of ultrafine oxidized cellulose mats via electrospinning’, Biomacromolecules, 5(1), 197–201. Yarin A L, Koombhongse S and Reneker D H (2001). ‘Taylor cone and jetting from liquid droplets in electrospinning of nanofibers’, J Appl Phys, 89(9), 4836–46. Zong X, Kim K, Fang D, Ran S, Hsiao BS and Chu B (2002). ‘Structure and process relationship of electrospun bioabsorbable nnofiber membranes’, Polymer, 43(16), 4403–12.



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4

Regulatory issues relating to electrospinning

A. W i l s o n, CellData Services, UK

Abstract: This chapter introduces the concept of regulation of products sold or promoted for medical purposes. The regulatory frameworks for medical devices and medicinal products in Europe are described. Factors that cause a medical product to be classified as a device or a medicine are examined in the context of electrospun scaffolds and the influence of claims or the manufacturer’s intended use upon classification is discussed. Situations in which a product used for regenerative medicine may be excluded from the current regulatory regime are introduced and the progress at European Union level towards the closure of such regulatory gaps is mentioned briefly. Key words: advanced therapy, classification, medical device, medicinal product, regulation.

4.1

Introduction

The manufacture and marketing of medical products are subject to extensive control via a range of methods. Legal instruments such as European Regulations and Directives, and national laws of the Member States, establish a framework that controls all aspects of the medical products business. These instruments are supplemented by several means, including guidelines and international standards. The classification of a product as a medical device or a medicinal product determines almost every aspect of development, approval for sale and subsequent marketing, and therefore an early and correct determination of how it will be regulated is of paramount importance.

4.1.1 Regulation of medical products This chapter introduces the concept of regulation of products sold or promoted for medical purposes. Such products are generally regulated as medicines or medical devices; their fundamental purpose or intended use relates to prevention, diagnosis or treatment of a disease, or management of a physiological condition. Any product sold or promoted for such purposes may potentially be subject to regulation as a medicine or a device and the claims the manufacturer makes for it can have a critical effect on whether the product falls under the device or medicinal products framework or is not regulated under either. These frameworks are very different and manufacturers need a clear understanding of how their products are regulated in order to develop both the product and its supporting data to meet the relevant 67 © Woodhead Publishing Limited, 2011


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regulations. Borderlines with other regulatory frameworks also exist, in which ambiguities in the directives lead to disagreements between manufacturers and regulatory authorities and even between different regulatory authorities about how a product should be classified. This situation arises from the sometimes unavoidable grey areas that can exist between the legal definitions of products regulated under different directives. As an example, until recently the status of whitening toothpaste was the subject of much discussion in Europe: should this common consumer product be regulated as a medical device or be subject to the requirements of the European cosmetics directives? A short description of the regulatory frameworks will be presented, followed by some discussion of the factors that influence classification of products as medicinal products or devices. The key features of scaffolds used in regenerative medicine will be analysed from the regulatory viewpoint and the use of standards and regulatory guidelines in the development process will be considered as a means of guiding the process and meeting regulators’ expectations. From a strategic perspective the need for clear regulatory input at appropriate stages in product development is essential: there is little possibility of meeting all of the requirements for commercialising a medicine when the product’s development programme is based on the incorrect assumption that it is a medical device. Beyond the simple need to meet the requirements, many of the regulatory ‘hurdles’ that the manufacturer has to clear are in fact tools that can greatly assist in streamlining the development process. An understanding of how obligations such as design control, risk analysis and preparation of detailed instructions for use can guide and clarify the development and will greatly improve the potential for success in a commercially viable timeframe. In considering the regulation of electrospun scaffolds in regenerative medicine, these principles will be explored to illustrate how scaffolds are regulated and how strategic regulatory input can help to optimise the development process.

4.1.2 History of regulation for tissue regeneration products Products used for tissue regeneration have been subject to a wide range of different regulations. Although electrospun scaffolds themselves are fairly straightforward from the regulatory perspective (Section 4.2.2), the history of other products and technologies in this field is informative. Many tissue regeneration products are themselves tissue-based, whether derived from materials of human or animal origin. Europe has only recently introduced a regulation to define ‘tissue engineering’, with the publication in December 2007 of the Advanced Therapy Medicinal Products Regulation (EC) 1394/2007. 1 This regulation ended the uncertainty and established a formal definition



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for bringing products that meet this definition firmly within the scope of medicinal product regulation. Prior to this regulation, tissue regeneration products containing human cells and tissues were regulated at the national level in Europe since there was no appropriate harmonised EU legislation covering them: in the absence of European regulation, the responsibility for regulation defaults to the individual member states. This regulatory gap existed because the Medical Devices Directive2 contained, and still does contain, an exclusion preventing the use of human-derived materials in medical devices. (A limited use of human-derived materials is permitted by Directive 2000/70/EC which amends the medical device directive to include the use of stable derivatives of human blood and plasma, such as heparin and albumin, in medical devices.) Put another way, if a device contains or is made from human-derived materials it cannot be CE-marked as a medical device. The consequence of this exclusion is that a number of medical products, which are not medicines and would otherwise be regulated as medical devices, have no legal route to the European single market. The difficulty of marketing such a product via approvals in 27 different countries has meant that some human tissue products such as demineralised bone and decellularised skin, whilst readily available in other developed markets, are rarely sold in European countries. The majority of electrospun scaffolds in themselves will generally be regulated as medical devices if sold for the purpose of tissue regeneration, since they do not contain human-derived materials. The absence of an appropriate regulatory system for devices containing human tissue does not have an impact on these scaffolds unless they are intended to be sold in combination with human cells or tissue-derived materials. As the field develops, however, the use of human tissue-derived scaffolds will necessitate the development of EU-level legislation covering non-viable human tissues. It is anticipated that a future amendment of the Medical Devices Directive will include a provision for including such products as medical devices. Section 4.2.2 examines the factors that cause electrospun scaffolds to be classified as medical devices and discusses situations in which the scaffold may form part of a medicinal product.

4.2

Regulation of materials in regenerative medicine

4.2.1 Overview of European regulatory frameworks for medical devices and medicinal products Regulation in the context of medical products Products sold or presented for purposes such as treatment or diagnosis of disease, management of disease or other medical uses are subject to a © Woodhead Publishing Limited, 2011


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complex framework of legislation that is primarily aimed at protection of the public. The overall intention of the regulations is to ensure that products are not harmful under normal conditions of use, or that their benefits exceed the potential risks, and that they have been shown to provide the benefits claimed by their manufacturer. Healthcare products are, broadly, regulated in one of two ways in the European Union: they are classified as either medicinal products or medical devices. The two systems are completely separate in their legal basis and in their operation and the vast majority of products are readily classified into the appropriate framework. The two systems differ markedly in terms of the procedures to be followed in order to place a product onto the European market and in terms of the costs and timescales associated with their development. However the fundamental health protection principles underlying their regulation and control can be summarised in terms of a few simple questions: ∑

Are the components and the final product acceptably safe for the intended use? ∑ Is the manufacturing process capable of delivering a consistent product? ∑ Does the product perform in the manner intended by the manufacturer? and most critically: ∑

Are the risks associated with the use of the product outweighed by the clinical benefit of using it?

In one sense, the process of product development is the process by which the manufacturer generates the data necessary to answer these questions to the extent required by the regulatory authorities. The process of regulatory approval is the mechanism by which the manufacturer of the healthcare product demonstrates that the legal and scientific requirements of the regulatory system have been met. It is important to appreciate that regulatory requirements go far beyond the testing needed to gain approval to market the product: both frameworks establish postmarketing obligations for the manufacturer to maintain an appropriate product on the market. These may encompass formal post-approval safety or efficacy studies, establishment of surveillance systems to seek and report adverse events associated with the clinical use of the product, and the need to evaluate and seek approval for changes to the product that the manufacturer may wish, or need, to implement after the original version has been placed on the market. The costs and challenges of the post-marketing (often called post-approval) phase are as important to the commercial business case as the pre-approval ones and their impact needs to be factored into the commercial model as rigorously as the initial development costs. Thus, the successful



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development of any healthcare product necessitates a thorough understanding of the regulatory environment in which the product is to be developed and marketed. Regulatory framework for medicinal products The term ‘medicinal product’ is defined in the primary legislation3 governing such products and is synonymous in Europe with the terms ‘pharmaceutical’ and ‘drug product’. In the EU the term also encompasses biological medicinal products such as vaccines, therapeutic proteins and the gene, cell and tissuebased products used in regenerative medicine. The USA recognises ‘biologics’ as a separate class of products subject to different regulation from chemical drugs, but the term has no separate meaning in the EU and biological medicinal products are subject to the same general framework and requirements as other types of medicines. Medicinal products are subject to a pre-approval system in which the manufacturer or company developing the product is required to obtain a licence, known as a marketing authorisation (MA), before the product can be sold in the EU. The licence may be a national one that is valid in the Member State that issued it, or an EU marketing authorisation that is issued by the European Commission following a centralised assessment by the European Medicines Agency (EMA). This latter type of licence is often called a ‘centralised’ MA after the centralised approval procedure that generates such licences and it is valid in all Member States of the EU and the European Economic Area. The system sets out criteria that establish which type of licence the manufacturer can apply for: new medicinal products, those produced by biotechnological processes and products for certain specified indications must be authorised by the centralised procedure. Some products that do not fall within the mandatory scope of the centralised procedure may still be authorised in this way but the manufacturer is required to seek confirmation of the product’s eligibility for the procedure from the EMA before the marketing authorisation application (MAA) can be submitted. Products that are outside the scope of the centralised procedure are authorised nationally by each Member State in which the manufacturer wishes to sell the product: a licence must be obtained in each Member State separately. The regulations provide for two assessment routes in this situation: the decentralised Procedure (DP) and the mutual recognition procedure (MRP). Both require the submission of a complete dossier of information to each Member State in which a licence is sought. Finally, a national procedure exists for products to be sold in one Member State only: this is increasingly less likely as the escalating costs of medicinal product development are less and less able to be recouped by anything short of global marketing. It is important to recognise that the extent of scientific data and the burden of proof expected is the same for each procedure; none of them represent a soft



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option in terms of product development. Within each of these procedures there arises scope for submission of a shortened application in certain legally defined situations. These ‘abridged’ applications may be submitted for products that are essentially similar to a product already authorised in the EU. They contain less pre-clinical and clinical data than a full application but they are acceptable only in very particular circumstances: either the owner of the pre-approved product must give the regulatory authority its agreement to refer to the original data when reviewing the second applicant’s submission, or the pre-approved product must have been authorised for a minimum of 10 years before the second applicant can make reference to the data. This second option is the basis of generic drug approvals. It is also possible to support an application solely by reference to published literature for pre-clinical and clinical data: this route (the ‘bibliographic’ application) is only allowable for medicines with a well-established authorised medical use and is quite impractical in most circumstances because of the difficulty of demonstrating that the data available from a publication, which may be several years old, meets current standards for study design, analysis and good laboratory or clinical practice. The scientific and clinical data required to support a licence application for a new medicinal product is largely harmonised across much of the world. The International Conference on Harmonisation (ICH) has established both the general data requirements and the application dossier format to be used when seeking an authorisation for a medicine in the EU, the US and Japan. Many other major countries, including Canada and Australia, have also adopted the ICH rules, making the compilation of a global dossier at least a possibility, although the legal requirements for marketing and the procedures under which the data are assessed remain very disparate in the different territories. The use of an unapproved medicinal product, for example in a clinical trial, or its use in an unauthorised indication or patient population, requires approval from the competent authority for medicines (regulator) in the Member State(s) in which the trial is to take place. The manufacture of clinical trial products and the conduct of clinical trials are regulated by specific European directives that are separate from the procedures required for authorisation to market the product. A full discussion of the legal basis, procedures and data requirements for authorisation of medicinal products is beyond the scope of this chapter; further information sources are identified in Section 4.4. Regulatory framework for medical devices Medical devices sold in the EU are subject to one of three medical device directives: the general medical devices directive (MDD),2 the active implantable medical device directive (AIMDD)4 and the in vitro diagnostic



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device directive (IVDD),5 depending on the type of device under consideration. For the purposes of this discussion only the general MDD will be considered. The scope of the other two directives is not likely to be relevant to tissue regeneration products. The medical device directives form a regulatory regime that is fundamentally quite different to that for medicinal products. The authorisation procedure for medical devices may take several different routes ranging from ‘self-certification’, in which the manufacturer makes no submission for external assessment but instead signs a declaration that their product is in compliance with the requirements of the relevant medical devices directive, to prior authorisation of the product by one or more external assessment agencies before they can make the declaration that their product complies with the directives. The central principle of the medical device directives is the establishment of ‘essential requirements’ (ERs) to which the device must conform, but the solutions the manufacturer may use to demonstrate conformity with the ERs are not specified in the legal documents. The use of ERs is a cornerstone of the ‘New Approach’ adopted by the European Commission in order to establish requirements for the safety and suitability for purpose of a wide range of goods without being unduly restrictive. This approach is designed for fast-moving and innovative industries which develop and introduce new technologies rapidly and that would be disadvantaged by the imposition of fixed requirements upon a variable and rapidly evolving technical field. The manufacturer demonstrates their assertion that they have met all of the ERs relevant for their product by affixing the ‘CE mark’ to the product or its labelling (Fig. 4.1). This symbol is interpreted to mean that all requirements applicable to that product have been met and that the product and its manufacturer are in compliance with the relevant directives. It is the responsibility of the manufacturer to determine which directive(s) apply to their product. The

4.1 The CE mark.



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CE mark can be seen on the body or label of all medical devices (e.g. on a box of first aid plasters) and on other goods covered by the New Approach directives, including toys and personal protective equipment such as steel toe-capped boots and horse-riding hats. The rules by which medical devices are classified by the Medical Device Directive have the effect of stratifying them in terms of the amount of risk their use poses, although the terms ‘high risk’ and ‘low risk’ are not used in the classification rules. Devices are grouped into classes by aspects such as route of exposure (intact skin, breached or injured skin, natural body orifices or introduction into the body via a surgical procedure) and the way in which they work (e.g. supply of ionising radiation, delivery of medicinal products, intended to be absorbed by the body). Additional ‘special’ rules apply and these have the effect of placing the device in a specified class regardless of its other attributes. For example, all devices containing or derived from non-viable animal tissue are in Class III (highest risk) regardless of how they function or how they contact the body. The directive sets out the ways in which the manufacturer can meet the ERs and these ‘conformity assessment routes’ vary according to the classification of the device. For all but the simplest non-sterile devices, the CE-marking process involves a notified body, an organisation appointed by the competent authority for medical devices, which carries out assessments in accordance with the chosen route to compliance before manufacturers can place their products on the market. The extent of notified body involvement varies with device classification, from confirmation that the manufacturer operates an appropriate quality system and keeps the necessary technical file on the product, through to a pre-marketing assessment of a product’s design, manufacture and testing before the manufacturer can CE mark the device. A key underlying principle for medical devices is the use of risk management to identify and remove risks associated with the manufacture and clinical use of the product. Unavoidable risks are required to be mitigated as far as possible and the manufacturer must take steps to ensure that any remaining unavoidable risks are managed, for example, by including appropriate instructions for use and necessary warnings for the patient, the clinician and others who may come into contact with the device in its intended use. In the same way as for medicinal products, the maintenance of a medical device on the European market necessitates implementation of systems for seeking information on adverse events arising from the product in clinical use, regular reassessment of the information underpinning the device’s risk assessment and communication with the notified body if the product is modified in any significant way. The sale of a non-CE-marked device is only permitted in certain circumstances, including conduct of clinical investigations prior to CE-marking the device and the production of a custom-made device manufactured to a specific physician’s order for a particular patient.



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As for medicinal products, a full discussion of the legal basis, conformity assessment procedures and data requirements for CE marking of medical devices is beyond the scope of this chapter; further information sources are identified in Section 4.4.

4.2.2 How tissue regeneration products are regulated What materials are regulated? Products used in tissue regeneration are regulated according to the scope of the directives into which they fit. In common with all other medical/ healthcare products, the regulation of tissue regeneration materials, and thus the legislation into which the product fits, is determined basically by the way in which the product achieves its intended purpose. There are two important principles contained within this statement: first that the regulations apply to ‘products’, i.e. commercially available items in their own right and not components of some larger item. Individual materials including scaffolds, matrices, chemical or biological agents, are not regulated as medicines or medical devices unless they are marketed specifically as a finished product with a medical purpose that meets the definitions in either the medical device or medicinal product legislation. The second principle is that the product must be intended by its manufacturer/marketer to have a medical purpose and not just a technical function. For example, a plastic film on which cells are cultured can be freely sold for research and manufacturing purposes but it is not regulated in itself as a medicine or a device because its intended purpose is to provide a surface for culture of cells in the laboratory. If, however, that same piece of plastic film were to be marketed as a commercial product for the intended purpose of covering and protecting wounds, it would be regulated as a medical device, since it now has a medical purpose. The situation becomes more complex when we consider how the plastic material might be sold if it were to be combined with living cells before marketing; the plastic component would be considered part of a combined advanced therapy medicinal product. The next section will examine briefly the factors influencing classification of medical products, using a generic electrospun scaffold as an example. Device or medicine? Tissue regeneration is often mentioned in association with ‘tissue engineering’. This is a term that, despite having been around since the early 1980s, has only recently been defined by EU legislation. Tissue engineering is generally accepted to be the manipulation and culture of human cells or tissue ex vivo to produce therapeutic products. In December 2007 the Advanced Therapy



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Medicinal Products Regulation (EC) 1394/2007 (ATMP)1 established a formal definition for ‘tissue engineering’, bringing products that meet this definition firmly within the scope of medicinal product regulation. Article 2.1(b) ‘tissue engineered product’ means a product that: – contains or consists of engineered cells or tissues and – is presented as having properties for, or is used in or administered to human beings with a view to regenerating, repairing or replacing a human tissue. A tissue engineered product may contain cells or tissues of human or animal origin, or both. The cells or tissues may be viable or non-viable. It may also contain additional substances, such as cellular products, biomolecules, biomaterials, chemical substances, scaffolds or matrices. Products containing or consisting exclusively of non-viable human or animal cells or tissue, which do not contain any viable cells or tissues and which do not act principally by pharmacological, immunological or metabolic action, shall be excluded from this definition. Cells or tissues shall be considered ‘engineered’ if they fulfil at least one of the following conditions: – the cells or tissues must have been subject to substantial manipulation, so that biological characteristics, physiological functions or structural properties relevant for the intended regeneration, repair or replacement are achieved. The manipulations listed in Annex 1 [of Regulation EC No 1394/2007], in particular, shall not be considered as substantial manipulations. – the cells or tissues are not intended to be used for the same essential function or functions in the recipient as in the donor The definition is a complex one and several conditions must be met in order for the product to be a tissue engineering product within the scope of the ATMP regulation. The reach of this definition is quite extensive and the majority of products containing viable cells will be classified as an advanced therapy medicinal product (ATMP). Electrospun scaffolds do not in themselves meet the definition of an ATMP because they contain no engineered cells. However if one is sold already seeded with engineered cells, the combined whole product will be regulated as a medicinal product under the ATMP regulation. In this situation the developer of the ATMP will be required to license the whole product via the centralised procedure discussed in Section 4.2.1 before it can be marketed in the EU. The producer of the electrospun scaffold may be the same as, or a different company or organisation from the manufacturer of the ATMP. If they are both part of the same enterprise, access to the data required to support the scaffold will not be an issue, but manufacturers who agree to sell a scaffold to another



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entity who will then use it as part of an advanced therapy medicinal product should be aware that the ATMP manufacturer will require a significant amount of detailed information on the composition, manufacture and quality control of the scaffold. Conversely, if the scaffold is CE-marked and sold on the open market, any ATMP manufacturer who purchases it for their own product development will need to negotiate with the scaffold manufacturer for provision of relevant data and, of course, the scaffold manufacturer is under no obligation to provide such information. If the scaffold is to be a separate commercial product in itself, there are two options for the scaffold manufacturer. The scaffold can be sold as an ordinary material, perhaps as a material for research into a variety of applications, with a description of its potential end uses that could be described in terms such as ‘encourages cell proliferation’, ‘provides a matrix into which cells can migrate’ or ‘guides development of nerve fibres’. These are general terms that in themselves make no specific claims of a medical purpose; they could just as easily apply to a tool for laboratory-based research. If, however, the scaffold manufacturer wants to sell the product for an intended clinical effect in humans, ‘a medical purpose’, then the scaffold becomes a medical device and will thus require a CE-mark. Manufacturers intending to produce and market an electrospun scaffold as an individual product (i.e. not already combined with cells) for a medical purpose will find that their product requires a CE-mark as a medical device before it can be legally sold in the EU. For example, intended uses such as guidance of nerve or tendon fibre to regenerate a patient’s tissue following damage or disease, or as a matrix to encourage proliferation of fibroblasts in wound repair, will be considered medical purposes. In many cases the ability of the scaffold to encourage ingrowth and/or proliferation of the host cells will be related to the bulk chemical composition of the scaffold and to its gross, microscopic or nano-scale surface characteristics. As discussed in other chapters there is intense research into optimisation of both chemical composition and physical structure for the production of electrospun scaffolds for a variety of regenerative medicine end-uses. The issue of the importance of physical and chemical attributes in the design and function of scaffolds dovetails neatly with the question ‘When does an electrospun material become a medical device?’ We have already seen that when combined with engineered cells and sold as a complete regenerative medicine/tissue engineering product, the scaffold becomes a component of an advanced therapy medicinal product: the scaffold itself is not required to be approved separately from the cell-based product. Could the scaffold itself be regulated as a medicinal product? In regulatory terms, a product becomes a medicine or a device when it meets the legal definition set out in the relevant directive. Remembering that both medicines and medical devices must have a medical purpose such as



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prevention or treatment of disease, how do we determine which category a product falls into? The most fundamental concept that differentiates medicines and devices is the idea of mechanism of action: the means by which the principal intended purpose of the product is achieved. Medicines are understood to be acting largely by pharmacological, immunological or metabolic means, with the assumption that they are interacting with the body at a cell receptor or molecular level. In contrast, according to the definition established in the general Medical Devices Directive 93/42/EEC:2 ‘medical device’ means any instrument, apparatus, appliance, software, material or other article, whether used alone or in combination, including the software intended by its manufacturer to be used specifically for diagnostic and/or therapeutic purposes and necessary for its proper application, intended by the manufacturer to be used for human beings for the purpose of: – diagnosis, prevention, monitoring, treatment or alleviation of disease – diagnosis, monitoring, treatment, alleviation of or compensation for an injury or handicap – investigation, replacement or modification of the anatomy or of a physiological process – control of conception and which does not achieve its principal intended action in or on the human body by pharmacological, immunological or metabolic means, but which may be assisted in its function by such means so medical devices are expected to achieve their main purpose principally in a physical, structural or mechanical way. Thus the key attributes of the electrospun scaffold, which has been designed to provide an optimised surface environment for growth of the recipient’s cells to facilitate regeneration of tissue, are physical and surface characteristics that place its mechanism of action firmly within the scope of medical device legislation. It is possible that the manufacturer may wish to coat the scaffold surface with a material that will enhance cell proliferation. Such materials may include human-derived proteins such as growth factors, or materials such as collagen or fibronectin which may be human or animal-derived. The inclusion of such materials can have a huge impact on the way the product is classified for the purpose of regulation. Referring back to the last part of the definition of a medical device: ‘… which does not achieve its principal intended action in or on the human body by pharmacological, immunological or metabolic means, but which may be assisted in its function by such means’, this phrase recognises that medical devices may be combined with pharmacologically active materials to assist their function provided that



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the pharmacological action is secondary to the main (physical) function of the device. This is referred to as ancillary medicinal action. In considering the classification of a product consisting of an electrospun scaffold and an additional material, the first point to evaluate is the mechanism of action of the added material. If the coating or addition material has the potential to act on the human body in a pharmacological, immunological or metabolic way and meets the definition of medicinal product set out in Directive 2001/83/EC,3 then it is considered to be a medicinal substance within the context of device regulation. All devices that contain as an integral part a medicinal substance with action ancillary to that of the device itself, such that the device function is assisted by the ancillary action of the medicinal substance, are regulated as Class III (highest risk) medical devices. These devices require prior examination of their technical data by a notified body and in addition the safety and usefulness of the ancillary medicinal substance must be verified by a competent authority for medicines before the notified body can issue its approval. One of the most important concepts in regulation of devices and medicines is the intention of the manufacturer and the claims or statements they make in respect of the purpose of the product. For products consisting of a device element and a medicinal element, the description or claims that are being made for the product may be central to the way the product is classified. In order to keep the product within the scope of the medical devices regulations, the overall intended purpose of the product must be consistent with the concept of structural or physical mechanism of action. The medicinal substance must contribute only a secondary function to the product. If the medicinal function of the combination is of primary importance, for example the delivery of a biologically active protein to the patient’s tissues, then the combination is a medicinal product. In many cases the product may have multiple functions and it can be quite difficult to determine how the product achieves its primary intended function. Consider, for example, the case of an electrospun scaffold coated with a peptide containing the exposed arginine-glycine-aspartic acid (RGD) sequence known to be a cell recognition site for numerous adhesive proteins present in the extracellular matrix. Surface immobilised RGD groups enhance cell attachment, so the tissue regeneration biologist may wish to improve the performance of the scaffold by adding the RGD sequence peptide to its surface. Is the combination of scaffold and peptide regulated as a medicinal product as a medical device? The scaffold has a structural mechanism of action, providing guidance and a permissive substrate for cell migration and proliferation. The RGD sequence is known to interact with cells on a molecular level and may well meet the criteria for a medicinal substance when used for a therapeutic purpose. The developer of such a product would need to determine a number of factors before classifying the product, such as:



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∑

Does the peptide containing the RGD sequence meet the definition of a medicinal product as set out in directive 2001/83/EC?3 ∑ If it does, is it liable to act on the body when used in the scaffold? ∑ Are the scaffold and peptide combined together in an integral product? ∑ How is the combination primarily intended to function: as a physical guide and support for the patient’s cells or by providing a molecule that enhances cell attachment? The classification of products as medicinal products and medical devices is a complex activity that requires not only a thorough understanding of the directives and associated guidelines and their interpretation but also a clear grasp of the intended purpose of the product and the means by which the product achieves its functions. Standards or guidelines? The two regulatory frameworks make extensive use of standards and guidelines, but these documents have different functions and play very different roles in the development of medical products. Medical device regulation uses standards to facilitate the development of devices. Standards are produced by national standards organisations such as the British Standards Institution (BSI, UK) and the Deutsches Institut für Normung (DIN), by regional standards bodies such as the European Committee for Standardisation (CEN) and by international standards organisations such as the International Standards Organisation (ISO) and the International Electrotechnical Commission (IEC). Technical standards are established norms or requirements. They are formal documents that establish uniform engineering or technical criteria, methods, processes and practices and are in wide use in almost every sphere of industry and research. ISO 90016 is a commonly seen standard relating to quality management in the design and manufacture of products, goods and services and many companies across all industries and services have achieved certification by this international standard. In medical device regulation, standards have a much more significant function. Section 4.2.1 introduces the concept of ‘essential requirements’ to which a medical device must comply in order for the manufacturer to claim conformity with the directive. The manufacturer can demonstrate compliance with what are termed ‘harmonised standards’ as evidence that the product conforms to the directive. A harmonised standard is a European standard, prepared by CEN (or the European Committee for Electrotechnical Standardisation CENELEC for electrical products) under the mandate of the European Commission with the



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purpose of supporting the essential requirements of a New Approach directive. So for the medical device manufacturer, compliance with harmonised standards provides a straightforward way of meeting the requirements of the directive: if the requirements of the standard are met, there is a legal presumption that the product will also meet the equivalent requirements of the directive. In the medical device field, standards cover specific aspects, for example the quality of surgical gloves or safety and performance requirements for pulse oximeters. There are also standards covering general aspects applicable to all devices such as requirements for biological evaluation of medical devices, or the labelling of devices. For example, if a device is to be supplied terminally sterilised by gamma irradiation, meeting the requirements of the harmonised standards in Table 4.1 will allow the manufacturer to demonstrate compliance with the essential requirements relating to sterility of medical devices. Depending on the class of device and the route of conformity assessment chosen by the manufacturer, some combination of the design, manufacture and testing of medical devices should be performed in accordance with a quality system. The use of the harmonised standard ISO EN 13485,7 which adapts the general quality system requirements of ISO 90016 to specific requirements for medical devices, will ensure that the manufacturer’s processes and systems are compatible with the requirements relating to quality systems for CE-marking of devices. Table 4.1 Harmonised standards in relation to gamma-sterilised medical devices Reference

Title of Standard

EN 556-1:2001

Sterilisation of medical devices – Requirements for medical devices to be designated ‘STERILE’ – Part 1: Requirements for terminally sterilised medical devices

EN ISO 11137-1:2006 Sterilisation of health care products – Radiation – Part 1: Requirements for development, validation and routine control of a sterilisation process for medical devices EN ISO 11137-2:2007 Sterilisation of health care products – Radiation – Part 2: Establishing the sterilisation dose EN ISO 11607-1:2009 Packaging for terminally sterilised medical devices – Part 1: Requirements for materials, sterile barrier systems and packaging systems EN ISO 11607-2:2006 Packaging for terminally sterilised medical devices – Part 2: Validation requirements for forming, sealing and assembly processes EN ISO 11737-1:2006 Sterilisation of medical devices – Microbiological methods – Part 1: Determination of a population of microorganisms on products EN ISO 11737-2:2009 Sterilisation of medical devices – Microbiological methods – Part 2: Tests of sterility performed in the definition, validation and maintenance of a sterilisation process



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The references of harmonised standards are published in the Official Journal of the European Communities and set out in the relevant website sections for each sector covered by the New Approach, such as medical devices, toy safety, or recreational craft. The standards are updated as necessary to take account of technical progress and manufacturers must ensure they are working with the current version of each standard. Compliance with harmonised standards is not mandatory (legal obligation) but by developing a medical device in accordance with the relevant standards the manufacture can greatly simplify the process of demonstrating conformity with the directive. In the medicinal products field standards have a much reduced role. The concept of harmonised standards is only legally applied to products covered by New Approach directive and since the medicinal products directives are not New Approach there are no harmonised standards and no presumption of conformity associated with complying with any standard. The manufacturer of a medicinal product may make use of general standards such as ISO 134857 for quality systems or ISO 149718 for risk analysis procedures but observance of these standards conveys no special advantage compared to other approaches that might be taken. Medicinal products directives are supported by guidelines issued by drug regulatory authorities and the International Conference on Harmonisation (ICH) guidelines are technical in content and do not address regional administrative and procedural requirements for approval of medicines, with the notable exception of the Common Technical Document (CTD). The CTD guideline sets out a standard format for organisation and content of the dossier for approval of the medicine and has been accepted in most major markets globally. Additionally, the guideline establishes an electronic CTD (eCTD) specification, the use of which facilitates compilation, review and maintenance of electronic dossiers for medicinal products. The regional and national authorities produce guidelines covering scientific and technical aspects of drug development and guidance relating to the submission and evaluation procedures for medicinal products. Manufacturers developing medicines in Europe need to determine which guidelines are relevant for their product type and either address the guidelines’ recommendations or prepare a justification for omitting them. Compliance with guidelines is not mandatory but the manufacturer will greatly increase the burden of proof needed to approve a medicinal product if the guidelines are not addressed. The EMA provide a useful compilation of scientific and procedural guidelines for all types of medicinal product on its website, see Section 4.4.

4.2.3 Regulatory input into development projects The preceding sections give an insight into the complexity of correctly classifying a product as a medicine or a medical device. Failure to do so correctly at an early stage in development can result in the development



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process being completely inadequate to support the required regulatory approvals, to the extent that the process may need to be completely re-worked under the correct system. It is thus apparent that determining whether the product is likely to be a device or a medicinal product is absolutely critical for development of any product to be used in regenerative medicine, since this will provide the overall pathway for the development process. The types of studies that will be required to support the approval of the product are only a small part of the regulatory consideration since the classification as device or medicine dictates everything in the development. The classification informs the following critical aspects of development: ∑ ∑ ∑ ∑ ∑ ∑ ∑ ∑ ∑ ∑ ∑

∑

which external authorities will need to be consulted and at what stage(s) during the development what approvals will be necessary and the timescales involved in achieving them what data are needed to support the regulatory approval of the product: technical/chemistry data, non-clinical safety and toxicology data, clinical safety and clinical performance/efficacy the requirements, routes of approval, costs and timescales for conducting any necessary clinical studies standards to which studies must be performed identification of relevant guidelines whether it is possible to utilise published information on other related products, or whether all of the necessary support data must be performed by the manufacturer whether the legislative framework gives the innovative developer any protection from copying or generic manufacture what licences the manufacturing site must possess: a manufacturing authorisation for production of medicinal products or a recognised quality system for medical device manufacture what standard of manufacturing environment is necessary; is pharmaceutical good manufacturing practice (GMP) required? the kinds of follow-up procedures necessary to maintain the product in compliance with the appropriate regulatory framework: which systems for vigilance (follow-up reports of adverse reactions and safety problems associated with the product in clinical use) and for assessing changes to the product or its manufacture process the controls and restrictions on advertising and promotional activities.

It can be seen from the foregoing list that the overall timescales, costs and likely returns on the development, value inflection points and options for commercial licensing deals, and in fact the entire business case for a new product, will depend on correct regulatory classification and appropriate development pathway. Regulatory advice should be obtained as soon as



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there is a reasonably well-defined concept and a picture of how the product may be used. Appropriate guidance on claims at an early stage may be sufficient to keep the development of a scaffold material within the medical device framework instead of inadvertently deciding on a strategy that could result in the need for medicinal product licensing, a route associated with much longer development and approval timescales and orders of magnitude increases in costs compared to medical device development. It is particularly important in advanced fields of technology that innovative companies take every opportunity to discuss their development with regulatory authorities. Obviously the company benefits from being able to ask for scientific advice in respect of particular issues and development problems, but it is also true that the regulators themselves can appreciate the chance to find out about new technologies or applications before the dossier lands on their desk. Regenerative medicine is a field undergoing very rapid development and it can greatly facilitate the development of appropriate regulations and sensible interpretations of them if the regulatory agencies themselves are aware of what is on the horizon and how the technology is likely to develop. The development of a positive and proactive relationship with the regulatory agencies is a key responsibility of regulatory affairs professionals. Those who understand how to make full use of both formal and informal interactions with the authorities can significantly improve the efficiency of the development process.

4.3

Future trends

As discussed, the majority of electrospun scaffolds marketed without cells will be regulated as medical devices, for which a mature and well-established regulatory framework exists. The Medical Device Directive2 underwent a significant revision in 2007 to update it, including changes to the essential requirements, the corresponding conformity assessment procedures and the classification of some devices. The revision was published as directive 2007/47/EC.9 Nevertheless feedback to the commission from Member States, clinicians and industry indicates the current system does not always offer a uniform level of protection of public health in the European Union. New and emerging technologies are sometimes difficult to fit within the existing framework. Tissue engineering products now being developed are a case in point which necessitated the development of a new regulation, the Advanced Therapies regulation. Rapid changes in technology and medicine highlight gaps and can lead to a scarcity of expertise in some areas. In addition, recognising that the medical devices market is a global one, the European Community approach to medical devices will need to more closely align with systems in other global regions in order to keep European industry competitive. The commission also recognises that the legal system has been criticised



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as being too fragmented and difficult to follow, and national variations in requirements have crept into individual Member States’ implementation of the device directives, undermining the whole intent of a harmonised European framework. The commission therefore initiated a process for considering the revision of the medical devices legal framework. This process, known as the ‘re-cast’ of the medical devices directives, was initiated in 2008 with a public consultation aimed at European national regulatory authorities, industry, notified bodies, health professionals and patient groups. The process is ongoing at the time of publication and focuses on aspects such as improvements in device vigilance, facilitating the assessment of new technologies and streamlining the regulatory framework for devices. In the context of regenerative medicine it is important to consider some of the applications of cells, tissues and materials that will not be regulated under the current EU framework established by the advanced therapies regulation, as many of those excluded from the regulation will continue in the legislative vacuum that the introduction of the ATMP Regulation was designed to fill. Because of the prohibition of use of human cell and tissuederived materials in medical devices, some applications of electrospinning technology which may seek to combine the benefits of electrospun materials with human cell-derived proteins or other materials will continue to be the subject of a negative impact by the absence of a formal EU-level regulatory system. Table 4.2 illustrates some types of regenerative medicine product and how they may be regulated. Products consisting exclusively of non-viable cells and tissues without a primary immunological, metabolic or pharmacological mode of action are excluded from the regulation. Non-viable human tissues such as demineralised bone matrix will therefore be excluded from the ATMP regulation.1 This means that until an alternative means of regulating these products, such as amendment of the Medical Device Directive,2 is introduced, they will remain subject to national rules or unregulated, as is currently the case. Non-viable human tissues or cells that do demonstrably have a primarily immunological, metabolic or pharmacological mode of action (for example a devitalised matrix designed to deliver cytokines or other proteins produced during culture) would be regulated under the ATMP Regulation.1 A combination ATMP is now defined as an integral combination of a device as defined in 93/42/EEC2 or 90/385/EEC4 and a human cell or tissue part in which the cells or tissues are viable or the cells or tissues are non-viable but their action is primary to that of the device part of the combination. A combination of device and non-viable cells or tissues is only an ATMP if the cells/tissues contribute a primary immunological, metabolic or pharmacological mode of action of the product. If the cells/tissues contribute only an ancillary function then the combination would be unregulated/subject to national rules,



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Table 4.2 Regulation of regenerative medicine products Product type Example Regulatory framework Products containing • ‘engineered’ cells or tissues: • Cells or tissues that • have been expanded, cultured or manipulated to produce the desired • biological effect • Cells or tissues administered to perform a function other than their normal function in the donor

Autologous chondro- Medicinal product cytes expanded in (advanced therapy) culture for reimplantation Allogeneic keratinocytes expanded in culture for burns treatment Fibroblast-derived dermal matrix for vascular repair

Products consisting of non- • viable human or animal tissues, in which the intended function of the product is achieved by pharmacological, immunological or metabolic means

Extracellular matrix Medicinal product derived from culture of (advanced therapy) human cells, for delivery of cytokines or other proteins

Products containing a • Extracellular matrix Medicinal product combination of non- cultured on an (combined viable human or animal implantable scaffold advanced therapy) tissues and one or more medical devices, in which the intended function of the product is achieved by pharmacological, immunological or metabolic means Products containing • Unexpanded population human cells or tissues of stem cells isolated that do not meet the from bone marrow definition of ‘engineered’ or blood cells

Not an advanced therapy medicinal product, but may still meet the primary definition of a medicinal product in 2001/83/EC

Products containing • Demineralised or non-viable human cells purified human bone or tissues, that do not act principally by pharmacological, immunological or metabolic means

Unregulated at the EU level. May be subject to specific national regulations in some Member States

Combination of non- • Electrospun scaffold • Unregulated at the viable human tissue or coated with peptide EU level. May be derivative and medical or growth factor to subject to specific



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Table 4.2 Continued Product Type Example Regulatory framework device in which the intended function of the product is achieved • by mechanical or structural means

enhance cell adhesion national regulations or proliferation in some Member Joint implant coated States with human tissue that encourages ingrowth of recipient’s bone cells

Products containing • Bovine collagen non-viable animal wound dressing cells or tissues in which the intended function of the product is achieved by mechanical or structural means

• Medical device

Combination of non- • Porcine collagen or viable animal tissue derivative coated and medical device onto a scaffold mesh in which the intended function of the product is achieved by mechanical or structural means

• Medical device

Note: this table is for illustration only. Professional regulatory guidance should be obtained for determination of the classification of all products used in regenerative medicine.

since the Medical Device Directive2 currently excludes human cells/tissues and therefore such a combination could not be CE marked at present. Thus the earlier example of a combination of an electrospun scaffold, coated with a human-derived peptide containing the RGD sequence to enhance cell adhesion to the scaffold, is currently unregulated at the EU level. Combinations of a device and non-viable cells/tissues that act in a manner secondary/ancillary to that of the device component are excluded from the ATMP regulation. The re-cast of the medical device directive seeks to address this regulatory gap by proposing an amendment removing the exclusion of non-viable human cells, tissues and derivatives from the scope of the directive. It is to be hoped that this amendment will be timely in order to support the commercialisation of the widest possible range of products based on electrospinning techniques.

4.4

Sources of further information and advice

This section contains references to source materials including the text of relevant directives and access points for guidelines and standards documents.



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Medicinal product directives and regulations Medicinal products – EU Commission website: http://ec.europa.eu/enterprise/sectors/pharmaceuticals/documents/eudralex/ vol-1/index_en.htm Text of Directive 2001/83/EC as amended: http://ec.europa.eu/enterprise/sectors/pharmaceuticals/files/eudralex/vol-1/ dir_2001_83_cons/dir2001_83_cons_20081230_en.pdf Text of Regulation 1394/2007/EC: http://ec.europa.eu/enterprise/sectors/pharmaceuticals/files/eudralex/vol-1/ reg_2007_1394/reg_2007_1394_en.pdf Medical device directives Medical Devices – EU Commission website: http://ec.europa.eu/enterprise/sectors/medical-devices/regulatory-framework/ index_en.htm Text of Directive 93/42/EEC as amended: fttp://eur-lex.europa.eu/LexUriServ/LexUriServ.do?uri=CONSLEG:1993L 0042:20071011:en:PDF Re-cast of medical devices directives – EU Commission website: http://ec.europa.eu/enterprise/sectors/medical-devices/documents/revision/ index_en.htm Medicinal product guidelines European and ICH scientific guidelines – EMA website: http://www.ema.europa.eu/htms/human/humanguidelines/background.htm European procedural guidelines – EMA website: http://www.ema.europa.eu/htms/human/raguidelines/intro.htm European Commission website – application procedures and guidance: http://ec.europa.eu/enterprise/sectors/pharmaceuticals/documents/eudralex/ vol-2/index_en.htm Advanced therapies – EU Commission website: http://ec.europa.eu/enterprise/sectors/pharmaceuticals/human-use/advancedtherapies Advanced therapies – EMA website: http://www.ema.europa.eu/htms/human/advanced_therapies/intro.htm ICH homepage; http://www.ich.org/home.html



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Medical device guidance and standards Index of device guidance documents – EU Commission website: http://ec.europa.eu/enterprise/sectors/medical-devices/documents/guidelines/ index_en.htm EU Commission information on standards in relation to New Approach directives: http://ec.europa.eu/enterprise/policies/european-standards/links/index_ en.htm Current list of harmonised standards for medical devices published on EU Commission website: http://ec.europa.eu/enterprise/policies/european-standards/documents/ harmonised-standards-legislation/list-references/medical-devices/index_ en.htm

4.5 1

2 3 4 5 6 7 8 9

References

Regulation (EC) No 1394/2007 of the European Parliament and of the Council of 13 November 2007 on advanced therapy medicinal products and amending Directive 2001/83/EC and Regulation (EC) No 726/2004. Council Directive 93/42/EEC of 14 June 1993 concerning medical devices. Directive 2001/83/EC of the European Parliament and of the Council of 6 November 2001 on the Community code relating to medicinal products for human use. Council Directive 90/385/EEC of 20 June 1990 on the approximation of the laws of the Member States relating to active implantable medical devices. Directive 98/79/EC of the European Parliament and of the Council of 27 October 1998 on in vitro diagnostic medical devices. ISO 9001: Quality Management Systems – Requirements. ISO 13485: Medical Devices – Quality Management Systems – Requirements for Regulatory Purposes. ISO 14971: Medical devices – Application of risk management to medical devices. Directive 2007/47/EC of the European Parliament and of the Council of 5 September 2007 amending Council Directive 90/385/EEC on the approximation of the laws of the Member States relating to active implantable medical devices, Council Directive 93/42/EEC concerning medical devices and Directive 98/8/EC concerning the placing of biocidal products on the market.



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5

Bone tissue regeneration

A. B a s s i, J. G o u g h, M. Z a k i k h a n i and S. D o w n e s, The University of Manchester, UK

Abstract: The area of bone tissue engineering is a broad field expanding at a rapid rate. Electrospinning is becoming the first choice of many researchers who are considering producing scaffolds that could be used as synthetic bone graft substitutes. This versatile, fibre-forming technique is ideal as the resulting submicrometre fibres mimic the structure of the natural extracellular matrix found in bone. Poly(e-caprolactone) is commonly used in bone tissue engineering; the bioactivity of this polymer can be further improved through the incorporation of bioactive agents such as hydroxyapatite. Osteoporotic patients are often treated in the same way as healthy individuals when surgical intervention is required. For these patients, this can lead to fracture or failure at the site of reduced bone mass. The design of a material which reduces the activity of osteoclasts in osteoporotic patients would be favourable in increasing bone mass. Key words: bone graft substitute, osteoporosis, poly(e-caprolactone), submicrometre fibres.

5.1

Introduction

Bone is a highly vascularised component of the skeletal system and is considered to be a living tissue which undergoes continuous remodelling. The bone remodelling process is a homeostatic balance between bone deposition and bone resorption by osteoblasts and osteoclasts, respectively. Bone plays an essential role in protecting the internal organs of the body from damage and trauma, as well as ensuring the skeleton has adequate load-bearing capacity. For most bone defects, such as fractures, bone has the capacity to regenerate itself spontaneously. However, for larger defects such as non-unions, healing can be problematic and can often require a bone graft or bone graft substitute to aid in the healing (Doblaré et al., 2004). Currently autografts are considered the ‘gold standard’, where healthy bone tissue is extracted from the patient (most commonly from the iliac crest in the hip) and implanted at the required site. Autografts integrate reliably within the host tissue and possess osteoinductive and osteoconductive properties (Finkemeier, 2002). Definitions of osteoinductive, osteoconductive and osteointegration are given in Table 5.1. Autografts, however, impose significant limitations, including pain and donor site morbidity and there is a limited supply of healthy tissue 93 © Woodhead Publishing Limited, 2011


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Table 5.1 Definitions of common terms used to describe the biological properties of materials used in bone tissue engineering Term

Definition

Osteoinductive

A material capable of inducing the recruitment, stimulation and differentiation of pluripotent stromal cells into osteoblasts.

Osteoconductive

Materials which promote bone formation and newly deposited bone conforms to the original bone matrix and exhibits the same composition, shape and surface texture.

Osteointegration

The direct anchorage of an implant by the formation of bone tissue around the implant.

Source: Wilson-Hench, 1987; Albrektsson et al., 1981

(Burchardt, 1983). Donor site morbidity was recently estimated to be as high as 44% (Hierholzer et al., 2006). The second choice for surgeons is the use of allografts, where healthy tissue is extracted from a donor patient and implanted into the host patient. There is a potentially unlimited supply of allograft material, however, there can be problems with the transmission of diseases, poor graft resorption and rejection by the host body (Perry, 1999; Lewandrowski et al., 2001 and Buck et al., 1989). Demineralised bone matrix (DBM) is a common choice for surgeons and is prepared from allograft bone. The use of DBM as a graft was first discovered by Urist (1965) in a revolutionary study where new bone formation was elicited from an intramuscular injection of DBM. Manufacturers of DBM often add different carriers such as glycerol, hyaluronic acid and gelatin to improve the osteoinductive properties of the matrix. However, one of the major drawbacks of DBM is the lack of mechanical support provided at the site of implantation, therefore limiting potential load-bearing applications. Although autografts and allografts are successfully used for the treatment of osseous defects, more recently, research has concentrated on the formation of synthetic bone graft substitutes, materials which are composed of non-bone substance but have osteoconductive properties. Commonly used synthetic bone graft substitutes include calcium phosphates, hydroxyapatite, bioactive glass, polymers and composites. Synthetic bone graft substitutes eradicate the problems associated with the transmission of diseases but do not possess the osteoinductive properties exhibited by autografts and demineralised bone matrix. The lack of osteoinductive properties can limit the repair of more demanding bone defects (Lane et al., 1999).

5.2

Principles of bone biology

Bone is composed of cells and extracellular matrix (ECM). The organic matrix of bone is composed primarily of type I collagen (approximately 90–95%); © Woodhead Publishing Limited, 2011


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type I collagen serves as a template upon which mineral is deposited. Collagen is organised in a three-dimensional porous network composed of collagen fibres that form a hierarchical structure (Kadler et al., 1996). The final 5% of organic matrix is composed of proteoglycans and noncollagenous proteins such as osteocalcin. Osteocalcin comprises up to 15% of non-collagenous protein and is the most abundant protein found in mature bone (Price et al., 1976). It is thought that osteocalcin is involved in bone turnover and mineralisation (Ducy et al., 1996). Within the organic matrix there is an inorganic phase, which is primarily composed of hydroxyapatite nanocrystallites. The inorganic phase of bone provides the compressive strength whereas the organic matrix provides the corresponding tensile properties. Bone is classed as an organic–inorganic nanocomposite, which has features in the nanoscale (Olszta et al., 2007). Within bone there are three types of cells present: osteoblasts, osteoclasts and osteocytes.

5.2.1 Osteoblasts and bone formation Osteoblasts are mononucleate cuboid cells that are responsible for bone formation. Osteoblasts originate from immature mesenchymal stem cells, which can also differentiate and give rise to chondrocytes, muscle, fat, ligament and tendon cells (Aubin and Triffitt, 2002). Mesenchymal stem cells undergo several transcription steps to form mature osteoblast cells. Bone morphogenic proteins (BMPs) are thought to control the commitment of mesenchymal stem cells to the osteoblast phenotype. BMPs are members of the transforming growth factor-b (TGF-b) superfamily of proteins, which act as morphogens influencing fundamental processes such as neurogenesis, development of the kidney, gut and tooth, as well as bone formation (Urist, 1965). Once osteoprogenitor cells start to differentiate into osteoblasts, they begin to express a range of genetic markers: they secrete collagen I which is essential for later mineralisation of hydroxyapatite (Young et al., 1992). The collagen excreted forms osteoids, the osteoblasts cause calcium salts and phosphorous to precipitate from the blood and bond with the newly formed osteoid to mineralise the bone tissue. Osteoblasts also produce alkaline phosphatase; alkaline phosphatase is an enzyme that is involved in the mineralisation of bone. It is an early marker of osteoblast differentiation and its increased expression is associated with the progressive differentiation of osteoblasts (Atsushi et al., 2003). Osteoblasts have oestrogen receptors that allow them to promote the number of osteoblasts in order to increase collagen production. When osteoblasts cease to make new bone they can become trapped within the matrix and differentiate into osteocytes (Abu-Amer and Tondravi, 1997).



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They cease to generate osteoid and mineralised matrix and instead act in a paracrine manner on active osteoblasts. Some osteoblasts remain on the surface of the new bone and differentiate into inactive bone-lining cells and the remainder undergo apoptosis and disintegrate (Abu-Amer and Tondravi, 1997).

5.2.2 Osteoclasts and bone resorption In contrast to osteoblasts, osteoclasts are multinucleated cells (Fig. 5.1) derived from the haematopoietic stem cell line. Osteoclasts have the ability to resorb mineralised bone fully. Osteoclast formation requires the presence of proteins, RANK-L (Theill et al., 2002) and macrophage colony stimulating factor (M-CSF) (Ross and Teitelbaum, 2005). Neighbouring osteoblast cells produce these membrane bound proteins; direct contact is required between osteoblast cells and osteoclast precursor cells. M-CSF acts by binding to its receptor on the osteoclast precursor cell, increasing the number of precursor cells available to mature into osteoclasts. RANK-L interacts with its receptor RANK on the osteoclast precursor cell. This leads to the activation of nuclear factor-kb (Anderson et al., 1997) and NFATc1 (Asagiri et al., 2005), which stimulate the development of differentiated osteoclasts. Osteoblast cells also release a protein called osteoprotegerin (OPG). OPG acts as a decoy and binds to RANK-L, preventing its interaction with RANK (Ludwika et al., 2008) and subsequently preventing osteoclast formation. The release of OPG also prevents excessive degradation of the matrix. At an active bone resorption site, osteoclasts form a specialised cell membrane called the ruffled border, which is in contact with the surface of

(a)

(b)

5.1 (a) Light microscope image of differentiated osteoclasts after 7 days in culture on glass coverslips. The white arrows indicate the presence of multinucleated osteoclast cells. (b) Tartate resistant alkaline phosphatase (TRAP) positive stained osteoclast cells at day 7. TRAP is secreted specifically by osteoclasts and is illustrated in the darker regions of the above image, the positive stain confirms the presence of osteoclasts (x10 magnification, scale bar = 100 µm).



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the bone tissue (Faccio et al., 2003). The ruffled border facilitates removal of the bone matrix and is a morphological characteristic of an osteoclast that is actively resorbing bone. The osteoclast attaches to the underlying bone to form a sealing zone (Faccio et al., 2003). The sealing zone is an actin-rich area surrounding the ruffled border. Within the sealing zone, the ruffled border closely interacts with the bone matrix (Zou et al., 2007). The area beneath the osteoclast is often referred to as the ‘resorptive pit’. Once attached, osteoclasts release hydrogen ions into the resorptive pit (Vaananen et al., 2000). This causes the mineralised bone matrix to become acidic and it is consequently broken down into Ca2+, H3PO4, H2CO3 and water (Vaananen et al., 2000).

5.3

Strategies for bone regeneration

In the past, research has focused on the development of bioinert materials (Langer and Vacanti, 1993), that is materials that have no effect on the surrounding tissue. More recently, focus has shifted to bioactive materials, which interact with the surrounding tissue and encourage bone regeneration and osteointegration (Hench and Polak, 2002). There are a number of synthetic bone graft substitutes available, however none have been found to match or better the properties associated with autologous bone grafting and demineralised bone matrix. The ideal bone graft substitute must have certain characteristics and properties in order for effective bone formation to occur, as shown in Table 5.2. Among these characteristics, the design of an osteoinductive scaffold is the most challenging.

5.4

Fabrication of scaffolds for bone tissue engineering

A number of techniques such as phase separation (Ma and Zhang, 1999), drawing (Ondarcuhu and Joachim, 1998) and self assembly (Whitesides and Grzybowski, 2002) have been used to produce three-dimensional scaffolds. These techniques give rise to highly porous structures but fail to mimic closely the extracellular matrix found in bone. Over the last decade, electrospinning has become an emerging field in producing scaffolds for bone tissue engineering. Electrospinning is a versatile and inexpensive technique, which gives rise to fibres with diameters in the range of 10–1000 nm. Electrospinning parameters and solution properties can be tailored to produce fibres with the desired fibre diameter and pore size. Fibre orientation can be modified through the choice of collector plate; a grounded collector plate gives rise to random orientated fibres whereas a high speed rotating mandrel gives rise to highly aligned fibres (Murugan and Ramakrishna, 2007). Many degradable and non-degradable polymers have been successfully electrospun including poly



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Table 5.2 Desirable properties required by a bone tissue engineered scaffold for successful bone regeneration Desirable properties

Advantages

Biocompatible

∑ Non-immunogenic ∑ Non-toxic

Porosity

∑ ∑ ∑ ∑

Biodegradable

∑ Temporary structure for the attachment and proliferation of cells ∑ Degrades as new bone is deposited ∑ Eliminates the need for further invasive surgery ∑ Degradation products are non-toxic and easily excreted through metabolic pathways

Ideal porosity ª 90% Allows the infiltration of cells and nutrients Removal of waste products Greater surface area for cell attachment and proliferation

Osteoconductive

∑ Support bone growth and encourage ingrowth of surrounding bone

Osteoinductive

∑ Induces the differentiation of stem cells into mature osteoblasts

Mechanical properties

∑ Stable for surgical handling ∑ Able to withstand the forces exerted by the body

Source: Agrawal and Ray, 2001; Puppi et al., 2010; Lutolf et al., 2003; WilsonHench, 1987; Albrektsson et al., 1981

(e-caprolactone), poly(lactic acid) (PLLA), poly(methyl methacrylate) and poly(styrene) for applications in tissue engineering (Bosworth et al., 2008; Yang et al., 2005; Chen et al., 2009; Baker et al., 2006). When designing a material to regenerate bone it is important that the extracellular matrix of bone is mimicked as closely as possible, that is both morphologically and mechanically (Ma, 2008). Electrospinning is becoming an increasingly popular method for producing scaffolds for bone tissue engineering, as the structure of the extracellular matrix of bone is comparable to that of an electrospun mesh, as shown in Fig. 5.2. The similarity in structure provides cells with a familiar environment for attachment and differentiation. The response of osteoblasts has been found to be influenced by surface morphology. For example, osteoblasts have been found to respond most advantageously to random orientated structures (Dalby et al., 2007). For this reason randomly collected fibres, as opposed to aligned fibres, are most favourable for use in scaffolds that are to be used in bone tissue engineering. Pore size is an influential factor when designing scaffolds. Pores must be large enough to allow cell infiltration and proliferation. When the pore



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Bone tissue regeneration (a)

99

(b)

5 µm

5 µm

5.2 Structure of the extracellular matrix secreted by osteoblasts in mouse calvarial bone (a, indicated by the white arrows) is comparable to that of electrospun polymer fibres (b) (Fibre diameter, 259±78 nm). The fibre diameter of the extracellular matrix secreted in mouse calvarial has been determined to be around 70 nm.

diameter is too small, cells may be obstructed and cellular penetration is prevented within the scaffold. Rout et al. (1988) reported that a pore size ranging from 75 to 100 mm resulted in the ingrowth of non-mineralised osteoid tissue. Some researchers have suggested that a pore size between 200 and 400 mm would be ideal for bone tissue engineering in vitro (Leong et al., 2003; Schwartz et al., 1995). As well as encouraging osteoblast infiltration it is necessary to stimulate neovascularisation; it should be noted that pores smaller than 150 mm prevent neovascularisation (Ripamonti et al., 1992). This highlights the importance of satisfactory pore sizes for bone regeneration. As well as the importance of pores for cell infiltration, electrospun substrates provide a large surface area for maximum osteoblast attachment and proliferation. This is particularly important as osteoblasts are said to be anchorage-dependent cells. In addition to having a suitable pore size, pores must also possess interconnectivity. Pore interconnectivity directly influences the diffusion of physiological nutrients and gases to cells, as well as the removal of metabolic waste and by-products (Mikos et al., 1993).

5.5

Potential materials for scaffolds

5.5.1 Polymeric materials The choice of material for bone tissue engineering is paramount to the success of a scaffold. Synthetic, biodegradable polymers are a popular choice as the mechanical and degradation characteristics of the material can be tailored to suit the required application, as well as having the ability to fabricate scaffolds in various shapes with the desired morphology. Polymers such as PLLA, poly(glycolic acid) and their co-polymers are inappropriate for use as synthetic bone graft substitutes owing to their short degradation times (1–2



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months) (Shin et al., 2006). The ideal material would degrade at the same rate as new bone being formed and therefore transfer load to the natural tissue. Electrospun poly(e-caprolactone) (PCL) is commonly used in bone tissue engineering (Fujihara et al., 2005; Shin et al., 2004; Thomas et al., 2006). This is because PCL has a long degradation time and phagocytosis of the polymer by macrophages does not occur until the molecular weight of the polymer decreases below 3000 g mol–1 (Pitt et al., 1981). The slow degradation rate of PCL also gives the polymer stable mechanical properties immediately after implantation. Electrospun PCL-based fibrous scaffolds have been found to have desirable physical and biological characteristics; these scaffolds maintain their physical structure in biological fluid and can support the attachment and proliferation of cells. Although PCL has been shown to support the growth of rat bone marrow stromal cells (BMSCs), the initial attachment of cells can be inhibited owing to the hydrophobic nature of the polymer (Jang et al., 2009). Blending PCL with natural polymers or bioactive agents can improve its biocompatibility (Kim et al., 2007); for example, combining gelatin with PCL nanofibres has been shown to improve the penetration of BMSCs into the fibre matrix (Zhang et al., 2004).

5.5.2 Composite materials As mentioned earlier, the inorganic matrix of bone is composed primarily of hydroxyapatite; the incorporation of inorganic materials that resemble this mineral phase is of great interest. Inorganic materials such as hydroxyapatite, calcium phosphate and bioactive glass have all been incorporated into polymeric electrospun fibres (Zijderveld et al., 2005; Nandakumar et al., 2009; Lu et al., 2009) . These systems have been successful in encouraging osteogenesis and improving bone formation; bioactive glass has been shown readily to encourage the formation of a bioactive hydroxycarbonated layer when immersed in biological fluid, which rapidly bonds to tissue (Stevens, 2008). Venugopal et al. (2007) have incorporated nanohydroxyapatite and collagen I into PCL nanofibres which were fabricated by the electrospinning process. The presence of collagen type I improved osteoblast ingrowth into hard tissue scaffolds. PCL/nanohydroxyapatite/collagen I scaffolds have been shown to enhance the osteoblast response, with mineralisation on these scaffolds improving by 55% when compared to PCL/hydroxyapatite alone, suggesting that collagen I has a greater effect on mineralisation (Venugopal et al., 2007). Similar results were obtained by Prabhakaran et al. (2009) who incorporated PLLA instead of PCL. Recently, Nandakumar et al. (2009) coated electrospun poly(ethylene oxide terephthalate) – poly(butylene terephthalate) fibres with a calcium apatite coating. Although there was no difference in the expression of



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alkaline phosphatase in vitro, there was evidence of bone formation in vivo during a subcutaneous implantation when compared to uncoated fibres. This suggests that the calcium phosphate coating improved the bioactivity of the polymer.

5.5.3 Functionalisation of fibres The most effective way of improving the osteoinductive properties of a material is through the incorporation of bone specific growth factors, which can stimulate bone formation. There are various growth factors involved in bone formation and fracture repair. The most important growth factors are part of the transforming growth factor-b superfamily; bone morphogenetic protein (BMP) is the most significant of this group. Growth factors such as BMP improve the commitment of mesenchymal pluripotent cells to the osteoblast phenotype and display the greatest osteogenic and osteoinductive potential. The incorporation of such factors is becoming of great interest in the area of bone regeneration. The potential use of BMPs in bone tissue engineering was first investigated by Wang et al. (1990). The study highlighted a dose–time dependent relationship in bone formation, where increased doses of BMP-2 led to faster bone formation in vivo. Recently, Li et al. (2006), fabricated silk/nanohydroxyapatite/BMP-2 fibres through the electrospinning process. These scaffolds elicited an enhanced osteoblast response, with an increase in calcium deposition and up-regulation of BMP-2 transcript levels, compared to control scaffolds of silk and nanohydroxyapatite.

5.6

Osteoporosis: a growing problem

Although extensive research is being carried out in the area of bone regeneration, little research is concentrated specifically on bone regeneration for osteoporotic patients. Osteoporosis is a chronic illness that affects an estimated 250 million people worldwide. The balance of bone deposition and bone resorption can often be interrupted which can lead to metabolic diseases such as osteoporosis. The disease is characterised by a decrease in bone density, which causes bones to become increasingly porous and brittle, and is caused by an imbalance in bone remodelling; the rate of bone resorption is greater than the rate of bone deposition, which leads to a bone deficit. For this reason, an osteoporotic patient is more likely to suffer a fracture under normal load conditions compared to a healthy individual. There are approximately 4 million osteoporotic fractures each year in Europe alone (Johnell and Kanis, 2006). Osteoporosis can often lead to a reduced quality of life and even increased mortality. When surgical intervention such as a joint replacement or fracture repair is required, osteoporotic patients undergo the same treatment as healthy



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individuals. This can often lead to failure at the implant site, owing to the lack of bone stock. This consequently makes revision surgery problematic owing to a continuing diminishment of bone stock. There is a clinical need for a synthetic bone graft substitute, which has been specifically designed for osteoporotic patients and can be used at sites of surgical intervention to improve bone stock and consequently reduce the risk of further fracture or failure.

5.7

Strategies for the treatment of bone defects

One possible approach to aid joint replacement and fracture repair in osteoporotic patients is to produce a bone graft substitute that increases bone growth at the site where it is required. This could be achieved through the functionalisation of electrospun fibres with a drug mimic to actively promote bone formation and reduce bone resorption. PCL fibres have been fabricated through the electrospinning process; the parameters used are outlined in Table 5.3. These fibres have been functionalised with a novel drug mimic, poly(vinyl phosphonic acid-co-acrylic acid) (PVPA) (Fig. 5.3). PVPA has an active phosphorous–carbon (P—C) pendant group (Fig. 5.4), which mimics the P—C—P backbone found in bisphosphonates. Bisphosphonates are a group of drugs administered orally to increase bone density by reducing osteoclast function in osteoporotic patients. The drug acts by binding to hydroxyapatite in bone via the P—C—P backbone. This active moiety is internalised by resorbing osteoclasts and consequently leads to cell apoptosis (Rogers, 2003). Table 5.3 Electrospinning parameters used to produce sub-micrometre poly(e-caprolactone) fibres. The average fibre diameter was 259±78 nm. Parameters

Parameters used

Solvent Concentration Voltage Needle to collector distance Flow rate

Acetone 10% w/v 20 kV 15 cm 0.05 ml min–1

H2 C

H2 C

H C PO3H2

H C

H2 C

COOH

H C n COOH

5.3 Structural formula of poly(vinyl phosphonic acid-co-acrylic acid) (PVPA). The phosphorous–carbon (P—C) pendant group is the active component of the polymer.



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5.4 Poly(e-caprolactone) (PCL) fibres functionalised with poly(vinyl phosphonic acid-co-acrylic acid) (PVPA). The average fibre diameter is 269 ± 102 nm (scale bar = 5 µm, magnification x5000).

(a)

(b)

20 µm

20 µm

5.5 Human osteoblast cells cultured on PCL/PVPA scaffolds at day 3 (a) and day 14 (b). At day 14 cells have covered the entire surface of the scaffold.

5.7.1 Effect of poly(vinyl phosphonic acid-co-acrylic acid) (PVPA) on bone formation Recent results have found not only that the PCL/PVPA polymer has an inhibitory effect on osteoclasts but it also has a positive effect on osteoblasts and consequently bone formation. Human osteoblasts attach and proliferate on the PCL/PVPA scaffold; after 14 days the entire surface of the scaffold is covered with cells (Fig. 5.5). By having fibres coated with PVPA as opposed to a smooth fibre surface, we observed effective attachment and spreading of osteoblasts; cells were able to spread their filopodia and there was extensive cell-to-cell contact. This observed level of cell adhesion could be due to the hydrophilic nature of the scaffold. Electrospun PCL alone is hydrophobic and often requires treatment in order for effective cell attachment



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and proliferation to occur. The functionalisation of PCL fibres with PVPA, significantly improves the wettability of the scaffold (Fig. 5.6) allowing the facilitated attachment of cells. Cell responses to surfaces are not mediated by a direct contact, but through an interfacial layer formed on the material surface once it is in contact with a physiological environment. Such a layer is created as a result of nonspecific adsorption of ECM proteins. It is suggested that an interfacial layer is rapidly formed on the material surface and PVPA is able to interact with the surrounding environment (Tan et al., 2005). Over the years, researchers have suggested that phosphonate-containing polymers such as PVPA interact with cell-adhesive, naturally occurring proteins, which consequently enhances proliferation and mineralisation; this hypothesis is becoming widely accepted (Gemeinhart et al., 2006). Phosphonate-containing polymers are also thought to encourage differentiation; pre-osteoblasts have been found to differentiate into mature osteoblasts when cultured in the presence of such polymers (Gemeinhart et al., 2006). Previous in vitro studies have demonstrated that cultures of osteoblasts in the presence of PCL/PVPA show increased levels of alkaline phosphatase activity, mineralisation and collagen type I secretion. The gradual development of the osteoblast phenotype is initialised by the deposition of osteoid and when collagen type I is expressed. Collagen type I is continually expressed during cell maturation and in addition alkaline phosphatase is produced. The increased expression of collagen type I in the present study suggests that osteoblast function is enhanced in the presence of PVPA. This further confirms the results of Reinholz et al. (2000) who demonstrated that cultures containing bisphosphonates expressed higher levels of collagen type I. It has been well documented that polymers containing phosphonate pendant groups exhibit higher rates of mineralisation both in vitro and in vivo, and they have been shown to mineralise in such a way that they resemble bone.

(a)

(b)

5.6 Water droplet on the surface of electrospun PCL (a) and PCL functionalised with PVPA (b). Surfaces functionalised with PVPA are hydrophilic (43.3°±1.2) compared to PCL alone (123.3°±10.8).



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The mechanism is not clearly understood, however, it is hypothesised that phosphate ions stabilise calcium ion interactions within the polymer surface, which combined with the presence of phosphorus in natural bone leads to an increase in mineralisation.

5.7.2 Effect of PVPA on bone resorption As mentioned earlier, PVPA contains P–C pendant groups; these pendant groups are in close proximity to each other and therefore mimic the P–C–P backbone found in bisphosphonates. It is proposed that the active P–C groups in PVPA chelate to bone mineral in the same way as bisphosphonates. The active moiety is selectively internalised by resorbing osteoclasts (Russell and Rogers, 1999). Once internalised it is suggested that bisphosphonates inhibit farnesyl pyrophosphate (FPP) synthase, which is a major enzyme in the mevalonate pathway. This inhibition leads to the disruption of osteoclast activities such as attachment, resorption and eventually leads to cell apoptosis (Reszka et al., 1999; Coxon et al., 2001; Rogers, 2003). There is likely to be local accumulation of the drug/drug mimic in the resorptive pit below the resorbing osteoclast. As mentioned earlier, there is a local release of hydrogen ions below the osteoclast cell, which causes mineral dissolution (Vaananen et al., 2000). As the drug mimic is bound to the mineral, osteoclasts are most likely to be exposed to the highest concentration of free PVPA (Russell and Rogers, 1999). Further degradation of the bone matrix occurs when several hydrolytic enzymes break down the organic components of the bone matrix and there is consequently release of PVPA which is internalised by the osteoclasts leading to osteoclast apoptosis.

5.8

Conclusions and future trends

Electrospinning is becoming a fast and emerging field for the development of bone tissue engineering. The electrospun matrix has a structure comparable to the extracellular matrix of bone, providing cells with a similar environment for attachment and proliferation. The properties of potential scaffolds for bone repair are paramount to the success of the scaffold. Electrospinning provides desirable scaffold properties such as high porosity, random fibre orientation and the potential to fabricate biodegradable polymeric scaffolds. Much research has been concentrated on the incorporation of bioactive materials, such as hydroxyapatite and bioactive glass into polymeric nanofibres to improve the osteoblast response. In treating bone deficits at sites of surgical intervention for osteoporotic patients there is the potential to functionalise polymeric nanofibres with drug mimics such as PVPA. These polymeric systems aim to reduce the activity of resorbing osteoclasts and therefore increase bone mass at the required sites. Not only do these scaffolds have



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the desired morphology but they also possess a hydrophilic surface, which encourages cell attachment and proliferation. This would be ideal for the treatment of bone defects in osteoporotic patients. One of the major goals in the future for bone tissue engineering is the design of a synthetic bone graft substitute that exhibits osteoinductive properties. The design of a material which has properties comparable to that of autologous bone grafting is ideal. Possible approaches to achieving this include further development of the incorporation of growth factors such as transforming growth factor-b and bone morphogenetic proteins into electrospun nanofibres. Such systems would not only provide osteoinductive properties for stem cell differentiation but would also provide a matrix that is comparable to the natural extracellular matrix. Another major area that requires attention in the field of bone tissue engineering is the formation of scaffolds which promote angiogenesis. Bone is a highly vascularised tissue, making blood vessel formation essential.

5.9

References

Abu-Amer, Y. and Tondravi, M. (1997). ‘NF-kB and bone: the breaking point’. Nature Medicine, 3, 1189–90. Agrawal, C.M. and Ray, R.B. (2001). ‘Biodegradable polymeric scaffolds for musculoskeletal tissue engineering’. Journal of Biomaterial Research, 55(2), 141–50. Albrektsson, T., Brånemark, P.I., Hansson, H.A. and Lindström, J. (1981). ‘Osseointegrated titanium implants. Requirements for ensuring a long-lasting, direct bone anchorage in man’. Acta Orthopaedica Scandinavica, 52, 155–70. Anderson, D.M., Maraskovsky, E., Billingsley, W.L., Dougall, W.C., Tometsko, M.E. and Roux, E.R. (1997). ‘A homologue of the TNF receptor and its ligand enhance T-cell growth and dendritic-cell function’. Nature, 390, 175–9. Asagiri, M., Sato, K., Usami, T., Ochi, S., Nishina, H., and Yoshida, H. (2005). Autoamplification of NFATc1 expression determines its essential role in bone homeostasis. The Journal of Experimental Medicine, 202, 1261–9. Atsushi, E., Korenori, O., Satoshi, I., Shigeyuki, E., Takayoshi, N. and Yukichi, U. (2003). ‘Effects of a-TCP and TetCP on MC3T3-E1 proliferation, differentiation and mineralization’. Biomaterials, 24, 831–6. Aubin, J.E. and Triffitt, J.T. (2002). ‘Mesenchymal stem cells and osteoblast differentiation’. In: Principles of bone biology. Bilezikian, J.P., Raisz, L.G. and Rodan, G.A. (eds) Academic Press, New York, 59–82. Baker, S.C., Atkin, N., Gunning, P.A., Granville, N., Wilson, K., Wilson, D. and Southgate, J. (2006). ‘Characterisation of electrospun polystyrene scaffolds for three-dimensional in vitro biological studies’. Biomaterials, 27, 3136–46. Bosworth, L., Clegg, P. and Downes, S. (2008). ‘Electrospun nanofibres of polycaprolactone, and their use for tendon regeneration’. International Journal of Nano and Biomaterials, 1, 263–79. Buck, B.E., Malinin, T.I. and Brown, M.D. (1989). ‘Bone transplantation and human immunodeficiency virus: an estimate of risk of acquired immunodeficiency syndrome (AIDS)’. Clinical Orthopaedics and Related Research, 240, 129–36.



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6

Cartilage tissue regeneration

T. H a r d i n g h a m, The University of Manchester, UK

Abstract: Disease and damage in articular cartilage makes it a major health care target for repair and regeneration. In this chapter the potential for cellbased repair is reviewed and the conditions for the culture, expansion and production of cartilage matrix by articular chondrocytes and bone marrow stem cells are presented. The design features of electrospun scaffolds for cartilage repair are set out, based on research experience and the advantages offered by different polymers are assessed, together with the issues of fibre dimensions, cell seeding and matrix porosities. Key words: cartilage, chondrocytes, chondrogenesis, differentiation, nanofibre scaffolds, stem cells.

6.1

Introduction

Cartilage is a tissue that forms the weight-bearing surface of articular joints. It has an important biomechanical function in distributing load onto the underlying bone and also providing a smooth low-friction surface for joint articulation. Cartilage is a wonderful example of a tissue predominantly formed from extracellular matrix (ECM) (Fig. 6.1). Being avascular and aneural it lacks some of the complexity of other tissues, as no vessels or nerves permeate it and it is composed of an expanded and highly specialised ECM, in which is embedded a single cell type, the chondrocyte. These cells, which form only 1–2% of the tissue mass, are very important for its health as they are responsible for producing and assembling the ECM and continue to do so throughout life. The ECM is predominantly formed by a dense network of collagen fibres embedded in a high concentration of proteoglycan. The collagen, mainly type II with lower amounts of type IX and XI, provides tensile strength and the proteoglycan, mainly aggrecan, endows the tissue with compressive resilience. This is achieved because of the biophysical properties of aggrecan, which forms supramolecular aggregates with hyaluronan and link protein and is thereby made largely immobile in the collagen fibre network. Aggrecan is a polyanion with about 8000 negatively charged sulphate and carboxylate groups per mole and at about 50 mg ml–1 (~25 mM) in the tissue this attracts mobile counterions and creates a large donnan osmotic pressure drawing water into the cartilage matrix, which extends the collagen network until tension in the collagen balances the tissue swelling pressure. Cartilage is ~75% water and its compressive resilience 111 © Woodhead Publishing Limited, 2011


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Electrospinning for tissue regeneration Articular surface Superficial zone

Chondrocyte

Middle zone

Pericellular matrix Intercellular matrix

Deep zone Tidemark Calcified cartilage Subchondral bone

6.1 Schematic of articular cartilage. Chondrocytes are embedded in a large extracellular matrix rich in collagen type II and aggrecan, forming 95% of the tissue.

and elasticity result from the combined functions of the fibrillar collagen, which is hydrated and extended by the polyanionic aggrecan. It is therefore important for the function of the articular cartilage that both the integrity of the collagen network is maintained and the high content of aggrecan that surrounds it is maintained in the tissue. In any strategy to replace the tissue by a biological repair, or by a temporary or permanent synthetic substitute it is important that these properties are replicated or mimicked. As a result of this composite structure, with sparse cells embedded in a large and well organised ECM, cartilage has a limited ability to be repaired once physically damaged, as it lacks the vasculature by which the traffic of systemic cells and their interaction with local cells and environment initiate a cascade of repair responses in other tissues. Thus injuries to cartilage, which are common in sports activities, frequently lead to more serious joint pathology such as osteoarthritis (OA), which is the highly prevalent and strongly age related clinically defined degenerative condition of joints, in which cartilage damage and loss is a frequent end-point. There is therefore an increasing need to develop treatments for the regeneration of articular cartilage. Loss of cartilage at the articulating surface during degenerative joint diseases such as OA leads to pain and immobility, which is presently treated by pain relief and eventually by replacement of the diseased joint by an implant. Whilst the outcomes of joint arthroplasty are often favourable, an increased incidence of prosthesis failure over time limits the usefulness of



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this procedure for individuals below the age of 50. Therefore, for younger people a desirable treatment for cartilage damage would be the resurfacing of the joint with a functional replacement tissue. Initially, research has focused on the repair of small acute localised defects in articular cartilage which may occur as a result of sports injury. Whilst the incidence of patients needing such treatment is significantly less than for those with arthritis, the localised defect presents an achievable first step in articular surface regeneration. Cell-based strategies for cartilage repair vary, but generally contain a number of common features. A source of chondrocytes needs to be identified. There are two possible approaches to this issue. One is to deliver cells, either alone or as part of a cell/tissue construct to the site of repair and the second is to deliver a construct without cells, but with an ability to attract chondrogenic cells in situ. The source of these cells may be from drilling through the base of the defect into bone marrow, which is a rich source of stem cells, or by encouraging trafficking of cells from the synovium into the defect. Current clinical practice involves the re-implantation of autologous primary chondrocytes, derived from samples of the patient’s cartilage tissue and delivered without a tissue construct. This has had some success, although limited to small lesions; however there is great interest in generating constructs suitable for larger defects. Autologous bone marrow-derived mesenchymal stem cells are an alternative source of cells.

6.2

Culture of chondrogenic cells for implantation

To implant a fully or partially formed tissue, the cells need to expand and when enough cells are obtained they need to have the potential to form cartilage tissue in a construct ex vivo prior to implantation and to retain the potential to complete the regeneration of cartilage in situ when delivered into the wound site. The phenotypic state of these cells and their potential to be fully chondrogenic has a profound effect on their ability to direct functional tissue restoration. This is an issue, as primary chondrocytes progressively lose their chondrogenic potential in monolayer culture, so cells after extended culture may have a limited capacity for tissue repair. Many studies both in vivo and in vitro have thus examined how to promote the retention of the chondrocyte phenotype, or more precisely, the ability of cells to regain a chondrocyte phenotype when placed in a suitable chondrogenic environment. Work in our laboratory has focused on the in vitro culture of both human articular chondrocytes and human bone marrow mesenchymal stem cells, to understand what controls their ability to synthesise an appropriate cartilagelike extracellular matrix containing a predominantly collagen type II-rich fibrillar network embedded in an aggrecan-rich ECM. This section focuses on the methodologies that have been found to be successful with cartilage and bone marrow sources of human cells and comments on the many factors



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which may enable improved phenotypic performance once the cells are in a fully chondrogenic environment. In order to set out the specification for the use of synthetic components to augment the generation of a chondrocyte-based biological repair, it is instructive to understand what determines the conditions under which chondrocytes are best able to form cartilage matrix.

6.2.1 Culture of human articular chondrocytes Human articular chondrocytes can be isolated from OA cartilage during joint replacement (see Tew et al., 2008). Although these cells are from the intact cartilage in an OA joint they differ in their pattern of gene expression from those of age-matched healthy cartilage (Brew et al., 2010). However, in monolayer culture all chondrocytes from healthy and OA cartilage downregulate the genes expressed by chondrocytes in cartilage (Fig. 6.2) in a process referred to as de-differentiation, and after passage the cells from OA cartilage are difficult to distinguish from cells from healthy cartilage. In early culture up to passage 4–5, which is 16 to 32-fold expansion, the chondrocytes are able to regain the ability to form matrix if placed in a cell pellet culture, or if cultured in a gel such as agarose or alginate. At later passage 8–10 (~1000 fold expansion) the cells show very little ability to produce matrix in similar cultures. Culture of chondrocytes in a monolayer thus involves fast changes in gene expression, which are reversible, but these are superimposed on slow progressive changes, which are not reversed by Passage number 0

1

2

3

4

5

6 Collagen II Collagen I Aggrecan Collagen X L-SOX 5 SOX 6 SOX 9 GAPHD

6.2 Loss of phenotype of chondrocytes in monolayer culture. Gene expression of cartilage matrix proteins and transcription factors SOX9, SOX6 and SOX5 decreases in monolayer culture with passage, and expression of collagen type I increases.



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simply rounding up cells in culture. The key transcription factor SOX9 is expressed in all chondrocytes and is necessary for matrix gene expression. Its expression steadily falls in monolayer culture (Fig. 6.2), but is able to increase when the cells are rounded up, although only at early passage. One of the key differences in chondrocytes cultured in a monolayer on plastic compared to when rounded up, is the altered organisation of the cytoskeleton. Focal adhesions form in chondrocytes on plastic with the development of strong actin stress fibres, whereas in a rounded culture the actin is cortically distributed in the cell with no focal adhesions. The expression of SOX9 follows these changes and when actin stress fibres are disorganised by agents such as cytochalasin D, chondrocytes round up and SOX9 expression increases (Tew and Hardingham, 2006). This process was also mimicked if Rho effector kinase (ROCK) inhibitors were used to disorganise actin and this revealed a mechanism of SOX9 mRNA stabilisation promoted by p38 MAPK pathway activation (Tew and Hardingham, 2006). The bottom line of these studies is that they provide a reason why chondrocytes in monolayer culture have decreased SOX9 expression and are poorly chondrogenic. The importance of these observations is that we can conclude that focal adhesion formation in chondrocytes inhibits chondrogenic matrix formation. This is potentially of great importance in using scaffolds to deliver chondrocytes for cartilage repair as it is clear that scaffolds with fibre diameters larger than cells will provide surfaces on which focal adhesion formation is promoted, whereas fibres with diameters much less than cell diameters may not. This is shown in Fig. 6.3 where it is suggested that macrofibres provide a poor chondrogenic support for chondrocytes because the fibre surface is rigid and the cells form stable focal adhesions and well organised actin stress fibres. This occurs when the fibre dimension is large and a single cell forms multiple attachments to the surface of a single fibre and therefore mimics chondrocyte attachment to a flat rigid culture dish. It contrasts with the situation in which the cell is surrounded by nanofibres with a much smaller diameter, such that each cell forms point contacts with many nanofibres (Fig. 6.3). As the nanofibres can flex independently of each other, the cell senses a more compliant matrix that does not encourage stable stress fibres to form. Nanofibre supports thus more closely mimic chondrocytes in cell aggregates where each cell forms contacts with many other cells and with the matrix and they retain a cortical actin distribution without stress fibres. This concept and the advantages of nanofibres have experimental support from comparisons of chondrogenic culture on fibres of different diameter, which have shown that chondrocytes performed better on nanofibre scaffolds (Li et al., 2006a, 2006b). Based on this evidence it is suggested that in selecting material support for cartilage repair, nanofibres would provide a scaffold with better chondrogenic properties than macrofibres. The only way this may be offset in large course scaffolds is to load the scaffolds with chondrocytes at a very



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Electrospinning for tissue regeneration (a) Macro-fibrous support 10 µm diameter or greater multiple contacts on same fibril, actin stress fibres generated

(b) Nano-fibrous support 1 µm diameter or less multiple contacts on different fibrils, no actin stress fibres generated

Cell Nucleus Actin Focal adhesion

Macrofibre

Nanofibres

6.3 Schematic of chondrocytes attached to macrofibres and nanofibres. (a) Chondrocytes form multiple stable focal adhesions on macrofibres with actin stress fibres formed, which inhibit chondrogenesis. (b) On nanofibres, chondrocytes form multiple attachments to separate fibres giving less stable adhesion and no inhibition of chondrogenesis. The results correlate with the stiffness of the macrofibre surface, contrasting with the compliance of a nanofibre matrix (Discher et al., 2005).

high cell density, such that most cells are in cell–cell contact in aggregates rather than each individually attached to the fibres. Observations similar to those in chondrocytes have also been identified in stem cell responses to their mechanical environment, where it was reported that differentiation could be strongly guided by the mechanical compliance of the matrix to which the cell was attached (Engler et al., 2005; Wang et al., 2009). The principle behind these observations was that cells developed cytoskeletal tension to sense the cell matrix environment and that the compliance of the cell’s matrix regulated different patterns of gene expression. This effect was also extended to the effects of nanofibrous matrices on the differentiation of bone marrow-derived mesenchymal stem cells (MSC), where it was recognised that nanofibres presented a more compliant matrix than macro surfaces, which in comparison were very stiff (Oh et al., 2009; Reilly and Engler, 2010).

6.2.2 Culture of human stem cells The conditions for generating cartilage from human articular chondrocytes apply equally to sources of stem cells directed to differentiate into



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chondrocytes. Stem cells with the capacity for differentiation into different cells of mesenchymal origin, like chondrocytes, osteoblasts and adipocytes have been identified in many adult tissues. Those identified first and worked on the longest are from bone marrow and they form an attractive supply of cells for autologous repair of articular cartilage, that is, using the cells from each patient to heal their own tissues. They also have the potential for expansion over many passages in culture, permitting more than a one million-fold increase in cell numbers and thus potentially forming a bank of cells for use in repairing many different patients’ tissue, using an allogeneic source of cells from a single donor. This approach to transferring cells from an unrelated donor into a recipient patient has not yet been established as a viable technique. There remains some concern about inflammation, immune reactions and rejection, but the background science has shown that bone marrow stem cells are hypo-immunogenic and may even induce tolerance. Attempts are therefore likely to be made to use allogeneic adult stem cells in a number of tissue engineering applications and repairing tissues such as articular cartilage will be one of those that will be explored. Limited experiments with allogeneic and even xenogeneic (cells from a different animal species) have shown that (human) bone marrow stem cells in an electrospun nanofibrous poly(e-caprolactone) (PCL) scaffold tested in a pig model of cartilage repair gave encouraging results, with cell seeded constructs giving the best repair with little evidence of inflammation, rejection or further complication for up to six months (Li et al., 2009). The use of allogeneic cells would enable the banking of cells and facilitate production of tissue-engineered constructs on a larger scale. This would reduce costs and thereby greatly extend treatment to more patients. This development would also profoundly increase the potential for commercialisation of cell-based cartilage repair procedures and progress here is thus keenly anticipated. Other sources of stem cells such as embryonic stem (ES) cells and induced pluripotent stem (iPS) cells could also be used in this way. However, although progress has been made in developing protocols for the efficient differentiation of human ES cells to chondrogenic cells (Oldershaw et al., 2010), their application for cartilage repair and the evaluation of any risks associated with the clinical use of ES cells, leaves the horizon for their use further away than for adult stem cells. Stem/progenitor cells can be easily prepared from the mononuclear fraction of adult human bone marrow aspirate as plastic adherent colony forming cells that proliferate in culture. These cells show multi lineage differentiation potential and this is retained during expansion over many passages in culture. Their potential for differentiation into chondrocytes is greatly helped if they are grown in culture with added basic fibroblast growth factor (FGF2). For differentiation they require factors, such as transforming growth factor beta, (TGFb), bone morphogenetic proteins (BMPs), together



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with the corticosteroid dexamethasone, but as with articular chondrocytes, they are poorly chondrogenic in monolayer culture and require a rounded up morphology to express full chondrogenesis. Human bone marrow stem cells when rounded up in cell aggregate cultures supported by differentiation factors, such as TGFb and dexamethasone, become chondrogenic. If the cells are seeded and cultured as multilayers in Transwells, which is a configuration with a good nutrient supply that favours cartilage matrix production, they form discs of scaffold-free cartilage in two weeks (Murdoch et al., 2007) (Fig. 6.4). The differentiation involves an initial wave of Notch signalling triggered by transient expression of the Notch ligand Jagged-1, but signalling has to be switched off for differentiation to be completed (Oldershaw and

(a)

(b)

6.4 Cartilage formed by the differentiation of human bone marrow stem cells (hMSC) in vitro into chondrocytes. hMSC differentiate into chondrocytes when cultured in chondrogenic medium (serumfree with TGFb3, dexamethasone, ITS+1) for 14 days in a cell multilayer on a porous membrane insert. Immunostaining shows strong collagen type II and weak collagen type I localisation: (a) immunostained for collagen type II; (b) immunostained for collagen type I (Murdoch et al., 2007).



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Hardingham, 2010). In standard protocols bone marrow stem cells rapidly form cartilage matrix, but they also show up-regulation of type X collagen, which is expressed in growth plate chondrocytes and associated with cell hypertrophy, tissue calcification and cell death, which has lead to concern over their use in articular cartilage repair. However it remains unclear whether if the in vitro conditions were different, a more articular phenotype could be generated and some evidence suggests that parathyroid hormone related peptide PTHrP can be used to suppress type X collagen expression, without blocking chondrogenesis. The use of bone marrow or other stem cell sources is attractive for cartilage repair, as it would forgo the need to isolate patient cartilage tissue and to expand the cells, which creates a two stage repair operation with the ensuing costs and delay. More important would be development that enables the use of allogeneic stem cells, that is cells from a separate donor. This would enable the exploitation of stem cell self renewal permitting vast expansion of one source of cells and enabling them to be stored as a cell bank and used for multiple tissue repair operations. This would greatly reduce the cost of each operation and make the procedure more widely available to those in clinical need.

6.3

Electrospun nanofibre scaffolds

It has frequently been said that tissue engineering scaffolds based on nanoscale fibres can mimic the natural extracellular matrix. However, they clearly do so more in terms of the dimensions of the fibres rather than in any detail of surface chemistry or properties (Bonzani et al., 2006; Nisbet et al., 2009). The more important issue is how far cell responses which have evolved to interact with biological ECM can adapt to synthetic scaffolds and enable and promote natural cell behaviour and full phenotypic function. In this context although the surface chemistry of the electrospun fibres may influence the interaction with cells, it will also determine how quickly the fibre surface is covered by adsorbed proteins, from serum, if it is included in the culture, and secreted matrix proteins from the cells. Protein adsorption typically happens in a time scale of minutes so that cell interaction with fibres quickly involves cell surface ligand interactions with proteins, such as fibronectin/ vitronectin and secreted collagens and laminins. These well characterised biological interactions will thus influence cell behaviour on nanofibre scaffolds in combination with the effects of topography and a three-dimensional (3D) nanofibre array. The importance in providing a scaffold for the cell-based repair of articular cartilage is three fold: (1) to provide a suitable environment for the survival and proliferation of living cells; (2) to generate an implant which is robust, easily handled and implanted by the surgeon, with a mechanical competence to withstand joint articulation soon after implantation; (3) to deliver a scaffold with a planned programme of biodegradation, to allow



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loss of the scaffold to be phased in with the generation of natural cartilage tissue to complete a biological repair. The articular cartilage forms a low friction-bearing surface which distributes load onto the underlying subchondral bone. Any repair thus needs to mimic some of these properties for generating a smooth surface with compressive resilience, as well as have the potential for good physical integration with the underlying calcified cartilage or subchondral bone and with any residual adjacent cartilage. The scaffold also needs to be easily seeded with chondrocytic cells and to provide an environment where they survive, proliferate, produce and assemble a fully functional cartilage matrix. Electrospinning of nanofibrous scaffolds has many features which suit the production of constructs with various 3D architectures that match the needs of cartilage repair.

6.3.1 Choice of materials for electrospun scaffolds for cartilage repair The use of a broad range of natural and synthetic polymers has been proposed for tissue engineering cartilage (Table 6.1). This has included biodegradable synthetic polymers widely used to produce implanted medical products, such as poly(l-lactic acid)(PLLA), poly(lactide-co-glycolide)(PLGA) and PCL, which all have FDA approval for use in vivo. More recently, emphasis amongst these has been on PCL as it has a slower in vivo degradation rate with a half life of months, which is better suited to cartilage repair than a half life of only a few weeks or less (e.g. PLGA, Nisbet et al., 2009). Table 6.1 Polymers tested in electrospun scaffolds for cartilage repair Polymer

Fibre diameter (nm)

Reference

Polycaprolactone (PCL)

~700 400–800 ~438, ~519 ~700 ~500, ~3000

Li et al., 2005 Li et al., 2006a Li et al., 2007 Nam et al., 2007 Wise et al., 2009

PCL nano with PLA macro

400–500 nano ~30 mm macro

Thorvaldsson et al., 2008

Polyglycolic acid (PGA)

300–600

Li et al., 2006a

Polylactic acid (PLA)

1000–1500

Li et al., 2006a

Polylactic-co-glycolic acid (PLGA) 500–1000 ~760

Li et al., 2006a Xin et al., 2007

Collagen II

~70, ~496

Shields et al., 2004

Chitin/chitosan

~40 – ~2000

Bhattarai et al., 2005

Silk

200–500

Zhang et al., 2009



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Electrospun PCL scaffolds have been produced and tested for supporting the chondrogenic differentiation of bone marrow stem cells (Li et al., 2005). Randomly oriented fibres were generated with an average diameter 700 nm. The scaffolds were tested for stability in an aqueous culture medium for at least 8 weeks at 37°C. After seeding with bone marrow stem cells, differentiation was tested over 21 days in chondrogenic medium and showed expression of chondrocyte genes and cartilage matrix production. This PCL scaffold also supported other differentiation stem cell pathways (osteogenic and adipogenic). The generation of osteoblasts together with chondrocytes in separate regions of a construct presented the potential for generation of biphasic cartilage/bone constructs for more complex osteochondral repair (Li et al., 2005). Human stem cell attachment and growth, or differentiation into chondrocytes, was also demonstrated on oriented PCL fibres over several weeks of culture (Wise et al., 2009). The density of the fibres and their orientation thus had no major effect on cell proliferation or differentiation into chondrocytes. Stem cell chondrogenesis and osteogenesis have also been reported to be successful on PLGA (85:15) nanofibre scaffolds (Xin et al., 2007). The mean diameter of the fibres was 760 nm with random fibre organisation. Human bone marrow stem cells survived and proliferated on the scaffold and showed more elongated cell attachment by 7 days. Stem cells also showed evidence of differentiation into chondrogenic and osteogenic cells in a selective differentiation medium and began matrix deposition over 7 days. The progeny of stem cells predifferentiated into osteoblasts and chondrocytes also survived and proliferated on the scaffold. The PLGA (85:15) nanofibre scaffold was thus entirely compatible with stem cell culture and proliferation and supported differentiation (Xin et al., 2007). Other exploration has been in biopolymers such as chitosan (partly deacetylated poly(N-acetylglucosamine)) and chitosan/polyethylene oxide, which form fibre systems with a range of dimensions potentially suitable for cartilage repair. However it is unclear what lifetime they would have in vivo and by what route they would be degraded, as vertebrate systems contain no enzymes that can digest the chitin polysaccharide backbone. Cleavage and clearance of chitosan might thus depend on a slow loss of fibrillar structure and free radical cleavage to give slow dissolution and loss from the tissue. The prolonged lifetime of the fibres in the cartilage construct may inhibit the long term generation of a cohesive and integrated repair tissue in this way, although there is currently no evidence to show if this would adversely affect its functional properties. Experience so far of a number of scaffolds is that the different chemistry of their structure has little effect on cell survival and proliferation of many cell types and of chondrocytes in particular. There is also no affect on chondrogenic differentiation of adult stem cells, although there have been few tests on different materials in parallel to enable a more



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quantitative comparison of rates of differentiation or subsequent matrix production. Another natural polymer that has been explored, type II collagen, has the advantage of being the collagen type naturally most abundant in articular cartilage (Shields et al., 2004). However, the form in which it is available for use in electrospinning is as a pepsin-digested product that lacks telopeptide regions normally involved in native collagen fibre formation. The fibres that it forms when it is electrospun therefore have many differences from natural fibres. They also required chemical cross-links, such as with glutaraldehyde, to stabilise the fibres and prevent rapid dissolution. Any advantage implied from it being a cartilage-derived macromolecule is thus rather offset by its non-native structure and the presence of non-biological chemical crosslinks. However, if judged merely as an alternative polymer it was shown to generate a nanofibrous matrix (mean diameter 496 nm), which was slightly coarser (mean 1460 nm), when cross-linked by exposure to 24-hour treatment with glutaraldehyde vapour. The cross-linking was important to retain fibre stability in an aqueous medium. Although the mechanical strength was far less than articular cartilage, the electrospun collagen construct had a tensile strength of 3 MPa (cf. 9–18MPa for cartilage) and although the porosity was limited there was infiltration of seeded chondrocytes and it supported their growth and matrix deposition (Shields et al., 2004).

6.3.2 Choice of fibre diameter and control of construct porosity As outlined above in the biology of chondrocytes, there is evidence that chondrocytes are inhibited from full expression of their phenotype if they are flattened and adhere to a flat rigid surface with strong formation of their actin cytoskeleton into stress fibres anchored by focal adhesions. This strongly suggests that the use of fibres with large diameters comparable to cell dimensions (5–15 mm+), will encourage chondrocyte adhesion and be anti-chondrogenic. Observation of chondrocytes on supports with much smaller fibre diameter, where each cell may make contact with several independent fibres, suggests that this does not lead to focal adhesions and stress fibres across the cell. The benefit of this for chondrocytes has been shown by evidence that when chondrocytes were seeded on fine PCL fibres (average diameter 500 nm) they had fewer stress fibres and accumulated a matrix with more aggrecan and type II collagen, compared to cells of more coarse fibres (average diameter 3000 nm) (Wise et al., 2009). The message is clear that the benefits of nanoscale electrospun fibres are in providing an environment that leads to more cartilage matrix production by chondrocytes. However, there is a challenge in using small diameter fibres with cells as, for example, in creating a random felt structure by



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electrospinning, the porosity decreases as the fibre size becomes smaller and with sub-micrometre fibre diameters the porosity is much less than the cell dimensions. So although polymer concentration and fluid flow in electrospinning can be modulated to create fibres with suitably small diameters (Eichhorn and Sampson, 2005), this creates scaffolds with low porosity and it becomes difficult to load such constructs with cells by gravity or fluid flow such that they permeate the whole depth of the construct. The cells thus tend to sit on and remain within the surface layers, which is undesirable for cartilage generation. Some proliferation of cells into nanofibrous supports has been reported, but this did not appear to generate a uniform construct. To try and address this problem, other methods have been devised to create greater porosity, such as by combining electrospinning with salt leaching (Nam et al., 2007). A setup was devised that delivered salt grains during the electrospinning of PCL nanofibres and this resulted in a construct with an even distribution of salt grains which could subsequently be removed in aqueous solution leaving pores that were interconnected and allowed easy cell infiltration under flow. Other methods of developing electrospun scaffolds with high porosity have included the generation of nanofibre-coated microfibres. This was achieved by co-fabrication of electrospun PCL nanofibres (mean diameter 400–500 nm) onto continuous electrospun PLA fibres (mean diameter 30 mm) (Thorvaldsson et al., 2008). The scaffold generated had a high porosity determined largely by the macrofibre characteristics, but the surface covering of nanofibres would appear to have the potential to confer the biological advantage of a nanofibre support. The porosity was clearly adequate for cell seeding tested with human articular chondrocytes, but its potential requires more thorough testing for cartilage tissue formation and to assess if its survival and rate of biodegradation in vivo is suitable for cartilage regeneration.

6.3.3 Electrospun scaffolds with graded properties Some attempts have also been made to generate scaffolds that replicate some of the anisotropic graded properties found in natural articular cartilage from the superficial zone at the surface down to its interface with the calcified zone and its junction with sub-chondral bone. A combined extrusion/electrospinning approach was used to generate a PCL electrospun scaffold impregnated with a graded concentration of nanoparticles of b-tricalcium phosphate (Erisken et al., 2008). The application was aimed more at the bone/cartilage interface than at pure cartilage, but the potential of the technique would be to enable the graded deposition of other particulate components within an electrospun scaffold to achieve specific spatial or compositional features. Variation in the mechanical properties of electrospun scaffolds has been achieved by using a rotating mandrel to collect fibres during electrospinning (Li et al.,



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2007). The rotation of the target causes the collected fibres to be aligned and whilst the speed of rotation determined the degree of alignment, it caused little change in fibre diameter. However, the alignment of the fibres had a big effect on the mechanical properties of the scaffold with a 33-fold increase in tensile stiffness as the fibre orientation went from random (33% aligned) to maximally aligned (94% aligned). Cell attachment and cytoskeletal organisation were influenced by fibre alignment, but survival and proliferation were unaffected in the cells tested (adult stem cells and human meniscal fibro-chondrocytes). This more parallel organisation of the fibre scaffold was proposed as most relevant to the repair of cartilage tissues with highly anisotropic collagen fibre organisation, such as in meniscal fibro-cartilages and the intervertebral disc.

6.4

Future trends

It is clear that nanofibrous supports can provide an environment able to support excellent chondrocyte growth and differentiation to form a cartilage matrix. At present the clinical delivery of cartilage repair in focal injuries relies on injecting cells without scaffolds and this achieves some success. Future clinical developments using the current technology will proceed to larger areas of trauma-related cartilage damage and towards the early stage of osteoarthritis. It is here where, with larger areas of damage, the need for a material support may become more important and new procedures may be developed that generate material-based constructs for surgeons to implant. Nanofibrous materials may well be part of that plan, but there will be a real need to obtain clinical evidence that any new procedure delivers a clear advantage over current practice before it is likely to be widely adopted. The exploitation of any of the broad range of research strategies summarised in this chapter is thus highly dependent on, first generating clear evidence of superior cartilage repair in animal models and, subsequently, evidence of good outcome in pre-clinical trials. If superior performance can be combined with simplicity and ease of surgical delivery, this is the aim of those of us who would like to see good science and technology benefit as many patients as possible.

6.5

References

Bhattarai N, Edmondson D, Veiseh O, Matsen FA and Zhang M (2005) ‘Electrospun chitosan-based nanofibers and their cellular compatibility’. Biomaterials, 26, 6176–84. Bonzani IC, George JH and Stevens MM (2006) ‘Novel materials for bone and cartilage regeneration’. Curr Opinion Chem Biol, 10, 568–75. Brew CJ, Clegg PD, Boot-Handford RP, Andrew JG and Hardingham TE (2010) ‘Gene expression in human chondrocytes in late osteoarthritis is changed in both fibrillated



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and intact cartilage without evidence of generalised chondrocyte hypertrophy’. Annals Rheum Dis, 69, 234–40. Discher DE, Janmey P and Wang YL (2005) ‘Tissue cells feel and respond to the stiffness of their substrate’. Science, 310, 1139–43. Eichhorn SJ and Sampson WW (2005) ‘Statistical geometry of pores and statistics of porous nanofibrous assemblies’. J Roy Soc Interface 2, 309–18. Engler AJ, Sen S, Sweeney HL and Discher DE (2005) ‘Matrix elasticity directs stem cell lineage specification’. Cell, 126, 677–89. Erisken C, Kalyon DM and Wang H (2008) ‘Functionally graded electrospun polycaprolactone and b-tricalcium phosphate nanocomposites for tissue engineering applications’. Biomaterials, 29, 4065–73. Li WJ, Tuli R, Huang X, Laquerriere P and Tuan RS (2005) ‘Multilineage differentiation of human mesenchymal stem cells in a three-dimensional nanofibrous scaffold’. Biomaterials, 6, 5158–66. Li WJ, Cooper JA, Mauck RL and Tuan RS (2006a) ‘Fabrication and characterization of six electrospun poly(a-hydroxyester)-based fibrous scaffolds for tissue engineering applications’. Acta Biomaterialia, 2, 377–85. Li WJ, Jiang YJ and Tuan RS (2006b) ‘Chondrocyte phenotype in engineered fibrous matrix is regulated by fiber size’. Tissue Eng, 12, 1775–85. Li WJ, Mauck RL, Cooper JA, Yuan X and Tuan RS (2007) ‘Engineering controllable anisotropy in biodegradable nanofibrous scaffolds for musculoskeletal tissue engineering’. J Biomech, 40, 1686–93. Li WJ, Chiang H, Kuo TF, Lee HS and Tuan RS (2009) ‘Evaluation of articular cartilage repair using biodegedable nanofibrous scaffolds in a swine model: a pilot study’. J Tissue Eng Regen Med, 3, 1–10. Lim SH and Mao HQ (2009) ‘Electrospun scaffolds for stem cell engineering’. Advanced Drug Delivery Rev, 61, 1084–96. Murdoch AD, Grady LM, Ablett MP, Katopodi T, Meadows RS and Hardingham TE (2007) ‘Chondrogenic differentiation of human bone marrow stem cells in Transwell cultures: Generation of scaffold-free cartilage’. Stem Cells, 25, 2786–96. Nam J, Huang Y, Agarwal S and Lannutti J (2007) ‘Improved cellular infiltration in electrospun fiber via engineered porosity’. Tissue Eng, 13, 2249–57. Nisbet DR, Forsythe JS, Shen W, Finkelstein DI and Horne MK (2009) ‘A review of the cellular response on electrospun nanofibers for tissue engineering’. J Biomat Applic, 24, 7–28. Oh S, Brammer KS, Li YSJ, Teng D, Engler AJ, Chien S and Jin S (2009) ‘Stem cell fate dictated solely by altered nanotube dimension’. Proc Natl Acad Sci (USA), 106, 2130–35. Oldershaw RA and Hardingham TE (2010) ‘Notch signaling during chondrogenesis of human bone marrow stem cells’. Bone, 46: 286–93. Oldershaw RA, Baxter MA, Lowe ET, Bates N, Grady LM, Soncin F, Brison DR, Hardingham TE and Kimber SJ (2010) ‘Directed differentiation of human embryonic stem cells toward chondrocytes’. Nature Biotech, 28, 1187–94. Reilly GC and Engler AJ (2010) ‘Intrinsic extracellular matrix properties regulate stem cell differentiation’. J Biomechanics, 43, 55–62. Shields KJ, Beckman MJ, Bowlin GL and Wayne JS (2004) ‘Mechanical properties and cellular proliferation of electrospun collagen Type II’. Tissue Eng, 10, 1510–17. Tew SR and Hardingham TE (2006) ‘Regulation of SOX9 mRNA in human articular chondrocytes involving p38 MAPK activation and mRNA stabilization’. J Biol Chem, 281, 39471–9. © Woodhead Publishing Limited, 2011


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Tew SR, Murdoch AD, Rauchenburg RP and Hardingham TE (2008) ‘Cellular methods in cartilage research: primary human chondrocytes in culture and chondrogenesis in human bone marrow stem cells’. Methods, 45, 2–9. Thorvaldsson A, Stenhamme H, Gatenholm P and Walkenstrom P (2008) ‘Electrospinning of highly porous scaffolds for cartilage regeneration’. Biomacromol, 9, 1044–9. Wang N, Tytell JD and Ingber DE (2009) ‘Mechanotransduction at a distance: mechanically coupling the extracellular matrix with the nucleus’. Nat Rev Mol Cell Biol, 10, 75–82. Wang Y, Blasioli DJ, Kim HJ, Kim HS and Kaplan DL (2006) ‘Cartilage tissue engineering with silk scaffolds and human articular chondrocytes’. Biomaterials, 27, 4434–42. Wise JK, Yarin AL, Megaridis CM and Cho, M (2009) ‘Chondrogenic differentiation of human mesenchymal stem cells on oriented nanofibrous scaffolds: engineering the superficial zone of articular cartilage’. Tissue Eng, 15, 913–21. Xin X, Hussain M and Mao JJ (2007) ‘Continuing differentiation of human mesenchymal stem cells and induced chondrogenic and osteogenic lineages in electrospun PLGA nanofiber scaffold’. Biomaterials, 28, 316–25. Zhang X, Reagan MR and Kaplan DL (2009) ‘Nanofibers in regenerative medicine and drug delivery’. Advanced Drug Delivery Rev, 61, 988–1006.



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7

Muscle tissue regeneration

K. D. M c K e o n - F i s c h e r and J. W. F r e e m a n, Virginia Polytechnic Institute and State University, USA

Abstract: Muscle is unique among tissues because it generates its own force, presenting a unique challenge in tissue engineering. New tissue developed from the scaffold must not only have structural integrity, it must also be able to coordinate the protein-based changes necessary for contraction and force generation. In this chapter we will investigate the tissue engineering of muscle, specifically the use of electrospun nanofibres in skeletal muscle tissue engineering. We will focus on the advantages of nanofibrous matrices for muscle regeneration and how investigators combine nanofibres with different types of external stimulation to enhance tissue growth. Key words: muscle, nanofibres, polymers, tissue engineering.

7.1

Introduction to skeletal muscle

Approximately 40% of the human body is composed of skeletal muscle, which is made up of multiple fibres ranging in diameter from 10 to 80 mm (Guyton and Hall, 2006). Skeletal muscle cells can be categorised as myoblasts (undifferentiated cells), myotubes (multiple myoblasts fused together to form a multinucleated construct) and muscle fibres (oriented differentiated myotubes) (Okano and Matsuda, 1997; Cheema et al., 2003; Lefeuvre et al., 1996; Bach et al., 2004). The unidirectional orientation of muscle fibres allows a large force to be generated during contraction (Okano and Matsuda, 1998; Bach et al., 2004). Muscle fibres contain hundreds to thousands of myofibrils; these are composed of two contractile proteins: 3000 actin and 1500 myosin filaments (Guyton and Hall, 2006). Muscle contraction begins when a potential action travels down the motor neuron to the neuromuscular junction (NMJ)–motor neuron-skeletal muscle interface (Arrowsmith, 2007; Lai and Ip, 2003; Chiou-Tan, 2003; McConville and Vincent, 2002; Nogajski et al., 2009; Guyton and Hall, 2006). After the motor neuron is depolarised, acetylcholine is released from the presynaptic membrane and diffuses across to the post-synapse muscle end-plate (Chiou-Tan, 2003; McConville and Vincent, 2002; Arrowsmith, 2007; Ngo et al., 2007; Shear and Martyn, 2009; Nogajski et al., 2009; Guyton and Hall, 2006). Once the threshold level is reached, depolarisation of the muscle leads to calcium ion release that causes the actin and myosin filaments to slide along each other for contraction (Chiou-Tan, 2003; McConville and Vincent, 2002; Arrowsmith, 2007; Shear and Martyn, 2009; Guyton and Hall, 2006). 127 © Woodhead Publishing Limited, 2011


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7.2


Skeletal muscle injuries

This process can be interrupted by traumatic injuries to peripheral nerves (PNs) and/or skeletal muscle (Navarro et al., 2007; Shear and Martyn, 2009; Campbell, 2008; Rodríguez et al., 2004; Bruns et al., 2007; Bian and Bursac, 2009). These types of injuries result in a significant loss of muscle and require either surgical reconstruction or, in severe cases, amputation (Liao and Zhou, 2009), some of the most extreme examples are seen in military personnel during wartime (Clark et al., 2007). The Walter Reed Army Medical Center (WRAMC) recorded the number of surgeries performed in the field and at the hospital on severely injured soldiers either from Afghanistan during Operation Enduring Freedom (OEF) or Operation Iraqi Freedom (OIF) (Clark et al., 2007). Table 7.1 displays the results for both amputation and soft tissue exploration/debridement (removal of dead muscle tissue to prevent infection) (Clark et al., 2007; Anon, 2004). Once a muscle is injured, necrotic muscle fibres are removed by macrophages and satellite cells are activated to help regenerate the skeletal muscle (Bach et al., 2004). However, satellite cell incidence in the tissue is extremely low, only 1–5%, and is dependent on the age and composition of the muscle fibre (Bach et al., 2004). Satellite cells migrate and proliferate at the injured site, but this process results in formation of scar tissue and loss of muscle function (Bach et al., 2004). Autologous muscle transplants and exogenous myogenic cells, satellite cells and myoblasts have also been investigated with little success (Bian and Bursac, 2009). This transplantation also leads to morbidity, loss of function and decreased volume at the donor site (Bach et al., 2004). Intramuscular injections of skeletal myoblasts have also been investigated, but had little effect owing to inadequate distribution and low cell survival rate (Bian and Bursac, 2009). Loss of skeletal muscle caused by prolonged denervation and traumatic injuries currently has no satisfactory method of restoration (Bian and Bursac, 2009; Bach et al., 2004; Liao and Zhou, 2009). These limitations in current treatments for lost or damaged musculature highlight the need for a muscle replacement therapy. There is a serious need for a treatment that can provide function and regenerate lost tissue in order to improve the patient’s quality of life in both the short and long term. Table 7.1 Number of soldiers receiving soft tissue exploration/debridement and amputations both in the field (n = 931) and at the WRAMC (n = 634)

Field WRAMC

Soft tissue exploration/ debridement

Amputation

219 287

114 40

Source: Clark et al., 2007



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7.3

129

Mechanical properties of skeletal muscle

The mechanical properties of skeletal muscle are important for devising a functional skeletal muscle tissue implant. As muscle is an elastic tissue, it is necessary to find an alternative with similar mechanical properties. Toursel et al. (2002) excised the soleus muscle from adult male Wistar rats, chemically skinned it and tensile tested single muscle fibres. The soleus muscle fibres (n = 13) had an average elastic modulus of 93.40 ± 2.17 kPa (Toursel et al., 2002). In an effort to discover whole muscle mechanical properties, our laboratory excised the entire soleus (n = 4) and vastus lateralis (VL) (n = 4) muscles from the hind limbs of two juvenile male Sprague-Dawley rats. The elastic modulus and yield stress were measured at 10% strain/minute. Our findings display similar results for both muscles in respect to elastic modulus and yield stress. We found an average elastic modulus of 48.11 ± 16.19 kPa for the soleus muscle and 51.27 ± 28.48 kPa for the VL muscle. Yield stresses were 17.84 ± 5.83 kPa and 12.10 ± 5.58 kPa, respectively (Table 7.2). Although our measured elastic modulus is almost half of the fibre elastic modulus reported by Toursel, we believe this may be due to differences in strength and composition at the macro versus micro levels. Mechanical properties of skeletal muscle of larger animals are of interest as eventually biomedically engineered replacements will need to be implanted in a human. Morrow et al. (2010) excised the extensor digitorum longus (EDL) muscle (n = 6) from the hind limb of New Zealand White rabbits. The average elastic modulus for tension at 0.05% s–1 strain was 447 (± 97.7) kPa (Morrow et al., 2010). This value is very large compared to the previously discussed values for both rat single fibre and whole skeletal muscle. Our laboratory has similarly tensile-tested two hind limb medial skeletal muscles from a female Yorkshire cross pig. Each muscle was cut into eight strips, tested at 10% stain/minute and then averaged together for a final value, rat muscles were tested at 20% strain/minute. The elastic moduli of the two muscles were similar: 151.83 (± 50.30) kPa and 141.12 (± 42.34) kPa. The yield stresses were also close: 54.00 (± 21.17) kPa and 57.80 (± 11.92) kPa (Table 7.2). The rabbit elastic modulus for the EDL muscle is approximately four times greater than the elastic moduli measured for the medial hind limb skeletal muscles from the pig. Although we cannot rule out the fact that Table 7.2 Elastic modulus and yield stress measured under tension in a PBS bath at 37°C for rat muscles and medial pig skeletal muscle

Rat soleus muscle Rat vastus lateralis muscle Pig medial muscle 1 Pig medial muscle 2

Elastic modulus (kPa)

Yield stress (kPa)

48.11 51.27 151.83 141.12

17.84 12.10 54.00 57.80

± ± ± ±

16.19 28.48 50.30 42.34

± ± ± ±

5.83 5.58 21.17 11.92



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different muscles were measured, we believe it could be due to differences in where the muscle cross-sectional area was measured or the location from which the skeletal muscle was taken.

7.4

Tissue engineering

Tissue engineering (TE) provides a way of combining both engineering and biology in order to ameliorate a current medical problem by creating, repairing and/or replacing damaged tissues and organs (Lo et al., 1996; Zünd et al., 1998; Leor et al., 2005; Okano and Matsuda, 1998; Liao and Zhou, 2009; Sarraf et al., 2002). TE accomplishes this by using a combination of cells, biomaterials and tissue-inducing agents, such as growth factors (Leor et al., 2005; Zünd et al., 1998; Lo et al., 1996). The chosen scaffold material is important in determining the biocompatibility, degradation rate, chemical properties and physical properties (Willerth and Sakiyama-Elbert, 2007; Maquet et al., 2000; McCullen et al., 2007). The material must also encourage cell adhesion, migration, growth and differentiation (Lannutti et al., 2007; Whitaker et al., 2001). Skeletal muscle cells have been grown in vitro on numerous scaffold materials including Saran Wrap®, collagen, acellular matrices, fibrin, poly(l-lactic acid) (PLLA), poly(d-lactic acid)–polyaniline (PDLA–PANi) and poly(e-caprolactone) (PCL) (McKeon-Fischer and Freeman, 2010; Strohman et al., 1990; Bach et al., 2004; McKeon et al., 2010; Conconi et al., 2005; Choi et al., 2008; Kin et al., 2007; Okano and Matsuda, 1998; McKeon-Fischer et al., submitted). TE can be utilised to create a functional implant to replace damaged skeletal muscle; however, there has been little success.

7.4.1 Collagen scaffolds Collagen is being investigated for skeletal muscle tissue engineering as it can be made into either a three-dimensional (3D) gel or sponge to be directly implanted into the muscular defect (Okano and Matsuda, 1997; Okano and Matsuda, 1998; Cheema et al., 2003; Kin et al., 2007). This biomaterial is commonly selected as a scaffold as it is a component of the native extracellular matrix (ECM). Okano and Matsuda (1997) formed 3D collagen gel discs and cultured C2C12 cells (murine myoblast cell line) for 11 days. The C2C12 cells formed myotubes on the collagen gel; however, the gels were very fragile and shrank over time (Okano and Matsuda, 1997). Kin et al. (2007) conducted an in situ study whereby 18 male rabbits had a defect created in their vastus lateralis (VL) muscle and a collagen sponge inserted. Although the collagen sponge degraded at a slower rate than the gel, only a small amount of the sponge remained two weeks post-surgery and had disappeared entirely by three weeks (Kin et al., 2007). The final



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conclusion from this study was that even after 24 weeks of implantation the defect (1.0 ¥ 1.0 ¥ 0.5 cm) was not filled in by skeletal muscle tissue (Kin et al., 2007). The mechanical properties of collagen gels have also been measured through tensile testing. Sheu et al. (2001) measured the elastic modulus of several collagen gels cross-linked with 0–0.2% gluteraldehyde. The resulting elastic moduli ranged from 4.23 to 7.23 kPa (Sheu et al., 2001). Madhavan et al. (2010) cross-linked collagen gels with ethyldimethyl aminopropyl and carbodiimide hydrochloride (EDAC). The elastic modulus increased to approximately 100 kPa, which is within range of the values given previously (Section 7.3) (Madhavan et al., 2010). These differences in elastic moduli could be due to the chemical cross-linker used or batch to batch variations in the collagen. Some other potential problems with using collagen as a scaffold include the possibility of disease transmission from animal to animal or human, its high cost and problems with currently accepted sterilisation techniques affecting collagen mechanical properties.

7.4.2 Electrospinning and myotube formation Unlike gels, electrospun scaffolds provide an extracellular matrix-like base to which cells can attach before secreting their own extracellular matrix (ECM) (Lannutti et al., 2007; Pham et al., 2006). The ECM aids in the attachment, alignment and differentiation of myoblasts (Liao and Zhou, 2009; Huang et al., 2006). In one study, Huang et al. (2006) cultured murine C2C12 myoblasts on electrospun PLLA fibres. They found that the myotubes produced were highly organised, grew along the nanofibre and were longer compared to C2C12 cells grown on unaligned mats. Choi et al. (2008) showed similar results with human skeletal muscle cells grown on PCL-collagen electrospun fibres. The aligned electrospun fibres displayed enhanced myotube formation compared to the unaligned electrospun fibres (Fig. 7.1) (Choi et al., 2008).

7.5

Contractile force

One of the major problems with culturing skeletal muscle cells is the lack of contractile force compared to native muscle. For example, only 1–2% contractile force has been produced when cells were cultured in vitro (Dennis and Kosnik, 2000; Kosnik et al., 2001; Dennis et al., 2001; Vandenburgh, 2001). Although several studies have investigated both mechanical and electrical stimulation to form more mature muscle fibres with better contractility, further research into this area is necessary (De Deyne, 2000; Fujita et al., 2007; Bach et al., 2004; Liao et al., 2008).



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(a)

(b)

7.1 Immunofluorescent human skeletal muscle cells on unaligned (a) and aligned (b) electrospun PCL scaffolds on day 7 (Choi et al., 2008). Reprinted from Choi et al. (2008) ©2008 with permission from Elsevier.

7.5.1 Mechanical stimulation De Deyne (2000) cultured primary rat skeletal muscle cells and introduced a 5% and 10% mechanical strain on the cultures. The 5% passive strain caused the rat muscle cells to align longitudinally but a 10% strain led to the loss of generated myotubes (De Deyne, 2000). A later study by Huang et al. (2006) found that electrospun aligned PLA fibres had the same effect on murine C2C12 myoblasts as seen when an external 5% strain was applied. Liao et al. (2008) also found similar results, longitudinal elongation and fusion of cells, with electrospun aligned PU fibres. Therefore, electrospun scaffolds can be utilised in place of an applied mechanical strain and reduce the external stimuli needed to engineer a functional skeletal muscle tissue.

7.5.2 Electrical stimulation De Deyne (2000) compared his results on the mechanical strain on skeletal muscle cells to ones that were electrically stimulated. He concluded that applying an external low frequency electrical stimulation was the most beneficial as it increased the number of myotubes in culture (De Deyne, 2000). Fujita et al. (2007) found that electrical pulse stimulation increased the C2C12 myotube formation rate in culture. They found that it took just under two hours for the cultured C2C12 cells to gain contractile activity in response to 40 V/60 mm, 24 ms and 1 Hz (Fujita et al., 2007). As a result, electrospinning and electrical activation can be used together to obtain highly aligned fibres that will cause muscle cell elongation, orientation, fusion and striation.



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Conductive elements

Electroactive polymers and metal micro- or nanoparticles can be utilised to increase the conductivity of a polymeric scaffold for possible use with electrical stimulation techniques. The use of conductive elements in scaffolds may allow a lower voltage and current to be used for skeletal muscle cell elongation, orientation, fusion and striation.

7.6.1 Conductive polymers Conductive polymers such as polyaniline (PANi) are of growing interest owing to their unique conductive property which increases cell attachment, proliferation, migration and differentiation (Huang et al., 2007; Huang et al., 2008; Willerth and Sakiyama-Elbert, 2007; Yu LMY et al., 2008; Li et al., 2006; Bidez et al., 2006). Several studies have shown that PANi can be easily electrospun either alone or with another polymer to create a scaffold (Veluru et al., 2007; Yu Q-Z et al., 2008; Khan et al., 2007; Li et al., 2006; Jeong et al., 2008). Several in vitro and in vivo tests have also shown that PANi is biocompatible (Jeong et al., 2008; Li et al., 2006; Wang et al., 1999; McKeon et al., 2010). For example, Li et al. (2006) showed that rat cardiac muscle cells were able to attach, migrate and proliferate on PANi–gelatin electrospun fibres. The average elastic modulus for each of the three PANi:gelatin ratios were 614 (± 75) MPa for 15:85, 1043 (± 184) MPa for 30:70, and 1384 (± 105) MPa for 45:55 (Li et al., 2006). However, these PANi–gelatin scaffolds have elastic moduli that are too large for skeletal muscle. McKeon et al. (2010) also demonstrated skeletal muscle cell attachment and growth on several different concentrations of PDLA–PANi scaffolds (Figs 7.2 and Fig. 7.3). However, the PDLA–PANi scaffolds degraded too quickly for use in skeletal muscle tissue engineering (McKeon et al., 2010). Wang et al. (1999) conducted an in vivo study implanting PANi in adult male Sprague Dawley rats both as a powder and a film. No significant inflammation was seen in any of the histological samples taken over the 50-week study (Wang et al., 1999). Although studies have shown that PANi is biocompatible both in vitro and in vivo, more research is needed to develop a scaffold with the correct characteristics for skeletal muscle.

7.6.2 Gold nanoparticles Polymer/nanoparticle composite scaffolds can be made by either (1) incorporating nanoparticles into a polymer matrix or (2) using the polymer material on the nanoscale level (Yang et al., 2007). Metal nanoparticles are more commonly researched because they introduce unique electrical, optical and catalytic properties (Kim et al., 2005; Wang et al., 2007; Kim and Ahn,



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Cells only 24% (83/17) 24% (80/20) 22% (75/25)

2.8 2.6 Absorbance at 490 nm

2.4 2.2 2.0 1.8 1.6 1.4 1.2 1.0 0.8 0.6 0.4 0.2

Day 1

Day 3

Day 7

Day 14

7.2 Rat primary skeletal muscle growth on 24% (83/17) PDLA–PANi scaffolds, 24% (80/20) PDLA–PANi scaffolds, 22% (75/25) PDLA–PANi scaffolds, and tissue culture polystyrene (TCP) as a control. No significant difference was seen between any of the groups at any of the time points. Data reprinted with permission (McKeon et al., 2010). (a)

(b)

1µm

1µm (c)

2µm

7.3 Scanning electron microscope (SEM) micrographs illustrating cellular attachment on day 14 for the (a) 24% (83/17), (b) 24% (80/20) and (c) 22% (75/25) scaffolds. Images reprinted with permission (McKeon et al., 2010).



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2008; Bai et al., 2008). Gold nanoparticles (Au NPs) have been shown to electrospin easily with several polymers (Bai et al., 2008; Du et al., 2007; Kim et al., 2005; Wang et al., 2007; Bai et al., 2007; Kim and Ahn, 2008). For example, Kim et al. (2005) was able to electrospin Au NPs successfully with poly(ethylene oxide) (PEO). However, for tissue engineering purposes, there is some concern that Au NPs could be toxic as they can easily enter cells (Lewinski et al., 2008). Several studies have researched the cytotoxic effects of Au NPs on different cell lines (Thomas and Klibanov, 2003; Patra et al., 2007; D’Britto et al., 2009; McKeon-Fischer and Freeman, 2010). Thomas and Klibanov (2003) researched the effect of polyethylenimines (PEI) conjugated to Au NPs on simian virus 40-transformed African green monkey kidney cells (COS-7). Cell viability was measured 48 hours after transfection with the PEI–Au NP complexes and ranged from 67–78% (Thomas and Klibanov, 2003). Patra et al. (2007) measured the cell viability of human caucasian lung carcinoma type II epithelial cells (A549 cell line) and human caucasian hepatocyte cells (HepG2 cell line) exposed to Au NPs in the following concentrations: 10 nM, 30 nM, 60 nM and 120 nM. The A549 cells had decreased cell viability with all four Au NP concentrations after 36 hours, whereas the HepG2 cells displayed no difference in viability for any of the concentrations after 72 hours (Patra et al., 2007). D’Britto et al. (2009) measured the effect of PEI films with Au NPs on Chinese hamster ovary (CHO) cells. The CHO cells grown on the PEI films incubated with Au NPs and the PEI films layered with Au NPs and lysine demonstrated much higher mitochondrial activity than the TCP control cells after 48 hours (D’Britto et al., 2009). Therefore, Au NP cytotoxicity is highly dependent on cell type. McKeon-Fischer and Freeman (2010) have incorporated Au NPs into an electrospun polymer scaffold for skeletal muscle tissue engineering to increase the conductivity of the scaffold. Three different concentrations of Au NPs, 7%, 13% and 21%, were electrospun with PLLA. The calculated conductivities for the 7% Au–PLLA scaffolds were 0.05 (± 0.015) S cm–1, 13% Au–PLLA scaffolds were 0.08 (± 0.004) S cm–1, and 21% Au–PLLA scaffolds were 0.09 (± 0.037) S cm–1. A cell study using primary rat skeletal muscle cells seeded on the three Au–PLLA scaffolds and tissue culture polystyrene (TCP) as a control was conducted over a four week period (Fig. 7.4). As all three scaffolds had significantly lower cell growth than the TCP control, a second cell study was performed to determine if the Au NPs were having a toxic effect (McKeon-Fischer and Freeman, 2010). The second cellular study measured the toxicity of cell media on rat primary muscle cells with Au NPs added at 7% and 13%. These were then compared to rat primary muscle cells growing on PLLA scaffolds and TCP. The Au NPs were not cytotoxic to skeletal muscle cells as no significant difference was seen between the 7% Au NP treated cells, 13% Au NP treated



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7% Au–Plla 13% Au–PLLA 21% Au–PLLA Positive control

*

Absorbance

1.2

* *

1.0 0.8 * 0.6 0.4 0.2 0.0 Day 7

Day 14

Day 21

Day 28

7.4 Rat muscle cell proliferation on the 7% Au–PLLA scaffolds, 13% Au–PLLA scaffolds, 21% Au–PLLA scaffolds and the cells alone as a positive control (n = 5). The solutions were diluted (1:4) with deionised H2O (*p < 0.05). Reprinted with permission (McKeon-Fischer and Freeman, 2010).

cells and the positive control (Fig. 7.5) (McKeon-Fischer and Freeman, 2010). Tensile mechanical tests were performed on the three Au–PLLA scaffolds and PLLA scaffolds on days 0, 7, 14, 21 and 28. The elastic modulus and yield stress for each of the time points are shown in Fig. 7.6. The 7% Au–PLLA elastic modulus was significantly lower than the 21% Au–PLLA scaffolds on day 0. Both the 7% Au–PLLA and PLLA scaffolds were also significantly lower than the 21% Au–PLLA scaffolds for elastic modulus on day 7. Yield stress for 13% Au–PLLA was significantly smaller than 21% Au–PLLA on day 7. The 7% Au–PLLA and 13% Au–PLLA scaffolds were smaller than 21% Au–PLLA on day 21. No other significant differences were determined for elastic modulus and yield stress. However, the mechanical properties of both the PLLA and Au–PLLA scaffolds are too high to be used for skeletal muscle tissue engineering (McKeon-Fischer and Freeman, 2010). Therefore, a different polymer is necessary for skeletal muscle tissue engineering.

7.6.3 Carbon nanotubes Carbon nanotubes are graphite sheets formed into hollow cylinders with diameters ranging from a few nanometres to several micrometres (Edwards



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13% Au

PLLA

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Positive control

2.0 1.8

Absorbance at 490 nm

1.6 1.4 1.2 1.0 *

*

0.8 0.6 0.4

*

*

0.2 Day 7

Day 14

Day 21

Day 28

7.5 Rat muscle cell proliferation with 7% Au added to each well, 13% Au added to each well, PLLA scaffolds and TCP as a positive control. Each solution was diluted (1:4) with deionised H2O (n = 5) (*p < 0.05). Reprinted with permission (McKeon-Fischer and Freeman, 2010).

et al., 2009; Saeed and Park, 2007). These cylinders can then be fashioned into multiwalled carbon nanotubes (MWCNTs) by putting together 2–50 cylinders separated by 0.34 nm (Edwards et al., 2009; Fraczek et al., 2008). MWCNTs are being used in tissue engineering as they have high electrical conductivity, high thermal conductivity and are chemically stable (Edwards et al., 2009; Saeed and Park, 2007; Fraczek et al., 2008). They have been successfully combined with several different polymers including PLA, PU, poly(ethylene oxide) (PEO), PCL and PTFE (Mei et al., 2007; Kimmer et al., 2009; McCullen et al., 2007; Saeed and Park, 2007; Yeh et al., 2009; Saeed et al., 2006; Show and Itabashi, 2008). Saeed et al. (2006) electrospun PCL with several weight percents (wt.%) of both acid functionalised and unfunctionalised MWCNTs (Saeed et al., 2006). They found that acid functionalised MWCNTs dispersed better as they were embedded and oriented within the PCL nanofibre (Fig. 7.7) (Saeed et al., 2006). Saeed and co-workers (Saeed et al., 2006; Saeed and Park, 2007) successfully electrospun PCL with MWCNTs ranging from 0.5 to 7 wt.%. Although the toxicity of the MWCNTs is a concern, Fraczek et al. (2008) found that MWCNTs implanted directly into the gluteal muscle actually encouraged muscle regeneration and only a low inflammatory response was observed.



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Electrospinning for tissue regeneration 130 120 110

PLLA 7% Au–Plla 13% Au–PLLA 21% Au–PLLA

Elastic modulus (MPa)

100 90 80 70 60 50 40

* * *

30 20 Day 0

1800 1600

Day 7

Day 14 (a)

Day 21

Day 28

PLLA 7% Au–Plla 13% Au–PLLA 21% Au–PLLA

Yield stress (kPa)

1400 1200 1000 800

*

600

*

*

400 200 Day 0

Day 7

Day 14 (b)

Day 21

Day 28

7.6 Graphical analysis of the averages of (a) elastic modulus and (b) yield stress for PLLA scaffolds, 7% Au–PLLA scaffolds, 13% Au-PLLA scaffolds and 21% Au–PLLA scaffolds (n = 4) on days 0, 7, 14, 21 and 28. (*p < 0.05). Reprinted with permission (McKeon-Fischer and Freeman, 2010).



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MWCNT

100 nm

7.7 Transmission electron microscopy (TEM) image of PCL fibre containing a functionalised MWCNT (Saeed et al., 2006). Reprinted from Saeed et al. (2006) ©2006.

Just as conductive polymers are discovering a place in skeletal muscle tissue engineering so are carbon nanotube imbedded polymers owing to their increased conductivity. McKeon-Fischer and co-workers have successfully electrospun an inner layer of PCL–0.05% MWCNTs coaxially with an outer layer containing polyacrylic acid (PAA) and polyvinyl alcohol (PVA) hydrogel; data is unpublished. Figure 7.8 displays the PCL–0.05% MWCNTs–PAA/ PVA scaffold surface and the cross-section of several fibres illustrating the inner and outer polymer layers. The calculated electrical conductivities for PCL alone, PCL–0.05% MWCNTs and PCL–0.05% MWCNTs–PAA/PVA electrospun scaffolds were as follows: 0.026 (± 0.004), 0.043 (± 0.007) and 0.039 (± 0.011) S cm–1. A cell study with rat primary muscle cells was conducted on PCL, PCL–0.05% MWCNTs scaffolds and PCL–0.05% MWCNTs–PAA/PVA scaffolds. On days 7 and 15, PCL, PCL–0.05% MWCNT and PCL–0.05% MWCNT–PAA/PVA scaffolds displayed similar cellular activity. However on day 21, the PCL–0.05% MWCNT–PAA/PVA scaffolds showed a considerable increase. The PCL scaffolds also showed a large increase on day 28. PCL–0.05% MWCNT scaffold cellular activity was low and remained fairly constant over the four-week period. TCP, as a positive control, was significantly higher for all three of the scaffolds at all four time points (Tukey’s test, p < 0.05). PCL–0.05% MWCNTs–PAA/PVA scaffold cellular



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20 µm (a)

5 µm (b)

7.8 PCL-0.05% MWCNTs–PAA/PVA scaffolds surface at 5000x (a) and the cross-section of several fibres illustrating the outer and inner layers at 20 000x (b). Reprinted with permission.

activity was statistically greater for PCL and PCL–MWCNTs scaffolds on day 21. Both PCL and PCL–0.05% MWCNTs–PAA/PVA scaffolds were significantly larger compared to PCL–0.05% MWCNTs scaffolds on day 28 (Fig. 7.9). Therefore, no adverse effects seen from the MWCNTs. The average elastic modulus (n = 5) and yield stress (n = 5) for PCL alone, PCL–0.05% MWCNTs and PCL–0.05% MWCNTs–PAA/ PVA scaffolds were also measured. PCL alone had an average elastic modulus of 443.78 (± 176.27) kPa, 7148.02 (± 1169.98) kPa for PCL–0.05% MWCNTs and 12 483.72 (± 725.34) kPa for PCL–0.05% MWCNTs–PAA/PVA. The average yield stress for PCL was 60.80 (± 15.12) kPa, 579.00 (± 114.51) kPa for PCL–0.05% MWCNTs and 644.80 (± 87.20) kPa for PCL–0.05% MWCNTs–PAA/PVA (Fig. 7.10). PCL scaffolds compared to the whole rat muscle and pig muscle strips were not statistically significant for elastic moduli. Both the PCL–MWCNTs and PCL–MWCNTs–PAA/PVA scaffolds had significantly larger elastic moduli and yield stresses than PCL scaffolds, rat whole muscle and pig muscle. Although PCL–MWCNTs–PAA/PVA values are higher than any of the skeletal muscle elastic moduli and yield stresses, the amount of MWCNTs incorporated into the scaffold can be optimised to match skeletal muscle tissue properties better.

7.7

Conclusion and future trends

Skeletal muscle makes up 40% of all muscle in the body and controls movement. Its fibres are composed of many myoblasts fused together and orientated in the same direction. This unidirectional configuration allows a large



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PCL–MWCNT

PCL–MWCNT-PAA/PVA

141

Control

1.6 a

a

a

1.4

a

Absorbance

1.2

c b

1.0

c

0.8 0.6 0.4 0.2 Day 7

Day 14

Day 21

Day 28

7.9 Average rat muscle cell activity on PCL scaffolds, PCL–0.05% MWCNTs scaffolds, PCL–0.05% MWCNTs–PAA/PVA scaffolds and TCP (control) (n = 4). Each solution was diluted 1:4 with deionised H2O. Similar letters (e.g. all the a’s) are not significantly different from one another (p < 0.05). a, b and c represent significant difference between the samples.

force to be generated during contraction and movement to occur. However, this process can be impeded by traumatic injuries to both the peripheral nerves and skeletal muscle. Although several muscle tissue treatments have been explored, none have been able to restore muscle function. Therefore, continuing research for synthetic muscle replacements is necessary. Polymers are an attractive material for skeletal muscle tissue engineering as their mechanical properties can be customised for skeletal muscle and they can degrade in the body leaving the newly grown tissue behind. Electrospun scaffolds initiate the alignment of muscle cells in a single direction and can encourage the necessary organisation of the cells to create a functional skeletal muscle. By including a conductive component in the scaffold, electrical stimulation and mechanical stimulation can be combined to facilitate skeletal muscle organisation. Although no polymeric scaffold has been completely successful in reproducing the characteristics necessary for skeletal muscle, we believe that continuing research into electrospun scaffolds with conductive elements is a promising new direction for skeletal muscle tissue engineering and that future work in this area is needed.



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Electrospinning for tissue regeneration 50% PCL 50% PCL–0.05% MWCNT 50% PCL–0.05% MWCNT (83/17) PAA/PVA 14000 12000

Elastic modulus (kPa)

10000 8000 6000 4000 2000 0

(a) 50% PCL 50% PCL–0.05% MWCNT 50% PCL–0.05% MWCNT (83/17) PAA/PVA

700

Yield stress (kPa)

600 500 400 300 200 100 0

(b)

7.10 Elastic modulus (a) and yield stress (b) for the 50% PCL scaffolds, 50% PCL–0.05% MWCNT scaffolds and the 50% PCL– 0.05% MWCNTs–PAA/PVA scaffolds (n = 5 for each). Reprinted with permission.



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L. A. B o s w o r t h, The University of Manchester, UK

Abstract: Tendons are highly fibrous tissues that are capable of withstanding particularly high tensile loads. Unfortunately, tendons can degenerate as a consequence of repetitive loading, which may lead to their eventual rupture. Current clinical interventions commonly result in a healed tendon that is both biochemically and biomechanically inferior and highly susceptible to re-rupture. An unmet clinical need remains and many researchers are seeking to address this with the use of degradable biomaterials to create tissue engineered tendons. This chapter focuses on how the electrospinning technique can create scaffolds which closely resemble the natural tendon structure. Key words: alignment, biopolymers, electrospinning, scaffolds, tendon, tissue engineering.

8.1

Introduction: tendon tissue

Tendons are highly fibrous connective tissues that are both flexible and elastic. They allow forces generated by muscle contraction to be transmitted to the attached bone, allowing movement (Kannus, 2000). Being able to absorb external forces, tendons are able to act as a buffer, helping to prevent injury to the attached muscle (Sharma and Maffulli, 2005). For every muscle there are two tendons, distal and proximal, and each tendon has two points of union: where the tendon attaches to muscle is referred to as the myotendinous junction (MTJ) and where it connects to bone, it is the osteotendinous junction (OTJ) (Sharma and Maffulli, 2005). Depending on their location and function within the body, tendons have a variety of anatomical structures, their fibrous ultrastructures exhibiting rounded cords, flattened ribbons, strap-like bands or fan shapes (Kannus, 2000). The length of the tendon also varies. Tendons which connect the quadriceps to the patella are short and wide, making them very strong; whereas tendons in the fingers, which perform more subtle movements, are long and thin (Kannus, 2000). The points of union and their attachment to bone also vary between muscles.

8.2

Tendon structure and composition

Like other connective tissues, tendons are mainly composed of extracellular matrix (ECM), which accounts for 80% of the total tissue volume, with cells 148 © Woodhead Publishing Limited, 2011


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forming the remaining 20% (Nordin et al., 2001). Tendon ECM undergoes continual formation and degradation, which is necessary for its growth and healing (Chuen et al., 2004). The main component of the ECM is collagen, whose fibres primarily run along the length of the tendon axis with few transversely orientated fibres (Sharma and Maffulli, 2005). This highly aligned structure is essential for conferring the necessary tensile strength along the tendon’s longitudinal (and main) axis. Another important structural characteristic of tendons are their hierarchical arrangement, where grouping collagen molecules together forms collagen fibrils, which group further to form collagen fibres, followed by subfascicle, fascicle and tertiary fibre bundles layers, before reaching the overall tendon unit. The cross-sectional size of each level varies from tendon to tendon, depending on its location and purpose, and also between individuals. The average diameter for each hierarchical level is shown in Table 8.1. The hierarchical configuration allows the tendon to withstand high loads whilst maintaining its flexibility and enables tendons to bend round bony prominences depending on their anatomical location (Screen et al., 2004). Subfascicle and larger bundles are encompassed by the endotenon, which is a thin, loose connective tissue layer, which lies within the tendon structure and binds the collagen fibre bundles together (Wang, 2006). The entire tendon is further covered by a thin sheet of connective tissue, the epitenon, which is contiguous with both the endotenon and the outer layer, the paratenon, which is composed of loose areolar connective tissue (Kannus, 2000). For sheath-free tendons, such as the Achilles, the paratenon is the very outer layer of the tissue and it functions to reduce friction between the tendon and neighbouring tissue, thus ensuring unhindered movement and preventing wearing (Sharma and Maffulli, 2005). Certain tendons, such as those in the hands and feet, which undergo increased levels of mechanical stress, are further enclosed within a membranous sheath. For these tendons, a thin film of fluid is present within the paratenon and sheath, which lubricates the tissues and facilitates free movement when the tendons are placed under high stress. Table 8.1 Hierarchical layers found within natural tendon tissue and their corresponding diameters Tendon hierarchical layers

Diameter range

Collagen molecules Collagen fibrils Collagen fibres Primary fibre bundle (subfascicle) Secondary fibre bundle (fascicle) Tertiary fibre bundle Tendon unit

1.5 nm 3.5 nm 10–20 nm 50–500 nm 10–50 µm 50–400 µm 100–500 µm

Source: Screen et al., 2004



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8.2.1 Tendon cells As mentioned earlier, the cellular content of tendons accounts for 20% of the overall tissue volume (Nordin et al., 2001). The main cell type present within tendons, which is responsible for the production and regulation of ECM, is the tenocyte (Maffulli et al., 2002). These are bipolar cells with elongated nuclei and spindle-shaped morphologies; they are regularly anchored in columns amongst the collagen bundles (Bernard-Beaubois et al., 1997). Other cells (though present in smaller numbers) include chondrocytes present at OTJ insertion sites, synovial cells composing the tendon sheath and vascular cells such as capillary endothelial and arteriole smooth muscle cells (Sharma and Maffulli, 2005).

8.2.2 Extra cellular matrix (ECM) The ECM is composed of ground substance (proteoglycans and glycoproteins) and fibrous proteins (elastin and collagen) (Sharma and Maffulli, 2005). Proteoglycans comprise a protein core and at least one covalently bound polysaccharide chain or glycosaminoglycan (GAG) (Kjaer, 2004). Proteoglycans enable the diffusion of water soluble molecules and migration of cells and are located amongst collagen fibrils and fibres (Yoon and Halper, 2005). Glycoproteins are macromolecules that contain a high number of protein and small glycan chains. Fibronectin is a glycoprotein present in tendons; it is responsible for regulating deposition and alignment of collagen molecules within the ECM and also controls cell binding to the ECM via fibronectin receptors found on cell membranes (Józsa and Kannus, 1997). These receptors are able to bind with actin filaments present within the cells’ cytoskeleton and the collagen present within the ECM enabling cell–matrix interactions (Burkitt et al., 1993). Several types of collagen make up 70% of tendon dry mass, with collagen type I fibres being the main structural component. However, collagen types II, III, IV and V are also present although in far smaller quantities. The parallel alignment of collagen fibres along the tendons’ longitudinal axis and covalent crosslinks between these fibres provide tendons with great tensile strength (Maffulli et al., 2000), for example the human Achilles tendon is capable of reaching strengths of 36.5 MPa when tested in vivo (although tendons were not measured to rupture) (Magnusson et al., 2003). However, the arrangement of collagen type I varies from loose to densely packed, depending upon the mechanical loads it is required to endure (Burkitt et al., 1993). In their relaxed state, tendons adopt a crimped configuration, although, on application of a load this wave formation disappears as the fibres become elongated and stretched. Elastin fibres, which account for only 2% of the



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tendon’s dry mass, are thought to be responsible for the recovery of the tissue to its natural crimped formation (Kannus, 2000).

8.3

Tendon pathology

All tendons have the potential to be affected by direct damage caused by accidental lacerations, or through participation in sports. Furthermore they can be susceptible to diseases, such as osteoarthritis. Tendon diseases were previously assumed to be caused by inflammation of the tissue and were termed ‘tendonitis’. However, Puddu et al. (1976) reasoned that the avascularity of tendons does not predispose the tissue to inflammatory processes and thus tendonitis is a misconception. Clinically, tendon disorders are commonly referred to as a ‘tendinopathy’ as this makes no assumption about the pathological processes within the tendon (Rees et al., 2006). Tendinopathies within the Achilles tendon which cause degeneration of the tissue are often the result of excessive and repetitive overloading of the Achilles tendon within both sporting and sedentary patients (Kader et al., 2002). In a study of 430 cases, 62% of Achilles ruptures were sports related (Kannus and Natri, 1997). Other tendons prone to pathology include the rotator cuff in the shoulder, where degeneration and tear size increases with age (Hijioka et al., 1993; Sher et al., 1995). This was recently demonstrated in a study by Yamamoto et al. (2010), where full thickness rotator cuff tears occurred in 20.7% of 1366 shoulders and prevalence increased with age. The patella tendon in the knee is commonly injured in athletes owing to excessive load bearing as a consequence of long-term training, which causes tenocytes to undergo apoptosis (or programmed cell death) (Lian et al., 2007). A variety of treatments are employed for tendinopathy management. In the early phase of the disease, conservative methods are customary and those who do not respond well to these treatments after six months often undergo surgical intervention (Maffulli et al., 2004). Conservative methods include the use of non-steroidal anti-inflammatory drugs (NSAIDs) and corticosteroids, although their use for tendon healing remains controversial owing to their conflicting benefits within the literature (Maffulli et al., 2004; Rees et al., 2006). Other conservative methods include cryotherapy and ultrasound, although there is little clinical evidence to prove their efficacy (Rees et al., 2006; Robertson and Baker, 2001). For patients presenting with acute rupture of the Achilles tendon, management depends on the surgeon’s own preference, which can fall within three main categories: open operative, percutaneous operative and non-operative (Maffulli, 1999). Open operative surgery involves the repair of the two ruptured ends by suturing them together. Percutaneous operative is a combination of open and non-operative techniques that was developed



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by Ma and Griffith (1977). It involves a number of small incisions used to suture the tendon without fully exposing the tissue. Non-operative treatments involve the immobilisation of the lower leg in a plaster cast for a period of 6–8 weeks (Cetti et al., 1993). Comparison between the results of these categories is limited owing to the lack of controlled, randomised trials and a reliable scoring system for the outcome following treated Achilles tendons (Maffulli, 1999). Owing to the often poor response to treatment and the resultant morbidity of tendon disease, there is a growing interest in novel techniques to repair such tissues. Following injury, tendon heals by production of scar tissue, which is organisationally, biochemically and biomechanically inferior to normal tendon matrix tissue (Corps et al., 2006; Leadbetter, 1992). Such inferior scar tissue leads to ongoing morbidity of affected patients (Kader et al., 2002). Previous strategies employed to improve the quality of tendon repair after injury have included xenograft tendons cross-linked with glutaraldehyde, as demonstrated by Smith et al. (1986). Tissue engineering approaches have utilised autologous tenocytes in biomaterial scaffolds (Cao et al., 2002). Collagen-based scaffolds have been investigated in an attempt to match the mechanical properties to native tendon (Garvin et al., 2003; Feng et al., 2006). Synthetic bioresorbable polymers such as poly(lactic-co-glycolic acid) (PLGA) and poly(glycolic acid) (PGA) have been developed to create scaffolds for potential use in tendon applications (Sahoo et al., 2006; Liu et al., 2006).

8.4

Clinical need

As mentioned, all tendons have the potential to be affected by direct damage or disease. The most commonly affected tendon is the Achilles, with ruptures occurring in approximately 18 per 100 000 patients each year and this figure is believed to be increasing (Leppilahti and Orava, 1998; Maffulli et al., 1999). Other tendons, such as flexor and extensor tendons of the hand have demonstrated a combined incidence level of 22.7 per 100 000 over a five-year period (Clayton and Court-Brown, 2008). The repair of tendons is a procedure with varying levels of complexity depending on the location of the injury, pathology, age and lifestyle of the patient, surgical expertise and preference of intervention. Current interventions for damaged and diseased tendons include prosthetic devices, allografts, autografts and operative or non-operative interventions. However, levels of difficulty are associated with each of these therapies. Prosthetic devices commonly experience problems of longevity and poor performance and thus do not provide an adequate solution for patients (Goldstein et al., 1989; Sahoo et al., 2006). Allografts involve removing tendon tissue from another person, most likely a cadaver, and implanting



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this tissue into the area of tendon damage (Robertson et al., 2006). Problems associated with this approach include increased chance of infection and tissue rejection and longer times for tissue remodelling (McDermott and Thomas, 2005). A common approach for patients with tendon injuries involves removing healthy tendon tissue from another region of the patient’s body and implanting this into the damaged area (autografting). Tendons used for the repair of ruptured Achilles have included the peroneus brevis (Pintore et al., 2001), flexor hallucis longus (Wilcox et al., 2000) and flexor digitorum longus (Mann et al., 1991). Whilst autografts eliminate the risk of tissue rejection, they do create secondary sites of morbidity, which increases the chance of infection and prolongs recovery time. For the treatment of Achilles ruptures, a comprehensive review by Wong et al. (2002) compared operative and non-operative management of this tendon for 5056 Achilles ruptures. Re-rupture rates for conservative approaches were determined to be 9.8% compared to 2.2% for operative management. Their conclusions determined operative repair and early mobilisation to be the most common management technique employed for spontaneous ruptures in younger athletic patients. Discrepancies between preferred treatments is likely to be a result of uncontrollable factors, such as patient individuality and age, the nature of the rupture, variation in treatment protocols and surgeon proficiency. Whilst ruptured tendons can be managed operatively, their repair invariably reduces the biological and mechanical properties of the tendon as a consequence of scar tissue formation caused by immobilisation of the sutured tendon (Koob, 2002). Owing to poor clinical outcome and a rise in the incidence of tendon pathologies (Riley, 2008), clinicians have recognised the need for alternative repair therapies that are capable of temporarily supporting and sustaining the tendon tissue, promoting tenogenesis and reducing recovery time but which do not elicit an immune response.

8.5

Tissue engineering

The majority of tissue engineering applications, regardless of their final location in the body, require the fabrication of a construct, with dimensions and structure that can fill the damaged area, promote tissue ingrowth and perform the original tissue’s role. Constructs, more commonly referred to as scaffolds, aim to replicate the original tissue’s ultrastructure. Scaffolds fabricated by the electrospinning method are being increasingly used in tissue engineering applications owing to their ability to create nanoscopic fibres from as small as 20 nm (Zhou et al., 2003). Biopolymers, which include both naturally occurring (e.g. collagen and elastin) and synthetic materials (e.g. poly(e-caprolactone)), are



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being increasingly used in the tissue engineering field because of their biocompatibility and ability to degrade over time without causing adverse effects on the body. However, for naturally occurring materials there are several limitations, including material availability and variation between batches as the material is sourced from human or animal tissue (Kim and Mooney, 1998). By contrast, synthetic materials are manufactured in bulk and are readily obtainable. Whilst synthetic materials can be processed by a range of techniques, including electrospinning, they are, however, hindered by a lack of cell-recognition receptors on the material surface. Nevertheless, the material surface chemistry can be specifically altered by functionalising with appropriate biomolecules, such as fibronectin and Arg-Gly-Asp (RGD) peptides, which makes the material surface far more attractive for the adhesion of cells (Kim and Park, 2006).

8.5.1 Tissue engineering of tendons If the scaffold is to provide a temporary repair and stimulate tissue regeneration, the scaffold must provide a structural topography that promotes appropriate contact guidance for the cells and an environment that encourages formation of the organised tendon tissue. Scaffolds for tendon tissue engineering must achieve four main aims: 1. The scaffold should resemble the ultrastructure of tendon tissue as closely as possible: the scaffold should be highly fibrous with purposefully orientated fibres. Ideally the scaffold should aim to mimic the hierarchical configuration of tendons. 2. The scaffold should stimulate an appropriate tendon tissue response from migrated tendon cells, this being production of aligned collagen type I fibres as opposed to scar tissue formation. 3. The scaffold should be sufficiently load bearing: the scaffold must be able to withstand forces (specifically tensile) that are placed on it, otherwise the graft and tissue will be compromised and it is likely to rupture. 4. The scaffold should be made of degradable materials: as new tendon tissue forms the scaffold should be degrading without toxicity and without compromising the neo-tissue. Material selection – tendon repair Previous tendon tissue engineering approaches have included scaffolds fabricated from either synthetic or natural materials, with collagen being a popular choice for natural materials possibly because tendons are principally composed of collagen (Garvin et al., 2003; Juncosa-Melvin et al., 2006; Nirmalanandhan et al., 2006; Feng et al., 2006).



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Whilst other naturally derived scaffolds researched for tendon applications have included chitosan-based hyaluronan hybrid polymer fibres (Funakoshi et al., 2005) and alginate–chitosan polyion hybrid fibrous complex (Majima et al., 2005), a large focus has been targeted on synthetic degradable materials, such as PGA (Cao et al., 2002; Liu et al., 2006) and PLGA (Ouyang et al., 2003; Sahoo et al., 2006; Moffat et al., 2009) owing to their mechanical properties (Montes de Oca and Ward, 2006; Ma et al., 2007). Table 8.2 addresses the various research projects that have been on-going in tendon tissue engineering over recent years. Whilst this table of prior research is not exhaustive, it highlights the wide range of techniques and materials used in the process of developing a construct with desirable properties for tendon tissue engineering. Effect of nanofibres on cell behaviour A number of the research groups included in Table 8.2 have incorporated the electrospinning technique as their method of fabricating scaffolds for this particular tissue application. This is because the fine fibres that are produced provide seeded cells with an environment similar to their natural surrounding ECM, consequently supporting cell viability, proliferation and phenotype. The engineered scaffold must take a three-dimensional (3D) form in order to simulate analogous cell responses as in vivo. A nanofibrous scaffold should provide this desired 3D environment owing to the greater surface area to volume ratio available and hence making a larger area of receptor sites available to the cell, thus aiding maintenance of cellular phenotype and normal cell functions (Stevens and George, 2005). Our own research focuses on the use of electrospinning to create 3D fibrous constructs which mimic the tendon ultrastructure. Whilst other research groups fabricate aligned nanofibrous sheets, our research takes this approach one step further in creating 3D bundles composed of nanofibres. This replicates the tertiary fibre bundle of the natural tendon tissue mentioned in Table 8.1, whilst the individual nanofibres resemble the subfascicle layer. Our intention is to use this single bundle as our primary scaffold, which can be woven with any number of other bundles to create larger structures with a hierarchical element that could be implanted into any size of tendon defect (Fig. 8.1).

8.6

Cell response to electrospun bundles

Poly(e-caprolactone) (PCL) nanofibres are naturally hydrophobic and do not possess natural surface receptor sites for the adhesion of cells (Bosworth and Downes, 2009). In order to make the surface more hydrophilic and biocompatible it must undergo some form of surface treatment. This can involve soaking the material in concentrated sodium hydroxide solution



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Table 8.2 Overview of research groups seeking to develop tendon tissue engineered constructs using a variety of different materials, techniques and models for assessment Scaffold material Model system and structure

Research outcome

Reference

Unwoven PGA fibres seeded with tenocytes for one week prior to implantation

In vivo – chicken second flexor digitorum profundus tendon

14 weeks post-operation, tendons Cao et al., with seeded PGA constructs 2002 were practically indistinguishable to normal tendon with tensile strengths reaching 83% of normal tendon

Mechanically loaded 3D collagen gels

In vitro

Phenotypically similar to native tendon compared to unloaded gels

Knitted PLGA fibres seeded with bone marrow stromal cells

In vivo – rabbit At 8 and 12 weeks post-operation Ouyang Achilles tendon both pre-seeded and unseeded et al., 2003 fibrous scaffolds demonstrated histology similar to natural tendon tissue. Seeded constructs yielded greatest tensile strength and stiffness compared to unseeded

Garvin et al., 2003

Self-organisation In vitro of tendon fibroblasts

Viable in vitro engineered tendon Calve without artificial scaffolding; et al., 2004 stress–strain response comparable to embryonic chick tendons

Polydioxane In vivo – rat tube with flexor tendon microgrooves on the interior

After 6 weeks, purposefully crushed tendons had functionally returned to normal when held within the tube

Chitosan-based hyaluronan hybrid fibrous scaffold seeded with fibroblasts

Type I collagen observed in cellFunakoshi seeded scaffolds only; mechanical et al., 2005 properties of this scaffold increased significantly 12 weeks post-operative

In vivo – rabbit infraspinatus tendon

Curtis et al., 2005

Alginate– In vitro chitosan polyion hybrid fibrous complex

Addition of chitosan to the scaffold complex increased the level of tendon fibroblast adhesion compared to alginate fibres alone

Majima et al., 2005

PLGA nanofibres In vitro electrospun on PLGA knitted scaffold

Higher expression of collagen I, decorin and biglycan genes on nanofibrous/knitted scaffolds compared to knitted

Sahoo et al., 2006

PGA unwoven fibres with dermal fibroblasts

26 weeks post-operation, Liu et al., dermal fibroblasts demonstrated 2006 histology similar to that of natural tissue and resulted in similar tensile strengths when compared to tenocyte-seeded scaffolds

In vivo – pig flexor digital superficial tendon



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Table 8.2 Continued Scaffold material Model system and structure

Research outcome

Collagen sponge In vivo – rabbit constructs patellar tendon containing MSCs prestimulated for two weeks prior to implantation Fibroblast– In vitro collagen gel with static and dynamic loading

Mechanical stimulation of Juncosaconstructs yielded greater Melvin mechanical properties compared et al., 2006 to non-stimulated for both in vitro and in vivo assessment

PLGA electrospun aligned nanofibres

In vitro

PLLA electrospun In vitro/in aligned vivo – mouse nanofibres gastrocnemius

Dynamically loaded gels resulted in mechanically weaker constructs compared to static samples; dynamic loading pattern to be addressed Contact guidance controlled cell alignment, distribution and matrix deposition compared to unaligned fibres; mechanical properties of aligned fibres were higher Scaffold provided an instructive microenvironment for tendon stem/progenitor cells to differentiate into the teno-lineage

Reference

Feng et al., 2006

Moffat et al., 2009

Yin et al., 2010

to induce basic hydrolysis and production of —COOH and —OH groups (Sun and Downes, 2009), or soaking the material in cell culture medium containing serum, which relies on serum proteins being adsorbed onto the material surface (Mo et al., 2004). Our in vitro investigations involved soaking the fibres in serum-containing medium prior to cell seeding. This option was preferable to sodium hydroxide treatment as this method induces material degradation and weakened our scaffold structure. Tenocytes were subsequently cultured on 3D bundles and woven structures for a two-week period. During this time cells aligned preferentially in the same direction as the nanofibres, which would suggest the fibres were conferring an element of contact guidance to the cells (Fig. 8.2). For the woven structures, cells demonstrated similar alignment as for the single 3D bundles, but were also found to bridge in-between the artificially created pores that were created as a consequence of the plaiting technique (Fig. 8.3). Direct penetration of the cells into the scaffold structure was not assessed; however, for these particular scaffolds (i.e. 3D bundles of PCL (dissolved in acetone)) cellular infiltration was restricted to the very outer surface of the fibres. Yet tenocytes cultured on fibres spun from PCL dissolved in hexafluoroisopropanol (HFIP) appeared to demonstrate some infiltration into the scaffold structure after just 24 hours in culture (Fig. 8.4).



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Step 1  Polymer solution electrospun under controlled parameters  Fibres deposited on suitable collector plate

Step 3  Sheets of aligned fibres are twisted together to create 3D bundles

20 mm

Step 2  Deposited fibres removed from collector plate as single sheet

Step 4  3D bundles can be scaled-up into different ‘rope-like’ formations

100 mm 5 mm

333 mm

8.1 Schematic demonstrating the electrospinning process and fibre manipulation to produce 3D fibrous bundles and larger woven structures.

8.7

Mechanical properties of electrospun bundles

The orientation and overall structure of an electrospun fibre scaffold will affect its mechanical properties. Significant changes in strength were observed depending on the fibre orientation, that is random or aligned fibres. Tensile properties were lowest for randomly orientated fibres; however, tensile strength and stiffness were significantly increased when applying a load in the direction of the fibres’ longitudinal axis. This is an observation similarly obtained by other research groups, including Yin et al. (2010) who tensile tested PLLA random and aligned electrospun fibres, with Young’s modulus demonstrating a 36-fold increase when moving from random to aligned orientation and Moffat et al. (2009) who similarly tested PLGA electrospun fibres and found the Young’s modulus to be three times as strong for aligned fibres compared to random ones. We also assessed the tensile properties of the electrospun 3D bundles, which involved loading along the bundle’s longitudinal axis. Both the Young’s modulus and ultimate tensile strength (UTS) of the bundles were significantly improved compared to the aligned electrospun mat (Table 8.3). However, as tendons are capable of withstanding sufficiently high tensile loads, for example, the tensile strength of human Achilles’ tendon is 36.5 MPa (not tested to rupture) (Magnusson et al., 2003), it is vitally important that the scaffold should be able to withstand similar strengths. Consequently, we assessed the tensile properties of electrospun bundles fabricated from other



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5.0 µm

50.0 µm (a)

50.0 µm

5.0 µm (b)

50.0 µm

5.0 µm (c)

8.2 Scanning electron microscope (SEM) micrographs demonstrating morphology and orientation of tenocytes cultured on 3D electrospun bundles after (a) 1 day, (b) 5 days and (c) 14 days. Cells and matrix appear to orientate along the main fibre axis (magnifications x200 (left) and x2000 (right)).

biopolymers and spinning solutions with different solvents. SEM images and fibre diameter distributions are indicated in Fig. 8.5; the corresponding tensile properties for these electrospun bundles are shown in Fig. 8.6. The results demonstrate the tensile properties of the electrospun bundles could be easily altered depending on the materials and solvent used. For example, PCL bundles made from an acetone-containing solution resulted in the weakest overall tensile properties. However, if the polymer (and solvent) was altered, in this case to PLGA(85:15) (dissolved in HFIP), there was a significant increase in Young’s modulus (up 11-fold) but no real change in the UTS. Similarly, if the PCL solvent was changed from acetone to HFIP, there was a considerable improvement in Young’s modulus and UTS, with mean values increasing five-fold and ten-fold, respectively.



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100 µm

10.0 µm (a)

100 µm

10.0 µm (b)

100 µm

10.0 µm (c)

8.3 SEM micrographs demonstrating morphology and orientation of tenocytes cultured on 3D electrospun bundles that have been plaited together after (a) 1 day, (b) 5 days and (c) 14 days. Cells and matrix appear to bridge the artificial pores created by the interweaving of the bundles (magnifications x200 (left) and x2000 (right)).

8.8


This chapter has highlighted the need for the development of a biomaterial/ tissue engineered device for patients with tendon injuries. Electrospinning appears to be a suitable technique for creating scaffolds which resemble the tendon ultrastructure, particularly 3D scaffolds which mimic the tendon’s collagen fibrous bundles and hierarchical arrangement. With surface treatments of synthetic materials, such as PCL, the electrospun fibres readily support cell attachment and proliferation and even confer contact guidance to the cells, thus directing cellular orientation. The selection of material and solvent warrants significant consideration as the tensile properties are significantly affected and are of considerable importance in this particular tissue application.



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2 µm

2 µm

8.4 SEM micrographs of poly(e-caprolactone) nanofibres, electrospun from a solution of hexafluoroisopropanol (HFIP) (concentration 10%w/v), with tenocytes cultured for 24 hours demonstrating alignment and infiltration of the cells into the scaffold structure (represented by arrowheads) (magnification x2000). Table 8.3 Tensile properties of different poly(ecaprolactone) electrospun fibrous scaffolds; represented as mean (± standard deviation) Scaffold structure

Modulus (MPa)

Tensile strength (MPa)

Random fibre mat

1.54 (± 0.26) 4.84 (± 0.13) 14.11 (± 3.76)

0.45 (± 0.09) 1.30 (± 0.14) 4.74 (± 1.64)

Aligned fibre mat 3D fibre bundle



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Fibre diameter (µm)


10 9 8 7 6 5 4 3 2 1 0

Number of fibres


(c)

Fibre ø 0.80±0.30

0.0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4 1.5

5.0 µm

20.0 µm


(d)

Fibre ø 0.56±0.13

0.0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4 1.5

16 14 12 10 8 6 4 2 0

5.0 µm

20.0 µm

8.5 SEM micrographs, fibre diameter distributions and mean fibre diameter (± standard deviation) for (a) poly(ecaprolactone) (PCL) (dissolved in acetone), (b) PCL (dissolved in hexafluoroisopropanol (HFIP)), (c) poly(lactic-co-glycolic) acid (PLGA85:15) (dissolved in HFIP) and (d) PCL–PLGA85:15 (50:50 blend dissolved in HFIP) (magnifications x50 (a and b top), x100 (c and d top) and x2000 bottom).

(b)

0.0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4 1.5

(a)

0

5

10

15

20

25

5.0 µm

50.0 µm

Fibre ø 0.44±0.11

Number of fibres

Fibre ø 0.33±0.18

0.0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4 1.5

0

5

10

15

20

25

30

5.0 µm

50.0 µm

Number of fibres

Number of fibres


163

In order to create fully tissue engineered devices for tendon implantation, incorporation of dynamic-loading bioreactors should be utilised. This will enable appropriate loading patterns to be applied to the cell-seeded scaffolds, which should elicit a more appropriate cell response, in terms of gene/protein expression, when compared to static cell culture. Future directions should begin to focus on the development of full tendon replacements, with consideration of the bone and muscle interfaces (OTJ and MTJ) and how these devices would be grafted and secured in place. 200

Modulus (MPa)

150

100

50

0

PCL (acetone) PCL (HFIP)

PLGA

80

Tensile strength (MPa)

PCL–PLGA

(a)

60

40

20

0

PCL (acetone)

PCL (HFIP)

PLGA

PCL–PLGA

(b)

8.6 Tensile testing data for 3D electrospun bundles of poly(ecaprolactone) (PCL) (dissolved in acetone), PCL (dissolved in hexafluoroisopropanol (HFIP)), poly(lactic-co-glycolic) acid (PLGA85:15) (dissolved in HFIP) and PCL–PLGA85:15 (50:50 blend dissolved in HFIP) in terms of (a) Young’s modulus, (b) ultimate tensile strength and (c) strain. One-way ANOVA with Bonferroni post-tests; n = 5. *p < 0.05 and # not significant.



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Electrospinning for tissue regeneration 4

Strain (mm mm–1)

3

2

1

0

PCL (acetone)

PCL (HFIP) (c)

PLGA

PCL–PLGA

8.6 Continued

8.9

Acknowledgements

The author wishes to thank RegeNer8 (the translational centre for regenerative medicine) for funding this research.

8.10

References

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9

Nerve tissue regeneration

C. W a n g, H. K o h and S. R a m a kr i sh n a, National University of Singapore, Singapore and S. L i a o, Nanyang Technological University, Singapore

Abstract: Injuries occurring in the peripheral and central nervous systems are healed by slow or limited endogenous repair. To accelerate this process, microsurgeons have used autografts and allografts to treat the injuries. However, donor morbidity induced by the autograft, limited donor availability and potential risk of immunorejection of the allograft have encouraged the development of potential alternatives such as bioengineered nerve grafts. In this chapter, regeneration of nerve tissue by electrospun nanofibrous grafts will be discussed as a means of solving some of the clinical problems encountered, such as the limited availability of allogeneic tissue that can be used in nerve conduits to achieve optimum mechanical properties and the size necessary for joining and regeneration. By mimicking native nerve tissue, nerve tissue engineering has been developed based on scaffold, cells and growth factors. Furthermore, we will discuss the nanofibrous scaffolds for use in the peripheral nervous system and their effect on neuronal differentiation of stem cells. A nanofibrous conduit and intraluminal guidance could be easily constructed by electrospinning. There is a great possibility that the superior adhesion and proliferation of stem cells, including nerve stem cells, on a nanofibrous scaffold will be able to facilitate stem cell based therapy clinically. Key words: bundle, conduit, electrospinning, intraluminal guidance, nanofibre, neurite, tissue regeneration.

9.1

Introduction

In the United States, there are over 250 000 people living with spinal cord injury (SCI) and approximately 11 000 new spinal cord injuries are registered each year, in addition to 1.5 million new cases of traumatic brain injury. Children and young adults are disproportionately affected by both types of injury, a grim statistic that is compounded by the lifelong disability that usually results. On the other hand, in 2002 more than 250 000 US patients suffered major traumatic wounds to peripheral nerves, including injuries from collisions, motor vehicle accidents, gun wounds, fractures, dislocations, lacerations or some other form of penetrating trauma. Owing to the difficulty and low efficacy of treating major peripheral nerve injuries, only 15% of patients were actually treated for their peripheral nerve problem (Atala, 2008; Yannas et al., 2007). 168 © Woodhead Publishing Limited, 2011


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Nerve tissue regeneration has been intensively investigated. At present, nerve autografts are still considered to be the ‘golden standard’ clinically in nerve tissue repairs for gaps of 20 mm or longer, and where direct suturing of the ends of two nerves will form tension at the suture line (Yannas et al., 2007). However, a donor nerve, from functionally less important regions, needs to be compensated for such a treatment, not to mention that it comes with many other implications like donor site morbidity, loss of function and/or formation of potentially painful neuromas at the donor site (Panseri et al., 2008; Jiang et al., 2010). In addition, donor nerves are limited and there may not be enough tissue for a large defect (Johnson and Soucacos, 2008). Thus, allografts are used as alternatives to autografts (Mackinnon and Hudson, 1992). However, the underlying problem remains unsolved as allografts are plagued by the risk of immunorejection and patients will have to rely on immunosuppression drugs until recovery or for life. Thus, many researchers have been looking into artificial nerve conduits to cope with the massive demand worldwide. In the following sections, we will give a detailed introduction to current clinical problems in nerve regeneration, discussions on the development of nerve tissue engineering and how nanofibrous scaffolds and stem cells (e.g. nerve stem cell (NSC), mesenchymal stem cell (MSC) and embryonic stem cell (ESC) interact with each other.

9.2

Clinical problems in nerve tissue therapy

9.2.1 Central nervous system Traumatic injuries to the central nervous system (CNS) come in a multitude of forms, most of which lead to long-term or permanent reduction in the quality of life. Traumatic brain injury and traumatic SCI are particularly devastating, both emotionally and financially. The complexity of the CNS makes efforts to repair injuries very difficult. Similarly Parkinson’s disease (PD), characterised by a progressive deterioration and loss of nigrostriatal dopaminergic neurons in the substantia nigra, presents clinical symptoms such as tremor, rigidity and bradykinesia. Dopaminergic drugs and functional neurosurgery largely reverse the classic motor features of Parkinson’s disease, tremor, rigidity and akinesia. As a result, most of the disability caused by Parkinson’s disease relates to symptoms such as gait dysfunction, loss of balance, swallowing and speech difficulties, autonomic disturbances and cognitive decline, which are less influenced by available therapies. Dementia is considered to be the major long-term cause of disability in people with Parkinson’s disease (Obeso et al., 2010) and arterial occlusion within the brain can lead to ischaemia and infarction of the brain parenchyma. The current treatment of stroke remains limited and focuses on neuroprotection



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to limit the expansion of the infarct and possibly recover the cells within the ischaemic penumbra. SCI remains a devastating ailment with little opportunity for treatment. The injury occurs from mechanical forces in an acute setting and is exacerbated by secondary inflammatory damage, both leading to neuronal death and demyelination. Accordingly, potential therapy would vary depending on the timeframe after injury, with minimising inflammation the primary concern early after injury and regeneration the major goal when injury is in its chronic phase (Jandial et al., 2007). After the 1990s, the stem cell was proposed for treatment of CNS diseases (Table 9.1). Furthermore, progress in tissue engineering in recent years has led to new approaches for the treatment of CNS injuries and renewed hope. We will talk about nerve tissue engineering in Section 9.3.

9.2.2 Peripheral nerve system Peripheral nerve injuries are common, with approximately 200 000 patients treated each year in the USA (Yannas et al., 2007). These injuries come in many forms, such as accidental trauma, external compression, disease, physical injuries or from certain surgical procedures requiring sectioning of nerve to gain access to adjacent surgical sites. Failure to restore these damaged nerves can lead to the loss of muscle function, impaired sensation and painful neuropathies. Peripheral nerve injury management has improved greatly in the past decades. Knowledge gained in basic science research is becoming increasingly important as it helps refine nerve surgical procedures, offering Table 9.1 Stem cell treatment for central nervous system (CNS) diseases Disease Graft

Animal model

PD

Rodent; Yes, challenges before monkey clinical translation; adverse side effects such as dyskinesias

(Soderstrom et al., 2006; Takagi et al., 2005)

Rodent/ Feasible murine models

(Park et al., 2002; Nakatomi et al., 2002)

Murine models

(KarimiAbdolrezaee et al., 2006; Cummings et al., 2005)

Stroke

SCI

Dopamine producing cells; fetal mesencephalic grafts; dopaminergic neurons derived from both ESC, NSC Combined use of stem cells together with polymer scaffolds; human NT-2 teratocarcinoma line; ESCs and human somatic NSCs NSCs or their derivatives (e.g. oligodendrocytes for myelination); human CNS fetal-derived stem cells; murine and human NSCs.

Clinical availability

More controlled differentiation in order both to maximise functional recovery and minimise side effects

Reference



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the promise of near-anatomic restoration of nerve structure and function. At the same time, a good understanding of the intraneural topography and anatomy of the peripheral nerve is also required to ensure better chances of functional recovery. Autologous nerve transplantation is the gold standard for repairing peripheral nerve injuries that result in gaps that cannot be repaired by simply suturing to reconnect the proximal and distal stumps. The sural nerve is the most common nerve graft used because this nerve is easy to harvest, providing approximately 25–30 cm of nerve tissue and leaving only minor sensory deficits (Schlosshauer et al., 2006). However, clinical reconstruction usually encompasses repairing the mixed or motor nerve deficits with a sensory nerve graft that might present inappropriate nerve alignment and size disparity (Nichols et al., 2004). Other biological bridging conduits, such as vein grafts, present limitations where the regenerated fibres penetrate through the vessel wall and entangle with the scar tissue together with collapse of the veins owing to their thin walls (Meek and Coert, 2002). A bioengineered nerve construct is an alternative to bridge nerve deficits to achieve functional recovery of the damaged nerve. These non-biological nerve guide implants can replace autologous nerve graft transplantations which are frequently limited by tissue availability, secondary deformities and the second surgical step needed to obtain the donor nerves. Although FDAapproved tissue engineered nerve devices (Table 9.2) are already available, these devices are reserved for small defects (from several millimetres up to 30 mm) and do not address the larger peripheral nerve gap injuries commonly found in clinics (Schmidt and Leach, 2003; Katayama et al., 2006). Tissue engineered constructs that facilitate axonal growth and glial cell proliferation, which include a scaffold, supporting cells, growth factors and ECM molecules, have been previously studied. However, results for these constructs have failed to equal the nerve regeneration achieved with autologous nerve grafts (Schmidt and Leach, 2003). Recent advances and contributions in the field of tissue engineering, in particular advances in the preparation of tissue scaffolds and better understanding of the biology of nerve regeneration, have focused research on the combination of materials and desired biomolecules that will create new composite materials that can actively stimulate the rate of nerve regeneration, increase the gap distance that the nerve construct can successfully bridge, target motor and sensory nerves reinnervation to their respective organs and minimise inflammation.

9.3

Nerve tissue engineering

9.3.1 Native nerve tissue Peripheral nerves consist of sensory and/or motor axons that are extensions of the nerve cell bodies. These nerve cell bodies are located in the dorsal root © Woodhead Publishing Limited, 2011


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Table 9.2 List of commercially available artificial nerve grafts Nerve guide (company)

Material

NeuraGen® (Integra)

Collagen type I

NeuroTube® (Synovis Micro)

Poly(glycolic acid)

Salubridge™

Polyvinyl alcohol hydrogel

Avance™ (Axogen Inc.)

Acellular nerve graft

Nerve guide image

Neuroflex™ (Collagen Matrix Inc.) Collagen type I

Neurolac® (Polyganics)

Poly(e-caprolactone)-co(dl-lactic acid) (50:50)

Source: Collagen Matrix, 2005; Schlosshauer et al., 2006; SaluMedical, 2000

ganglia (sensory neurons), autonomic ganglia (autonomic neurons) and the ventral horn of the spinal cord or brain stem (motor neurons) of the nervous system. In general terms, peripheral nerves can be simply divided anatomically in three layers: namely the endoneurium, perineurium and epineurium (Table 9.3). Axons are arranged in bundles and found in the endoneurium layer. In the endoneurium compartment (En), a single Schwann cell envelops several unmyelinated axons, and another Schwann cell provides multiple wrappings of the plasma membrane forming the myelin sheath of a myelinated axon. Schwann cells associated with both unmyelinated and myelinated axons are covered with a continuous basal lamina. Capillaries (Cap) are present within the endoneurial compartment and collagen fibres run primarily longitudinally



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Table 9.3 Anatomical layers of the peripheral nerve Layer

Primary component

Endoneurium

Basal lamina of type IV collagen, fibronectin and heparin sulphate proteoglycan Loose connective tissue of type I and type II collagen fibrils arranged longitudinally Fibroblasts, mast cells, macrophages and endoneurial fluid

Perineurium

Layers of type I and type II collagen fibrils Elastic fibres arranged circumferentially oblique and longitudinally Basal lamina with laminin, heparin sulphate proteoglycan and fibronectin

Epineurium

Bundles of type I and type III collagen fibrils Elastic fibres, fibroblasts, mast cells and fat cells

Table 9.4 Description of extracellular matrix (ECM) molecules of the peripheral nervous system Extracellular Active Distribution in ECM of matrix peptide nervous system molecules sequence

Major function

Collagen

RGD

Synthesised by Schwann cells, Schwann cell basement membrane

Major structural protein of lamina densa, stimulates ensheathment of neurons by Schwann cells

Fibronectin

RGD

Radial glia in developing cerebellum, skeletal muscle and Schwann cell basement membranes, ECM of neural crest

Guidance of migrating neuroblasts, and growing nerve fibres, Schwann cell mitogen

Entactin

–

Basal lamina of Schwann cells

Cell attachment

Heparan sulphate

–

Basement membrane of Schwann cells

Myoneural synaptogenesis, adhesion of culture neural cells to substrates, associated with a neurotrophic factor

Source: Carbonetto, 1984

between axons. Axons, Schwann cells, collagen and endoneurial fluid are bundled into a fascicle by the perineurium (Pe). Schwann cells ensheathing the axons actively deposit sheets of ECM. The molecular constituents found in the peripheral nerve ECM are listed in Table 9.4 (Carbonetto, 1984). ECM mechanically stabilises cells and is critical



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for myelination of axons. There are two major components: basal lamina surrounding the Schwann cells and axons, and endoneurial collagen, both formed during embryonic development. Basal lamina contain collagen type IV and V, laminin, fibronectin, entactin and heparan sulphate proteoglycan. Endoneurial collagen contains longitudinally arranged collagen fibrils that are mainly collagen type I, II, IV and V. Since ECM plays an important role in supporting the axons and Schwann cells, it may be advantageous to mimic the structure and biochemical signals found in the PNS ECM to enhance the performance of synthetic nerve constructs in bridging nerve deficits (Podratz et al., 2001). Haptotactic cues for nerve regeneration include the incorporation of contact-mediated signals such as ECM proteins, or short sequences of the ‘functional’ nucleotides of the ECM, to guide axons to extend to the synaptic targets (TessierLavigne and Goodman, 1996). It is well recognised that ECM components like collagen, laminin and fibronectin can improve the results of nerve repair. Collagen, a major component of the peripheral nerve’s ECM, consists of different isoforms that display different functions for the PNS. For example, collagen type I supports attachment and proliferation of Schwann cells through interaction with the cell surface integrins and collagen type IV is the main structural component of the basal densa. Numerous studies have evaluated collagen for nerve tissue engineering applications. Incorporation of collagen gel or collagen sponge in the lumen of nerve guides were shown to improve nerve regeneration (Nakamura et al., 2004; Dubey et al., 1999; Ceballos et al., 1999). Collagen type I filaments were found to enhance nerve regeneration in a 30 mm rat sciatic nerve deficit (Yoshii et al., 2002). Laminin, a major component of the basal lamina, is one of a series of structural proteins that activate the b1-integrin receptor. It plays a crucial role in cell migration, differentiation of Schwann cells and axonal growth. In vitro experiments have shown that neurite outgrowth can be enhanced on scaffolds that have been coupled with laminin (Yu et al., 1999; Rangappa et al., 2000; Labrador et al., 1998). In addition, short bioactive oligopeptide sequences of laminin (e.g. IKVAV and YIGSR) that can be recognised by specific receptors have been shown to improve neurite outgrowth (Huber et al., 1998; Luo and Shoichet, 2004). Fibronectin has been reported to affect the motility of Schwann cells and has been shown to influence axonal elongation, Schwann cell attachment and proliferation (Armstrong et al., 2007). In addition, fibronectin-containing matrices were formed in the lumen of synthetic conduits used to bridge nerve gaps (Longo et al., 1984). Experiments demonstrated the effectiveness of fibronectin mesh as a delivery vector of chemotactic cues in promoting nerve regeneration (Whitworth et al., 1996; Sterne et al., 1997), while providing support as a basal lamina scaffold.



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9.3.2 Current materials and scaffolds for nerve tissue engineering Lungborg (1988) is a pioneer in the evaluation of the use of non-biological conduits for nerve repair. Non-biodegradable tubes such as silicone or polytetrafluoroethylene (PTFE) conduits which were initially used to reconstruct the nerve after transection injuries have shown promising results. However these non-biodegradable conduits pose long-term complications, such as fibrosis and chronic nerve compression, that eventually require surgical removal of the conduits. Conduits made from bioabsorbable polymers (synthetic and naturally occurring) are promising alternatives to non-degradable tubes. Table 9.5 lists some of the common biodegradable materials used to construct tubular conduits that have been used in in vivo studies. The advantage of using natural biomaterials for nerve construct production is the presence of biological recognition which provides a better environment for tissue regeneration. However, natural biomaterials may exhibit inflammatory and antigenic responses, although collagen only exhibits these immunological responses minimally. In contrast, synthetic polymers can be produced reproducibly on a large scale with controlled properties of strength, degradation rate and microstructure. A drawback of synthetic polymers is the lack of biological recognition but this can be circumvented through the incorporation of cell recognition molecules or domains such as laminin or tripeptide RGD into these materials. In fact, natural materials, such as collagen, and synthetic polymers, such as poly(glycolic acid) (PGA), were mixed to fabricate tubes to bridge nerve gaps and this combination of materials has demonstrated excellent nerve regeneration over large nerve gaps (Matsumoto et al., 2000). The many key characteristics possessed by synthetic polymers, such as high porosity and surface area, structural strength and specific threedimensional shapes, outweigh the lack of biological recognition in their use in tissue engineering or other bioengineering applications, especially in nervous tissue repair.

9.3.3 Cells for nerve tissue engineering Glial cells, namely Schwann cells (SCs), are identified as being essential for successful peripheral nerve regeneration. During development, SCs actively deposit ECM comprising basal lamina sheets that surround individual axon-Schwann cell units. Following injury, denervated SCs will produce a number of neurotrophic factors to support the survival of injured neurons, promote macrophage infiltration to the injured nerve and provide substrate for axonal regeneration. Another important role of the Schwann cells involves the ensheathing and remyelination of regenerating axons. Although the mechanisms between axon and SC partnership during nerve regeneration



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General properties (e.g. degradation)a

b

a

Degradation between one and 9 months

Freeze-drying of hydrogel

Spinning of collagen fibres around a central mold to form tubular geometry

Dip coating

Interconnected pores with laminated appearance; inner layer consisting of densely packed fibres that are intertwined with multilayered structure

(Den Dunnen et al., 1995, 1998)

Porous

Semi-permeable (Li et al., 1992) (e.g. permeable to molecules the size of BSA)

Semi-permeable

(Bini et al., 2004b)

Semi-permeable

Electrospinning

Randomly arranged nanofibres provide relatively smooth surface Inner layer is dense to prevent ingrowth of fibroblast, outer layer is porous

(Evans et al., 1999, 2000; Widmer et al., 1998)

Porous, mean pore size of 12.1 µm

References

(Schlosshauer et al., 2006)

Permeability

Semi-permeable

Surface properties

Woven mesh – fabrication technique Solvent casting, Rough-looking owing to extrusion and open-pore structure particulate leaching

Fabrication technique

Degradation rate is dependent on the fabrication method. Commercially available nerve conduit

Alginate

Natural Collagenb

Poly(glycolic acid) Aliphatic polyester, crystalline, degradation in (PGA)b 6–12 months Poly(l-lactic) acid Aliphatic polyester, (PLLA) crystalline, slow degradation (e.g. > 24 months) Poly(lactic-coAmorphous, degradation glycolic) acid dependent on ratio of lactic (PLGA) acid and glycolic acid Amorphous, degradation Poly(lactic acidco-e-caprolactone) within year(s) (PLA-CL)b

Synthetic

Materials

Table 9.5 Some common materials used to fabricate nerve conduits that have been used in in vivo studies


177

are not exactly known, experiments have shown that a close relationship is present and regenerating axons could be misdirected by atypical SC processes (Thompson and Buettner, 2006; Chen YY et al., 2005). Guidance of Schwann cell migration, adhesion and maturation has been observed when cultured on various substrates (Schnell et al., 2007; Vleggeert-Lankamp et al., 2004). Incorporation of SCs in the nerve conduits achieved good regeneration in the sciatic nerve repair compared with the nerve conduits without SCs (Hadlock et al., 2000; Sinis et al., 2005). Coupling SCs in artificial nerve conduits is a potential strategy to improve peripheral nerve regeneration. However, application of SCs in clinics remains controversial because in vitro expansion of cultured SCs is limited and the potential rejection of non-autologous donor SCs has to be addressed. Stem cells are an alternative source of Schwann cells, such as embryonic stem cells (ESC), neural stem cells (NSC), mesenchymal stem cells (MSC) and so on. The list of various cell types differentiated from human ESCs (e.g. neurons, cardiomyocytes, hepatocytes) is continuously increasing. Pluripotent ESCs can be differentiated stepwise in the culture dish by recapitulating aspects of in vivo development and the use of relevant epigenetic factors. In contrast to pluripotent ESCs, somatic stem cells are believed to be multipotent and thereby capable of generating major cell types that are limited to the tissue of origin. Somatic/tissue-specific stem cells are the building blocks of organs during development and survive in specialised microenvironments (‘stem cell niche’) contributing to new cells throughout life (Fig. 9.1) (Gage, 2000). NSCs are (1) multipotent (have the ability to yield mature cells in all three fundamental neural lineages throughout the nervous system: neurons, astrocytes and oligodendrocytes), (2) have the ability to populate a developing region and/or repopulate an ablated or degenerated region of the CNS with appropriate cell types and (3) undergo ‘self-renewal’, that is the ability to produce daughter cells with identical properties. NSCs have been identified in vitro. However, no study has been able to demonstrate the existence of multipotent NSCs in vivo. NSCs (Heine et al., 2004) and trans-differentiated bone marrow stromal cells (Lopes et al., 2006; Keilhoff et al., 2006a, 2006b) that express glial phenotype show potential use of stem cell therapy for peripheral nerve tissue engineering. Grafting of glial-differentiated bone marrow stromal cells into peripheral nerve repair showed a beneficial effect on Schwann cell growth within the nerves, thus indirectly promoting axonal regeneration (Keilhoff et al., 2006b). However trans-differentiation of stem cells into neuronal or glial-like cells for use in clinical settings would require considerable evaluation to justify the use in the design of artificial nerve conduits. Besides stem cells, research into the use of olfactory ensheathing cells (Verdu et al., 1999), and genetic modified cells that over-express neurotrophins (Li et al., 2006; Ma et al., 2004; Menei et al., 1998; Tuszynski et al., 1998; Sayers et al., 1998) for



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Restriction

Cell

Source

Totipotent non-self-renewing

Zygote

Zygote

Pluripotent self-renewing

Embryonic stem cell

Blastocyst

Broad potential self-renewing

Multipotent stem cells

Embryo or adult brain, blood(?)

Limited potential limited self-renewal

Neural progenitor

Brain or spinal cord

Limited division non-functional

Committed neural progenitor

Brain subregion

Neuronal progenitor

Glial progenitor

Non-mitotic Differentiated functional

Neuron

Specific brain sites

Glia

9.1 Illustration proposing the classes of mammalian stem cells that can give rise to neurons, presented as a hierarchy beginning with the most primitive and multipotent stem cell and progressing to the most restricted. From Gage (2000). Reprinted with permission from AAAS.

nerve regeneration has also been described. Olfactory ensheathing cells that could be derived centrally from the olfactory bulb or derived peripherally from the olfactory epithelium possess similar phenotypes to Schwann cells



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and astrocytes. It is, however, suggested that in order to warrant the use of olfactory ensheathing cells in peripheral nerve regeneration, distinct advantages of these cells over Schwann cells would have to be shown (Schmidt and Leach, 2003).

9.3.4 Growth factors for nerve tissue engineering In the normal nerve state, cellular and molecular activities are dependent on the influence of neurotrophic factors secreted by the peripheral target cells. Neurotrophic factors are a family of polypeptides required for the survival of discrete neuronal populations. Different neurotrophic factors have been shown to exert biological action on various different sensory neurons. They mediate the chemotaxis processes that are required for successful re-establishment of functional synaptic connections. The action of neurotrophic factors is influenced by the neurotrophic receptors such that binding each individual neurotrophin with their specific receptor(s) will initiate a cascade of reactions. During nerve regeneration, axons can respond to specific neurotrophic proteins and grow back preferentially towards the original target organs to form functional synaptic innervations. Table 9.6 (Sterne et al., 1997; Zhang et al., 2004; Piotrowicz and Shoichet, 2006; Kapur and Shoichet, 2004; Chew et al. 2005, 2007; Terris et al., 2001; Aebischer et al., 1989; Sondell et al., 1999; Chen et al., 2006) shows the various neurotrophic growth factors and other related factors that have been studied for their potential use in peripheral nerve repair. Nerve growth factor (NGF), the first identified neurotrophic factor, is essential for the survival and maintenance of sensory and sympathetic neurons. It binds to the p75 neurotrophin receptor and the trk-A receptor, which are both associated with neurodegenerative disorders. Another growth factor that is closely related to NGF is brain-derived neurotrophic factor (BDNF). BDNF is expressed by skeletal muscle which primarily binds to the trk-B receptor and is important for the regulation of neuromuscular synapse development. Evidence obtained from in vitro experiments has shown that BDNF is an important biomolecule for the survival and differentiation of motor neurons and could assist in preventing neuron death after nerve injury (Vogelin et al., 2006). In vitro and in vivo studies have shown that neurotrophins-3 (NT-3) and -4 (NT-4) and ciliary neurotrophic factor (CNTF) bind to the trk receptors and the glycosyl-phosphatidylinositol-linked CNTF receptor, respectively, which are also required for the survival of certain motor neurons (Kingham and Terenghi, 2006; Kelleher et al., 2006). Glial cell line-derived neurotrophic factor (GDNF), which binds to tyrosine kinase receptor (RET) and GDNF receptor-alpha, has the ability to promote survival of both motor and sensory neurons (Kingham and Terenghi, 2006). Many studies have introduced neurotrophic factors into peripheral nerve



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BDNF

Brain-derived neurotrophic factor Glial-derived neurotrophic factor Fibroblast growth factor 2 Ciliary neurotrophic factor Vascular endothelial growth factor Insulin-like growth factor I Glial growth factor

GGF

IGF-I

VEGF

CNTF

FGF-2

GDNF

NT-3

NGF

Nerve growth factor

Neurotrophin-3

Abbreviation

Growth factor

Delivery vector(s)

Sensory neurons, small Microspheres, axons hydrogel, nanofibres

Major target

Sensory neurons, small Fibronectin scaffolds and medium size axons Neurotrophin Sensory neurons, large Collagen gel axons Neurotrophic factors Motor neurons Synthetic polymeric nanofibre conduit FGF Vascular endothelial Synthetic polymer cells conduit Nerve growth factors Schwann cells (injury Protein solution factor) Vascular endothelial Protein solution cells Inflammatory cells Covalently bound to (anti-inflammatory) conduit Neuregulin Schwann cells Alginate hydrogel

Neurotrophin

Neurotrophin

Family

Table 9.6 Neurotrophic factors for peripheral nerve regeneration

(Mohanna et al., 2003)

(Chen et al., 2006)

(Sondell et al., 1999)

(Zhang et al., 2004)

(Aebischer et al., 1989)

(Chew et al., 2007)

(Terris et al., 2001)

(Piotrowicz and Shoichet, 2006; Kapur and Shoichet, 2004; Chew et al., 2005) (Sterne et al., 1997)

References


181

tissue engineering strategies to evaluate the extent of nerve regeneration. Biodegradable polymers are usually used to produce the delivery vectors that provide sustained release of bioactive neurotrophic proteins. Since neurotrophins usually have a short plasma half-life, this must be considered when encapsulating growth factors into the design of the nerve conduit. Several concerns about the encapsulation of proteins for drug release include maximising therapeutic activity, maintaining protein stability and minimising toxic side effects. Release of neurotrophic factors such as nerve growth factor that were encapsulated in microspheres (Pean et al., 1998), matrices (Lee et al., 2003), scaffolds (Chew et al., 2005; Yang Y et al., 2005) or covalently bound to scaffolds (Chen P R et al., 2005) has demonstrated improved neurite outgrowth. The presence of exogenous nerve-related bioactive molecules such as NT-3 and BDNF has demonstrated enhanced nerve regeneration (Schmidt and Leach, 2003; Midha et al., 2003). Also, concentration gradients of NGF and NT-3 have been immobilised on hydrogel to promote neurite outgrowth (Kapur and Shoichet, 2004; Moore et al., 2006). Usually, protein release profiles are characterised by a marked initial burst release that results in a subsequently low amount of neurotrophic factors delivered to the injury sites (Xu et al., 2003). Efforts are being made to minimise this problem and to achieve controllable and continuous delivery of proteins using different biodegradable polymers or vectors such as microspheres or hydrogel. There is increasing interest in the use of nanoscale delivery vectors such as nanoparticles, nanotubes and nanofibres to provide improved sustained release of drugs (Goldberg et al., 2007). Among these, the nanofibres are promising as a dual functional scaffold that not only provides controlled drug delivery but also supports cell and tissue growth as the nano-scale topography closely mimics the natural ECM.

9.4

Biomimetic nanoscaffolds for peripheral nerve regeneration

9.4.1 Nerve conduit The ideal nerve conduit must be easily available, biodegradable, readily vascularised, have low antigenicity, be porous for oxygen and nutrients diffusion, and avoid long-term compression. Current approaches to nerve guide channel research include the use of biodegradable polymer scaffolds and incorporation of supporting cells such as SCs and macrophages, growth factors and ECM molecules in the nerve guides. A variety of natural materials such as vein, laminin, fibronectin and collagen have been used as nerve conduits and their advantages include, improved biocompatibility, decreased toxic effects and enhanced migration of supporting SCs. However, synthetic polymers are attractive for fabricating nerve conduits since variations in the chemical



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and engineering properties of these synthetic polymers allow alterations in the geometric configuration, biocompatibility, porosity, degradation and mechanical strength of these conduits. These tubular conduits have been fabricated from non-degradable synthetic materials such as silicon (Gibson et al., 1989), polycarbonate (Harvey et al., 1994) and poly(acrylonitrile-covinyl chloride) (Aebischer et al., 1990), as well as biodegradable polymers such as poly(lactic acid) (PLA) (Evans et al., 1999; Yang et al., 2004), poly(lactic-co-glycolic acid) (PLGA) (Hadlock et al., 1998; Widmer et al., 1998; He et al., 2009), poly(glycolic acid) (PGA) (Kiyotani et al., 1995), poly(lactic acid-co-e-caprolactone) (PLA-CL) (Den Dunnen et al., 1996; Nicoli Aldini et al., 1996) and polyhydroxybutyrate (PHB) (Hazari et al., 1999). Tiny tubes with an inner diameter of 1.27 mm and a length of 16 mm and fibre architecture were developed by a microbraiding technique so as to function as nerve guide conduit. The feasibility of in vivo nerve regeneration in the rat sciatic nerve model with a 12 mm gap was also investigated (Bini et al., 2004a). One month after implantation, 9 of 10 rats showed successful nerve regeneration, none of the implanted tubes showed tube breakage and there was no inflammatory response after implantation. Conduits fabricated from non-degradable synthetic materials remain in situ as a foreign body and may elicit an inflammatory response, limiting nerve regeneration. In contrast, conduits fabricated from biodegradable scaffolds offer the advantage of resorption when nerve regeneration is complete and reduce the chance of inflammatory responses. Researchers worldwide are investigating the use of nanofibres for a variety of applications. A key area for the use of electrospun nanofibres is their application as scaffolds for the tissue engineering of nerve, skin, bone and so on. Since the basic structural support of natural ECM consists of collagen fibres and fibrils with nanometre scale diameters, many researchers are turning to fabrication of the ideal ECM scaffold in tissue replacement by electrospinning (Teo and Ramakrishna, 2006). Research has shown that by using electrospun fibres with their ECM-mimetic nanotopography, certain signal transduction pathways vital in normal cellular activity can be activated in the absence of chemical cues (Nur-E-Kamal et al., 2005). Electrospinning offers many advantages over other methods of fabricating scaffolds. Other than its ability to fabricate nanometre diameter fibres, it can be used in a range of materials to produce continuous nanofibres. Synthetic non-biodegradable polymers, biodegradable polymers, natural polymers, composites and even ceramic nanofibres have been electrospun (Ramakrishna et al., 2005). Although a typical electrospun fibre scaffold is in the form of a two-dimensional nonwoven fibre mesh, other assemblies such as tubular scaffolds (Fig. 9.2) (Bini et al., 2004b), braided fibres (Bini et al., 2006) and aligned fibrous scaffolds (Yang et al., 2005) can be made through simple modification of the electrospinning setup (Teo et al., 2005). Research on



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Syringe PLGA polymer solution

+ ve High voltage

PLGA nanofibres

Rotating teflon tube

– ve High voltage

5kV X1, 000 10 µm

9.2 Schematic diagram of the electrospinning set up for the fabrication of nanofibrous tubes for peripheral nerve regeneration adapted from Bini et al. (2004b).

the use of the electrospun nanofibrous nerve grafts and the behaviour of nerve cells on them has been reported (Bini et al., 2005; Yang et al., 2004; Yang F et al., 2005; Christopherson et al., 2009). The flexibility and ease of incorporating other materials and additives (Prabhakaran et al., 2008a, 2008b; Ghasemi et al., 2008; Koh et al., 2008) into electrospun nanofibres means that the resultant scaffold can be tailored to the specific requirements of the particular tissue. For example, bilayered nanofibrous conduit and longitudinally aligned nanofibrous intra-luminal guidance channels were fabricated by electrospinning (Fig. 9.3).

9.4.2 Intraluminal guidance channel An important process in nerve regeneration for repairing transected nerves involves the migration of SCs into the synthetic nerve grafts from the proximal and distal nerve stumps (Brandt et al., 1999), subsequently forming cellular bridges across the interstump gap that are termed bands of Büngner (Fig. 9.4). These physical structures play a significant role in nerve regeneration as they act as guiding cues and substrates for growth cone adherence and axon outgrowth towards the distal end. The longitudinally oriented cell



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(a) (d) Outer layer of conduit

(b) Nanofibrous conduit

Inner layer of conduit

Single guidance channel

(c)

Nanofibrous nerve construct

Bundle of guidance channels

9.3 Macrographs of nanofibrous nerve construct. (a) Bilayered nanofibre nerve conduit, (b) intraluminal guidance channels, (c) nerve conduit containing intra-luminal channels and (d) scanning electron micrographs (SEMs) of nanofibrous nerve constructs (Koh et al., 2010).

matrix organises axon extension such that if regenerating axons accidentally leave the supporting matrix, axonal elongation usually terminates, resulting in the formation of neuroma (Lietz et al., 2006). Tubulation devices that utilise hollow conduits provide a physical substrate to guide sprouting axonal elongation towards the distal stumps, thus preventing aberrant outgrowth of axons and formation of scar tissue that impede nerve repair. In a hollow conduit, formation of fibrin matrix occurs which is subsequently infiltrated with SCs and non-neuronal cells. Formation of physical cellular bridges (i.e. bands of Büngner) will be created by the proliferated SCs that are shown to influence nerve regeneration and can act as scaffolds for advancing axons. If the nerve gap deficits are too great, formation of fibrin matrix and bands of Büngner might be compromised, thus affecting nerve regeneration and



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Distal

Proximal Axon Macrophage sprouts Schwann cell Growth cone

Basal laminae

Bands of Büngner

9.4 Formation of bands of Büngner by the Schwann cells during nerve regeneration (Lundborg, 1988). Table 9.7 Fillings and scaffolds in the lumen of nerve conduits Fillings

Materials

Features

Denatured muscle Allogenic and scaffolds acellular

Presents longitudinally oriented basal lamina of muscle that also contains ECM molecules such as laminin and collagen type IV. Hydrogel Naturally occurring Highly porous. Can be materials such as coupled with neuritecollagen, Matrigel, promoting biomolecules hyaluronic acid, etc. Magnetically Collagen Fibres that mimic the aligned fibrils natural ECM. Can be hydrogel coupled with neuritepromoting biomolecules Synthetic or Collagen and Micro-sized fibres (5 naturally occurring synthetic polymer µm–150 µm). polymeric (e.g. PLLA, PGA) Can be coupled with filaments (intraneurite-promoting luminal guidance biomolecules. channels)

References (Meek et al., 1996, 2004)

(Labrador et al., 1998)

(Ceballos et al., 1999)

(Ngo et al., 2003; Yoshii et al., 2002; Rangappa et al., 2000; Cai et al., 2005)

functional recovery. Researchers have thus introduced physical scaffolds within the lumen of the conduits potentially to enhance axonal regeneration (Table 9.7). Incorporation of fillings in the conduits aims to create a permissive regenerative environment that can regulate appropriate axonal growth and promote excellent functional activities of the non-neuronal cells essential for



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nerve regeneration. The use of denatured muscles has been shown effectively to direct outgrowing axons as the presence of longitudinally oriented basal lamina of the muscle tissue that contains biochemical molecules (Meek et al., 1996) was considered to be beneficial. Longitudinally aligned guidance scaffolds have been described to provide excellent guiding function for regenerating axons. A study has shown that aligned collagen fibrils in collagen conduits enhanced the regeneration of sciatic nerves in mice compared to conduits containing unaligned collagen gel (Dubey et al., 1999). Microgrooved polymer filaments inside nerve conduits allowed nerve regeneration to occur with highly oriented axon growth without meandering (Lietz et al., 2006). Incorporation of intraluminal guidance channels can thus potentially improve nerve regeneration in large nerve defect injuries since they can mimic the bands of Büngner (Schmidt and Leach, 2003; Ngo et al., 2003; Rangappa et al., 2000; Cai et al., 2005; Wen and Tresco, 2006). For 18 mm and 30 mm nerve lesion gaps, Cai et al. (2005) and Yoshii et al. (2002) have reported that conduits containing physical filaments could enhance axonal regeneration in animal studies. This assists in recreating an environment that more resembles the native nerve tissue organisation architecture. Bellamkonda’s group demonstrated that aligned poly(acrylonitrile-comethylacrylate) (PAN-MA), nanofibre mesh (10–12 layered) stacked in the conduit could enhance the growth of dorsal root ganglions (DRGs) in vitro for bridging long peripheral nerve gaps (Kim et al., 2008). Furthermore, nerve regeneration through a one-film guidance channel, containing a single continuous thin-film of aligned PAN-MA fibres, in comparison with a three-film channel in the conduit that provided two additional thin-film tracks was evaluated (Clements et al., 2009). Interestingly, the one-film channels supported more enhanced regeneration compared with the threefilm channels in terms of regenerated axon profile counts and measures of nerve conduction velocity. These results suggest that minimal levels of appropriately positioned topographical cues significantly enhance guidance channel function by modulating endogenous repair mechanisms, resulting in effective bridging of critically sized peripheral nerve gaps. However, optimised channel structure is far from being understood. We fabricated the intraluminal guidance channels to mimic the basal laminae of peripheral nerve that could act as substrates for axonal outgrowth and Schwann cells migration for nerve repair. Figure 9.3 shows the macrographs and scanning electron micrographs of the nerve construct. The intraluminal guidance channel is made from aligned nanofibrous yarns. In addition, biomolecules such as laminin and NGF were incorporated in the nanofibrous nerve construct to determine their efficacy in in vivo nerve regeneration. In some cases, the nanofibrous conduit with aligned nanofibrous yarn as an intra-luminal guidance channel performed better than the autograft in terms of muscle reinnervation and withdrawal reflex latency tests (Koh et al., 2010).



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However, the axon density count is highest in the autograft at mid-graft. Despite observing that 50% of the empty nanofibrous conduits collapsed in this study, the nanofibrous conduits fabricated generally possessed adequate mechanical properties to maintain a stable path for nerve regeneration across the injury site. Furthermore, introducing nanofibrous guidance channels in the lumen of the conduit would reduce the probability of conduit collapse, whilst not hindering the nerve regeneration process. A study performed by Ngo et al. (2003) on the effect of packing density of intraluminal guidance channels in conduits has suggested that a 7.5% packing density would be optimal for bridging nerve defects, even when a 3.5% packing density showed a greater number of myelinated axons in an in vivo experiment. Here, we used a packing density of intraluminal guidance channels of 10% to evaluate axonal regeneration in the transected sciatic nerve. On the other hand, the synergistic effects of nanotopography and biochemical cues possessed by the novel nerve construct described herein present a new avenue for repairing a large peripheral nerve gap (Koh et al., 2010).

9.5

Stem cell therapy with nanofibre for nerve regeneration

9.5.1 Neuronal differentiation of stem cells Transplantation of stem cells is one of the most promising strategies for neuroregeneration and has been exploited in a number of disease settings including CNS repair and multiple sclerosis (Payne et al., 2008). There are two types of stem cell: the embryonic stem (ES) cells, found in the inner cell mass of the early embryo, and tissue-specific adult stem cells found in the bone marrow blood, cornea and retina, brain, skeletal muscles, dental pulp, liver and skin and so on. Stem cells located in continuously renewing tissues, such as skin, gut or bone marrow, are able to regenerate or repair these tissues throughout life. Among adult stem cells, bone marrow derived-MSCs (BM-MSCs) represent a potential for autologous stem cell based replacement therapies, since they do not elicit graft versus host disease. MSCs are currently being investigated for the treatment of diverse disorders including graft versus host diseases, osteogenesis imperfecta and so on, because of their ease of propagation, functional attributes and/or poor immunogenicity (Bianco et al., 1999). The capacity of MSC for neuronal differentiation offers more hope for developing therapeutics for neuro-regenerative diseases. Recent studies have shown that BM-MSCs can be induced to express a neuronal phenotype in vitro under specific experimental conditions. SanchezRamos and co-workers investigated differentiation of bone marrow stromal cells into neural cells (Sanchez-Ramos et al., 2000; Sanchez-Ramos, 2002). For neural induction, MSCs were re-plated in culture dishes in the presence



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of a neuronal growth medium (N5) supplemented by 5% horse serum, 1% FBS, transferrin (100 mg ml–1), putrescine (60 mM), insulin (25 mg ml–1), progesterone (0.02 mM), selenium (0.03 mM), all-trans-retinoic acid (0.5 mM) and brain-derived neurotrophic factor (BDNF) at a concentration of 10 ng ml–1. Various chemical inductors such as butylhydroxytoluene (BHT), butylhydroxyanisole (BHA), dimethyl sulphoxide (DMSO), 2-mercaptoethanol, 3-isobutyl-1-methylxanthine (IBMX) and 5-aza-cytidine were used alone or in combination with cytokines for neural differentiation of MSCs. For example, Woodbury et al. (2000) observed that in the presence of b-mercaptoethanol and DMSO, BM-MSC may differentiate into cells expressing neuron specific enolase (NSE) and neurofilament. Moreover, it has been described that in the presence of epidermal growth factor (EGF) and BDNF, stromal cells differentiate into neural cells expressing both neuronal and glial markers (Neu-N, nestin, GFAP). Human and murine BM-MSC are capable of differentiating into cells that express several neural proteins and resemble immature neurons or glial cells. But the expression of one or even two neuronal proteins does not prove that the cell bearing these ‘neuronal markers’ is capable of all the complex functions of a neuron. Other research has been done in the field of MSCs differentiating into SCs, which provide neurotrophic support and are able to myelinate the growing axons (Keilhoff, 2006a). In this study, MSCs were expanded in growth medium for three days and an additional day with an additional 1 mM b-mercaptoethanol. Subsequently, the medium was replaced by the transdifferentiation medium (a-MEM, 10% FCS, 2 mM l-glutamine, 100 U ml–1 penicillin and 100 mg ml–1 streptomycin, 35 ng ml–1 all-trans-retinoic acid). For the final transdifferentiation step, cultures were incubated in the transdifferentiation medium supplemented by forskolin 5mM, bFGF 10 ng ml–1, PDGF-AA 5 ng ml–1, Her-b 200 ng ml–1 and insulin-like growth factor (IGF-1) 10 ng ml–1 for 8 days with a media change every other day. Although the differentiation progress is quite complex, BM-MSCs can still be considered as seed cells for peripheral nerve tissue engineering because of their ability to promote axonal regeneration (Keilhoff, 2006b; Hou et al., 2006). In vivo studies by Chow et al. (2000) showed that stem cells implanted in injured spinal cord differentiated into glial and neuronal cells. Kopen and colleagues (1999) showed the potential of MSCs to differentiate into cells with neural lineage by injecting them into the lateral ventricle of neonatal mice. This also shows that a greater control over stem cell differentiation either by in vivo treatments or by utilising stem cells restricted to neural lineages might allow the transplantation of stem cells for nerve regeneration. Considering the immunological limitation of allogeneic neural stem cells, the ability of MSCs to induce along neural pathways would be of great benefit in neural repair. Moreover, the multi-potency of adult stem cells has generated tremendous interest in their potential therapeutic value, because these cells



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may be used in autologous treatments and do not raise any ethical concerns. However, functional recovery with transplantation of induced neural cells on nanofibres for peripheral nerve defects in animal models remains challenging (Hou et al., 2006). On the other hand, bioelectricity present in the human body plays an integral role in maintaining normal biological functions such as signalling of the nervous system, muscle contraction and wound healing (Shi et al., 2008). Electrical stimulation is a relatively simple, flexible and feasible technique both in vivo and in vitro to be applied for both two-dimensional and three dimensional cultured cells (Sun et al., 2006). Alternatively, researchers have also sought to incorporate electrical signals directly into biomaterials (Rivers et al., 2002). Electrical stimulation has been shown to be an effective cue to enhance neurite and axonal outgrowth from in vitro study (English et al., 2007) and animal study (Huang et al., 2009), and several theories have been suggested to explain the effect of electric stimulation on nerve regeneration (Patel and Poo, 1982; Kotwal and Schmidt, 2001). DC electrical gradients of voltages are present within the tissues (endogenous electrical fields) and various conditions have been applied to electrical stimulation on structured materials.

9.5.2 Effect of nanofibres on neuronal differentiation Nano-scale topographic features such as fibre, steps, grooves, pillars and pits (or pores) modulate cell behaviour. Topographical features, including fibre dimensions and pattern, are important aspects in developing fibrous scaffolds for nerve tissue engineering. Previously, we have found that nanofibrous scaffolds support nerve stem cell growth significantly more than the microfibres of random or aligned patterned fibres (Fig. 9.5) (Yang F et al., 2005). It was found that the NSCs elongated and formed neurites and grew out along, and parallel to the fibre direction for the aligned scaffolds, whereas the fibre diameter did not show any significant effect on the cell orientation. Further investigations were done on the effects of topographical features by optimising the electrospinning parameters to fabricate aligned poly(l-lactide) (PLLA) fibres with average diameters of 307, 500, 679 and 917 nm and random fibres with average diameters of 327, 545, 746 and 1150 nm, respectively as shown in Fig. 9.6 (He et al., 2010). Neonatal mouse cerebellum C17.2 cells exhibited significantly different growth and differentiation depending upon fibre dimension and pattern. On aligned fibres cell viability and proliferation was better on 500 nm fibres and reduced on smaller or larger fibres. However, on random fibres cell viability and proliferation was significantly increased with the smallest (350 nm) and largest (1150 nm) diameter fibres. Yim and co-workers showed that topography with 350 nm width can



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80.00 µm

40.00 µm (b)

(a)

80.00 µm

40.00 µm (d)

(c) 120

Neurite length (µm)

*

80

40

0

ANF

AMF

(e)

RNF

RMF

9.5 Laser scanning confocal microscopy (LSCM) micrographs of immunostained neurofilament 200 kD in neural stem cells (NSCs) after 2 days of culture; (a) on aligned nanofibre-ANF, low magnification (x200); (b) on ANF, high magnification (x400); (c) on aligned microfibre-AMF, low magnification (x200) and (d) on AMF, high magnification (x400). (e) The average length of the longest neurite per cell was measured for 50 randomly selected cells per scaffold (mean for n = 3 ± standard error of mean) (*p < 0.05). RNF and RMF refer to random nanofibre and random microfibre respectively (Yang et al., 2005).



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Nerve tissue regeneration A3

A5

A7

A9

R3

R5

R7

R11

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9.6 SEM micrographs of aligned (A) and random (R) PLLA electrospun nanofibres with different diameters and arrangement; A3 to A9: aligned fibres with diameter of 300 nm, 500 nm, 700 nm and 900 nm respectively; R3 to R11: random fibres with diameter of 300 nm, 500 nm, 700 nm and 1100 nm respectively. (He et al., 2010).

also influence cellular response including enhanced neural differentiation of progenitor cells (Yim et al., 2007). They produced nanogratings with soft lithography and differentiated MSCs from nerve cells on this surface by adding retinoic acid. The combination of retinoic acid induction and nanotopographical induction showed the highest microtubule-associated protein (MAP2) expression in a quantitative polymerase chain reaction study. This study suggests that nanotopography could influence stem cell differentiation into specific non-default pathways, such as transdifferentiation of MSCs, and the topographical cues could be combined with chemical cues to create a synergistic condition for neuronal differentiation. We investigated the feasibility of functionalising electrospun nanofibrous scaffolds by coupling laminin onto poly( l-lactic acid) (PLLA) nanofibres. Laminin was successfully added to nanofibres using covalent binding, physical adsorption or blended electrospinning procedures (Koh et al., 2008). PC12 cell (derived from the pheochromocytoma of the rat adrenal medulla) viability and neurite outgrowth assays confirmed that the functionalised nanofibres were able to enhance axonal extensions. Especially, extensive neurite outgrowth was observed on blended laminin/PLLA nanofibres which suggest that electrospinning of laminin in the solvent at high voltage had not adversely affected the bioactivity of the ECM protein. Compared to covalent immobilisation and physical adsorption, blended



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electrospinning of laminin and synthetic polymer is a facile and efficient method of modifying nanofibres for their potential use in peripheral nerve regeneration applications. Alternatively, P(LA-CL) and P(LA-CL/collagen) nanofibrous scaffolds were fabricated by electrospinning (Prabhakaran et al., 2009). The differentiation of MSCs was carried out using neuronal inducing factors including b-mercaptoethanol, epidermal growth factor, NGF and brain derived growth factor in DMEM/F12 media for a period of 5 to 15 days on electrospun nanofibres. The proliferation of MSCs evaluated by MTS assay showed that the cells grown on PLA-CL/collagen nanofibrous scaffolds were comparatively higher (80%) than those on PLA-CL nanofibres as we expected. SEM results showed that MSCs differentiated on PLA-CL/ collagen nanofibrous scaffolds showed neuronal morphology, with multipolar elongations, and expressed neurofilament and nestin protein. Although more studies are required to understand the neuronal cell maturation and the molecular mechanisms responsible for long-term neuronal differentiations, we provided better alternatives for neurotransplantation of neuronal cells differentiated from stem cells and an accessible tissue in every individual, on biocompatible electrospun nanofibrous scaffolds. In order to facilitate electrical stimulation for the neural differentiation of the stem cells, we fabricated a novel nanofibrous scaffold containing polyaniline (PANI), PCL and gelatin (labelled PANI/PG). Proliferation and differentiation of NSCs (C17.2 neural stem cells) on electrospun electrically conducting nanofibrous scaffolds were studied to understand the effect of incorporation of a conductive polymer (PANI) into a biodegradable polymeric composite for nerve tissue engineering (Ghasemi-Mobarakeh et al., 2009). The overall results based on the proliferation and morphology of NSCs on PG and PANI/PG nanofibrous scaffolds indicated that the incorporation of PANI into PG does not hinder the biocompatibility behaviour of PANI/ PG nanofibrous scaffolds and NSCs interacted favourably with PANI/PG nanofibrous scaffolds. A constant voltage (1.5 V) for a period of 60 min was applied for the above scaffolds with seeded cells. The average neurite length for NSCs grown on PANI/PG with application of electrical stimulation for 1 hour and without electrical stimulation was found to be 30 ± 8 and 22 ± 7 mm, respectively, indicating that the application of electrical stimulus for 1 hour to NSCs cultured on PANI/PG significantly enhanced the neurite outgrowth.

9.6

Conclusion and perspectives

Functional nerve regeneration is a complex and delicate process. Even peripheral nerves consist of a relatively extensive network of blood vessels to maintain the structural and functional integrity of peripheral axons. For



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large nerve gap repair, providing sufficient adequate vascularisation may be critical for its success in providing sufficient nutrients for proliferating SCs and possibly maintain the survival of neurons. Together with vascular endothelial growth factor, platelet-derived growth factor-BB (PDGF-BB) may be required for formation of new matured blood vessels to aid in axon extension and maturation. The nerve conduit and intraluminal guidance channels described herein that possess longitudinally aligned nano-topograhical fibres have shown promising results in supporting directional neuronal growth across lesions in vivo to promote better functional recovery. However, further comprehensive analysis can be done to investigate the spatial and temporal progression of axons from the proximal to the distal ends when using this nanofibrous construct. This will provide important information about how the nanofibrous intraluminal guidance channels acted as scaffolds to support nerve regeneration. Intraluminal guidance channels can also be made with optimised nanofibre architecture (e.g. hollow inner structure). The hollow inner structure will provide more surface area for regenerating axons and SCs to adhere to and extend, which will limit the hindrance of nerve regeneration in the lumen of the conduit. There is no single approach in the design of nerve construct that can be effective for use in clinical peripheral nerve repair. To improve nerve regeneration and the outcome after various injuries there is a requirement to explore new research avenues (Dahlin, 2009). There are several other aspects which require attention from the clinical perspective. Better therapies may be required to improve the ability of the surviving neurons to regenerate and increase the speed of axonal outgrowth. For long gap repair, strategies could be devised to prevent the loss of basal laminae of the distal nerves because their maintenance is critical for regenerating axons to have a support to reinnervate and form synaptic contacts with the appropriate targets (Hoke, 2006). Further research in basic science, such as developmental biology, and applied sciences, such as bioengineering, is required to produce more suitable substitutes that can provide remarkable outcomes in the field of peripheral nerve repair and regeneration. The rapid development of stem cell biology represents a strong foundation for neural repair (Jandial et al., 2007). So far, experimental evidence suggests that stem cell therapy may be applicable for treating patients in the future, especially stem cell differentiation on nanofibres, although we should highlight current problems in the progressing stem cell field for clinical applications.

9.7

References

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10

Heart valve tissue regeneration

M. S i m o n e t, A. D r i e ss e n - M o l, F.P.T. B aa i j e n s and C.V.C. B o u t e n, Eindhoven University of Technology, The Netherlands

Abstract: Creating a functional valve equivalent to the appropriate tissue structure and mechanical properties of the heart is a key issue in the area of heart valve tissue engineering. The design and production of threedimensional scaffolds with both microscopically and macroscopically optimised material properties to regulate tissue development and ultimately valve functioning is believed to be critical. Electrospinning offers many ways of developing such scaffolds. In this chapter we aim to provide an overview of the various requirements for these scaffolds as well as presenting ways of achieving them using electrospinning. Key words: bioactive scaffold, cellular microenvironment, electrospun fibre morphology and organisation, functional scaffold design, heart valve tissue engineering.

10.1

Introduction

Most human tissues have a limited regeneration potential. In the case of tissue damage, the recovery of tissue structure and function is often incomplete, usually leading to scar tissue. For fully differentiated, loadbearing cardiovascular tissues, such as heart valves, regeneration is unlikely and replacement is a frequently applied therapy. Tissue engineering solutions are intensively studied for this purpose, mainly because of the option for creating a living tissue replacement that can grow and remodel in response to changing environmental conditions. Key research questions in the area of heart valve tissue engineering are: ‘Can we create a functional valve equivalent with appropriate tissue structure and mechanical properties?’ and ‘Can we control tissue formation and organisation by designing scaffolds and culture conditions that favour these processes?’. With recent advances in developmental and (stem) cell biology, tissue engineering is becoming increasingly oriented towards designing biologically inspired cellular microenvironments, aimed at guiding cell growth, differentiation and functional tissue organisation. The premise is that in order to unlock the full potential of the cells, at least some aspects of the three-dimensional (3D) tissue environment associated with their renewal, differentiation and organisation needs to be mimicked in the applied scaffold materials. The selection of the material, as well as the processing of the 202 © Woodhead Publishing Limited, 2011


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material into a 3D structure, therefore, becomes increasingly important to achieve both microscopic and macroscopic relevant material properties that favour tissue development and ultimate valve functioning. In this chapter we discuss the target tissue to be replaced by engineered heart valve equivalents, the specific tissue requirements that may dictate macroscopic and microscopic scaffold properties for this purpose and the different electrospinning settings and modalities that can be applied to meet these requirements.

10.2

Tissue to be replaced: heart valves

The human heart contains four heart valves, acting as one-way doors to guide blood flow into the proper direction in both the systemic and the pulmonary circulation. The atrioventricular valves (tricuspid and mitral valve) prevent backflow of blood from the ventricles to the atria, while the semilunar valves (pulmonary and aortic valve) prevent backflow from the arteries into the ventricles during diastole. All valves are believed to function in a passive pressure-driven manner. They open when pressure gradients force the blood forwards and close when backward pressure gradients push the blood backwards. The atrioventricular valves are anatomically and functionally different from the semilunar valves, the first being supported by papillary muscles that prevent leaflet prolapse during ventricular contraction. This chapter focuses on semilunar heart valves. Heart valve disease causes either insufficient opening (stenosis) or closing (regurgitation) dynamics of the valve, or a combination thereof, which ultimately will result in heart failure. Congenital heart disease affects 1% of all newborns and often has its origin in abnormalities of one of the valves or its function (Warnes, 2005). A common cause for acquired heart valve disease is rheumatic fever, currently still persisting in developing countries, thereby affecting children and young adults (Nkomo, 2007; Zilla et al., 2008). In industrialised nations, acquired heart valve disease is mainly considered a degenerative pathology, predominantly affecting the elderly (Iung et al., 2003; Goldbarg et al., 2007). In general, heart valve diseases are considered to be a worldwide major public health problem causing significant morbidity and mortality. The left-sided valves (aortic valve and mitral valve) are most prone to degenerative dysfunction in adult patients as these are located in the systemic circulation and are thereby exposed to harsh haemodynamic conditions. When heart valve disease is diagnosed, the affected valve can either be repaired or replaced. Despite the existence of an extensive toolbox for valve repair and excellent surgical techniques, seven out of ten diseased valves are not suitable for repair (Yacoub and Cohn, 2004) and have to be replaced. With the continuous growth and ageing of the world’s population, the social and economic impact of heart valve disease will increase (Nkomo



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et al., 2006). Therefore, the number of patients requiring a heart valve replacement is expected to triple from 290 000 in 2003 to over 850 000 in 2050 (Yacoub and Takkenberg, 2005). Nowadays, many valve replacement types are available, mainly classified as mechanical valves or bioprostheses, each with their own advantages or disadvantages (Zilla et al., 2008). Mechanical valves offer excellent structural durability, but are prone to thromboembolic events, thereby committing patients to daily anticoagulation. Bioprostheses are less susceptible to thromboembolic events, but undergo structural valve degeneration which necessitates re-operation. Homografts, better known as donor valves, represent the most ideal option for haemodynamic behaviour, but these have limited availability. The selection of appropriate heart valve replacement is dependent on patient characteristics, such as the patient’s age, lifestyle and tolerance to the use of anticoagulants (El Oakley et al., 2008). Although life expectancy is significantly improved by valve replacement, none of the currently available valve prostheses can fully restore native valve function as they lack growth, remodelling and adaptation capacity. Heart valve tissue engineering has the potential to overcome the limitations of today’s valve prostheses by creating an autologous living valve replacement that can grow or adapt to changing functional demands. Tissue engineering is an emerging interdisciplinary field, applying the principles and methods of engineering to the development of biological substitutes that can restore, maintain or improve tissue function (Langer and Vacanti, 1993). In the conventional tissue engineering paradigm, cells are isolated from the patient and expanded and subsequently seeded onto an appropriate scaffold material, after which the cells are stimulated to form tissue in bioreactor systems. These tissues can then be implanted into the patient from whom the cells were taken to serve as autologous living implants. The use of this paradigm for rendering living semilunar heart valve substitutes has been extensively studied using various cell sources and scaffold materials. The most obvious choice for a heart valve scaffold material is a homograft or xenograft depleted of cells, to render so-called decellularised matrices. Although these may provide the most natural template for valve remodelling and growth, homografts have limited availability, while xenografts suffer from the risk of zoonoses. In addition, successful ingrowth of cells and recapitulation of tissue growth and remodelling has not yet been proven in humans. The use of synthetic scaffold materials represents a promising alternative approach, allowing the tissue to develop while the scaffold degrades, without supply limitations and risk of zoonoses. Commonly used synthetic scaffold materials for heart valve tissue engineering are poly(glycolic acid) (PGA), poly(lactic acid) (PLA), poly(4-hydroxybutyrate) (P4HB), poly(hydroxyalkanoate) (PHA) and poly(hydroxyoctanoates) (PHO). The possibility of heart valve tissue engineering using synthetic scaffolds was demonstrated in sheep by replacement of a single pulmonary valve leaflet with



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a tissue engineered equivalent based on rapid degrading synthetic scaffolds (Shinoka et al., 1995; Kim et al., 2001). Later, full trileaflet valves were fabricated based on rapid degrading synthetic scaffolds, which demonstrated remodelling potential into native valve mimicking the structures in sheep (Hoerstrup et al., 2000; Sodian et al., 2000; Stock et al., 2000; Sutherland et al., 2005). Although the approach seems promising, it is obvious that the remodelling initially occurs via thickening of the engineered tissues (Gottlieb et al., 2010). This thickening may be the result of non-physiological valve characteristics at the time of implantation and might benefit from improvements in scaffold development to mimic the native valve more closely. Growth of tissue engineered prostheses based on synthetic scaffolds has been demonstrated for large blood vessel substitutes (Hoerstrup et al., 2006) and should be further explored for tissue engineered heart valves. To reduce mortality and morbidity risks with conventional heart valve replacement surgery, minimally invasive valve replacement techniques have rapidly evolved as an alternative treatment option (Lutter et al., 2004; Walther et al., 2007). The feasibility of using tissue engineered heart valves based on rapid degrading synthetic scaffolds integrated into self-expanding nitinol stents and implanted via a minimally invasive valve implantation technique was recently demonstrated in sheep (Schmidt et al., 2010). Only recently, the tissue engineering paradigm has become increasingly oriented towards a so-called in situ approach, using the body as bioreactor system for the tissue formation phase (Mol et al., 2009). This in situ tissue engineering approach represents an economically attractive approach, providing off-the-shelf availability and less risk of contamination as extensive in vitro culturing is omitted and a cheaper and shorter production process is available. The shift from the in vitro to the in situ tissue engineering approach emphasises the role of the scaffold, not only as a template for the cells to form a tissue, but also to enable cell invasion and to maintain valve functionality whilst the tissue is developing. This may require the use of slower degrading scaffold materials, such as poly(e-caprolactone) (PCL) and the inclusion of bioactive compounds into the scaffold to guide tissue development. To summarise, the valve scaffold should provide a natural template for cell invasion and tissue formation. Further, it should mimic native valve functionality and allow insertion via a minimally invasive approach. The development of such a valve scaffold represents a true challenge within the heart valve tissue engineering field and may be facilitated by electrospinning.

10.3

Specific tissue requirements as a blueprint for scaffold properties

The scaffold plays a pivotal role in both the in vitro and in situ heart valve tissue engineering approach as heart valves are geometrically, mechanically



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and functionally challenging structures. To be able to define scaffold requirements, one should be acquainted with the tissue requirements of the semilunar human heart valve from the macro- to the microscopic level. Here, information is provided on geometry, function, structure, composition and cell-matrix interactions of human semilunar heart valves to serve as a blueprint for scaffold properties.

10.3.1 Geometry and function of semilunar heart valves The semilunar valves consist of three flexible leaflets, three sinuses and a root. The average aortic valve diameter is 23 mm and the average pulmonary valve diameter is 26 mm for adult men. The valve diameter is directly correlated to the body surface area (Capps et al., 2000). Owing to their u-shaped attachment to the arterial wall, the leaflets form cusps that intimately fit together during closure of the valve. The average height of the aortic root from the bottom of leaflet attachment to the top of the commissures is 18 mm. An aortic leaflet measures on average 0.4–0.5 mm in thickness and 16 mm in length from attachment to the root to the free edge (Swanson and Clark, 1974). A coaptation area at the free edge of the leaflets prevents prolapse during valve closure. On average, the coaptation area of an aortic valve leaflet is 5 mm in length (Swanson and Clark, 1974). Figure 10.1 demonstrates the geometry of an adult aortic valve. This unique geometry enables lifelong opening and closure of the valve, on average 100 000 times a day or 3.7 billion times in a lifetime. In adult subjects, the pressure difference over the leaflets during valve closure, that is during the diastolic phase of the cardiac cycle, is about 10–12 kPa for the aortic valve and 1.5 kPa for the pulmonary valve. As tissue engineered Aorta

Commissure

Sinus

Commissure

Aortic root

Sinus

Free edge

Aortic ring (a)

Leaflets (b)

(c)

10.1 Schematic representation of a human aortic heart valve including the relative dimensions with respect to valve diameter: (a) side view of the complete valve, (b) after removal of one leaflet and the corresponding sinus and (c) view from the aortic side (adapted from De Hart, 2002).



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valves are living structures, which allow for growth and remodelling, they represent attractive valve substitutes for paediatric patients. Therefore, it is important to gain an insight into the haemodynamic parameters that apply to the valve and its diameter during development from neonate to adult. The aortic valve diameter increases from 9 mm at birth to 23 mm in adults. The cardiac output increases from 0.75 l min–1 at birth to 5.5 l min–1 in adults, with heart rate decreasing from 145 beats per minute at birth to 70 beats per minute in adults. Finally, the transvalvular pressure at diastole increases from 5 kPa at 1 year old to 10 kPa in adults. Figure 10.2 summarises these haemodynamic parameters and the valve diameter from neonate to adults, which have been collected from several databases, including The National Heart Lung and Blood Institute, MedScape and literature (Capps et al., 2000).

10.3.2 Structure and composition of semilunar heart valves Fibrous load-bearing tissues, such as heart valves, can be considered to be composite materials, mainly consisting of cells and extracellular matrix, organised to perform a specialised function. Mechanical functioning is 160

Valve diameter (mm) Transvalvular pressure (kPa)

140

Valve diameter (mm) Transvalvular pressure (kPa) Cardiac output (l min–1)

Cardiac output (l min–1) 20

Heart rate (beats per minute)

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25

120 100

15

80 10

60 40

5 20 0

0 0

5

10 Age (years)

15

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10.2 Haemodynamic parameters that apply on the aortic valve (transvalvular pressure difference, cardiac output and heart rate) and its diameter during development from neonate to adult. The data for this graph was obtained from several databases, including the National Heart Lung and Blood Institute, MedScape and literature (Capps et al., 2000).



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determined by the organisation and composition of the extracellular matrix (ECM) as well as the interactions between the matrix and the cells. The ECM is composed of a network of fibrous proteins, predominantly collagen and elastin, embedded in a gel of proteoglycans, glycoproteins and water. It provides mechanical support and anchorage for the cells and guides and restricts mechanical deformation of the tissue. The collagen fibres provide tensile strength to the tissue, whereas the elastin fibres give the tissue its resilience. The hydrated gel provides resistance when the tissue is compressed. These tissues are rich in collagen and their strength mainly depends on collagen architecture, that is collagen fibre length, fibre diameter, fibre organisation (e.g. alignment) and fibre cross-linking (Christiansen et al., 2000; Balguid et al., 2007). The main load-bearing parts of the valves are the leaflets. At first sight, these may look like simple structures, but they are complex and mechanically very robust tissues, serving a life-dependent function. Microscopically, the leaflets consist of three function-tailored tissue layers, maintained by the valvular interstitial cells, covered by endothelium. The fibrosa layer at the arterial side of the leaflets is mainly composed of circumferentially aligned collagen fibres. This layer is considered to be the main load-bearing layer of the leaflet and prevents excessive stretching during diastole. The ventricularis layer at the outflow surface of the leaflets mainly consists of radially aligned elastin. The elastin in this layer is believed to restore the collagen fibre geometry to its original corrugated geometry during systole allowing it to be stretched again in the next heart cycle. In between the fibrosa and the ventricularis, a loose spongiosa layer is present, containing an abundant amount of proteoglycans. This layer is suggested to serve as a shock absorber upon closure of the valve and enables sliding of the individual layers for optimal functioning. The typical layering of the aortic valve leaflet and the collagen fibre structure in the fibrosa of the aortic valve leaflet is demonstrated in Fig. 10.3.

10.3.3 Cell–matrix interactions in semilunar heart valves During development, heart valve formation goes through coordinated sequences of growth, differentiation and organisation that are orchestrated by spatial and temporal gradients of multiple regulatory factors. Accumulating evidence suggests a distinct role for mechanical and biochemical regulatory factors provided by the extracellular matrix environment (Schroeder et al., 2003; Schoen, 2008). The composition, architecture and mechanical properties of the extracellular matrix act in concert to provide the necessary cues to regulate tissue development. The sensors and effectors in this process are the valvular cells. During early tissue formation they respond to environmental stimuli by growing and dividing (proliferation) and by



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Cusp free edge

Free edge

Aorta

c

c

Fibrosa

209

Spongiosa Ventricularis

(a)

(b)

10.3 (a) The configuration of the fibrosa, spongiosa and ventricularis within the aortic valve leaflet (adapted from Mol et al., 2004). (b) Typical collagen fibre structure in an aortic valve leaflet (adapted from Sauren, 1981). The commissures are denoted by ‘c’.

laying down extracellular matrix components (Butcher and Markwald, 2007). Later, they respond by changing their own morphology and function (differentiation) and by modifying the composition and organisation of the extracellular matrix, a process referred to as matrix remodelling (Hinton et al., 2006). Thus, the cells are the key modulators of tissue formation and remodelling and in this way actively control and maintain tissue architecture and mechanical function. Under non-pathological conditions heart valves show an intriguing adaptive response to changes in their environment. In the case where the functional demand changes, the cells rapidly remodel the ECM to meet the new requirements. They presumably do this by changing the production of matrix components, by secreting matrix enhancing or degrading products (matrix metalloproteinases, MMPs) as well as their inhibitors (tissue inhibitors of metalloproteinases, TIMPs), and by applying traction forces to the deposited fibres (Baaijens et al., 2010). The interactions between the cells and the matrix environment are twofold. On the one hand, the matrix controls cell and tissue fate (cell shape, proliferation, differentiation, motility, alignment) through its 3D architecture and focal adhesion organisation. On the other hand, the cells influence the matrix by applying traction forces and by changing the matrix turnover. How the cells sense and affect their environment is only partly known. Several ‘mechanosensors’ and receptors for biochemical cues have been identified (Katsumi et al., 2004; Parker and Ingber, 2007; Kolahi and Mofrad, 2010) and various signalling pathways that translate the sensed signals into cellular actions have been proposed (Tomasek et al., 2002; Chiquet et al., 2009), though not specifically for heart valves. A thorough understanding of the interactions between cells and their microscopic matrix environment in



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developing and remodelling tissues, however, is lacking. It is nevertheless evident that scaffold materials are an important tool for mimicking these interactions during tissue engineering strategies and to study systematically these interactions during in vitro and in vivo experiments. On the microscopic scale a diversity of environmental factors contributes to the overall control of the cell fate, either through physical or through molecular interactions, aimed at guiding cell morphology and behaviour (adhesion, migration, proliferation, differentiation) as well as tissue morphogenesis. These environmental factors, which together constitute the so-called ‘cellular niche’, include the presence and density of cell adhesion molecules, growth factors and morphogens, local viscoelastic properties, local fibrillar matrix structure (i.e. fibre diameter and morphology, fibre interconnections and mutual organisation) and local mechanical stresses and strains. While many of these ECM related properties might be incorporated in scaffolds for heart valve tissue engineering, stresses and strains cannot. The scaffold material should, however, be designed to facilitate the application of such loads, for instance by providing elastic environments that allow for ongoing cyclic loading of the cells and neo-tissue. Cell adhesion plays a crucial role in regulating cell motility, proliferation and differentiation. Since the identification of the Arg-Gly-Asp (RGD) tripeptide of fibronectin and other adhesion molecules (Ruoslahti and Pierschbacher, 1987), several approaches for reconstituting the cell-adhesive character of the ECM in 3D scaffolds have been developed. In cardiovascular tissue engineering, Arg-Glu-Asp-Val (REDV) and Val-Ala-Pro-Gly (VAPG) peptide sequences have been used selectively to capture endothelial cells and smooth muscle cells in heterogeneous cell suspensions under shear stress (Plouffe et al., 2007). Whereas overall cell adhesion in these studies was dominated by the shear stress, the selective adhesion could be better controlled by the type of peptide. Apart from the type of peptide, the surface density and spatial distribution of peptides can be used to control cell fate (Maheshwari et al., 2000). In addition to synthetic materials, it is further possible to incorporate cell adhesion ligands into biologically derived materials to alter their adhesive character. This has been explored for decellularised heart valves in an attempt to improve the cell adhesive properties of these scaffolds (Shi et al., 2009). The micromechanical environment of the cell is another important regulator of cell function and viability, either through the application of external mechanical loads on cells, or through the development of cell contractile forces on materials with different mechanical properties. Two-dimensional (2D) cell culture studies on substrates with varying elastic moduli have demonstrated that (stem) cell differentiation can be influenced by the mechanical properties of the cell’s microenvironment (Engler et al., 2006; Nemir and West, 2010). The ability to reproduce these properties in a synthetic 3D scaffold may



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provide a potent means of controlling cell fate both in vitro and in situ. Kloxin et al. (2010) used a photodegradable hydrogel to study the effects of a wide range of elastic moduli on the differentiation of porcine valvular interstitial cells (VICs) into myofibroblasts. VICs cultured on high moduli differentiated into matrix producing myofibroblasts, whereas decreasing the substrate modulus suppressed this differentiation, demonstrating that myofibroblasts can be de-differentiated solely by changing the modulus of the underlying substrate. These findings are important for the rational design of biomaterials for heart valve regeneration and offer insight into fibrotic tissue development and counteractive therapy by changing the elasticity of the microenvironment.

10.4

Selection of scaffold material

In general, there a two groups of polymer materials that can be considered for use in electrospinning to obtain a valve scaffold: the naturally derived and the synthetic polymers. Natural polymers such as polysaccharides, gelatin, collagen (Matthews et al., 2002), elastin (Buttafoco et al., 2006), silk (Soffer et al., 2008) and fibrinogen (McManus et al., 2007) were electrospun into scaffolds for cardiovascular research. In contrast to synthetic polymers, these materials provide many of the natural instructive cues for cell attachment and proliferation. They are considered to be biodegradable, although biosafety should be ensured prior to their use. On the other hand, the natural polymers demonstrate batch-to-batch variation and the electrospinning process is less versatile since there is a limited choice of solvents which do not compromise the integrity of the polymers (Zhang et al., 2005; Zeugolis et al., 2008). This limited choice of solvents also restricts control over the mechanical properties and the design and biodegradability of the resulting scaffold. In contrast to these natural polymers many synthetic polymers can be dissolved and electrospun in a broader range of solvents as well as directly from their molten state. This offers more possibilities and design freedom for obtaining a scaffold tailored with the desired tissue requirements. The most commonly known and used synthetic polymers in tissue engineering are the poly(a-hydroxy acids), such as the lactic and glycolic acids as well as PCL and their copolymers (Piskin et al., 2007). In addition, there are manmade polymeric systems, such as polyurethanes (PU) (Saad et al., 1999; Courtney et al., 2006), polyphosphazenes (Carampin et al., 2007) or supramolecular polymers (Dankers et al., 2005), specifically designed to meet the desired mechanical properties, cell responses and integration of functionalities, such as peptides or growth factors. The poly(a-hydroxy acids), with poly(lactic acid-co-glycolic acid) (PLGA) being the most investigated polymer owing to its variable degradability (Vert et al., 1992), and PCL (Zong et al., 2005) are representatives of biodegradable



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polymers. They degrade mainly by hydrolysis in a few weeks to several months, depending on the molecular structure, molecular weight, fibre morphology, and so on (Dong et al., 2009). It is evident that the degradation products of the biodegradable polymers must be non-toxic and do not provoke any foreign body response (Böstman and Pihlajamäki, 2000).

10.5

Scaffold properties to meet tissue requirements

Electrospinning, which allows the creation of fibrous scaffolds with control over 3D scaffold architecture, is exploited for cardiovascular tissue engineering and heart valve tissue engineering in particular. Once familiar with the tissue requirements for the human semilunar heart valves, different electrospinning settings and modalities can be applied to meet these requirements on the macro- to microscopic level. These include the valve geometry, degradation, fibre alignment, diameter and morphology, and bioactivity.

10.5.1 Scaffold geometry Electrospinning allows direct fabrication of complex 3D structures, such as heart valve scaffolds. By selecting and manufacturing the appropriate shape and dimensions of the mould used for electrospinning, native valve geometries from various materials, such as PCL (Van Lieshout et al., 2006; Del Gaudio et al., 2008) and PU, can be obtained (Fig. 10.4).

10.5.2 Balancing scaffold degradation and tissue formation During the process of tissue engineering, the scaffold provides initial anchorage and support for the cells, until they have produced and organised their own ECM to obtain a functional tissue. A common hypothesis is that the rate of

1 cm (a)

1 cm (b)

1 cm (c)

10.4 Three-dimensional electrospun heart valve scaffolds made from polyurethane (Saad et al., 1999): (a) view from the aortic side, (b) view from the ventricle side and (c) view of the complete valve.



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scaffold degradation, and with it the loss of mechanical strength, should be in pace with the rate of matrix production. In addition, the engineered substitute should meet and maintain identified mechanical target properties upon implantation to resist in vivo haemodynamic loading. Subsequent remodelling is intended to recapitulate normal tissue architecture and function. Thus, degradation properties and concomitant changes in mechanical properties need to be incorporated into the scaffold design. So far mainly rapidly degrading synthetic scaffolds, such as PGA, have been used for in vitro tissue engineering. In such PGA scaffolds, ECM formation and tissue mechanical properties can be stimulated within weeks depending on culture conditions (Rubbens et al., 2009). However, these scaffolds are readily seeded with matrix producing cells, while this is not the case for more advanced in situ tissue engineering solutions. For the in situ approach, scaffolds should be designed to mimic the mechanical behaviour of native heart valves, although these may change between subjects and with age (Stephens et al., 2010). Only upon recruitment and/or bioactive stimulation of matrix synthesising cells can ECM be formed, requiring a slow and controlled degradation of the scaffold material to balance neo-tissue formation and to ensure haemodynamic functioning. Because of patient-specific variations in neo-tissue formation, scaffold degradation should be tightly controlled, preferably by host cells using, for example protease-mediated degradation (Davis et al., 2005) or by external stimuli for on-demand degradation (Chan and Mooney, 2008). Owing to the flexibility of the electrospinning process it would also be possible to co-electrospin a degradation trigger directly into the polymer scaffold (Stuart et al., 2010). This could allow, for example, an external induced stepwise degradation based on the actual rather than on the expected progress of the growing tissue. Controlled degradation might be possible by peptides built into the backbone of the polymers (Mann et al., 2001). Regulating the degradation rate of scaffold materials is furthermore important for the release of incorporated bioactive factors that direct tissue formation and may facilitate cell migration, tissue remodelling and the infiltration of blood vessels (Hubbell, 2003). One of the consequences of using slowly degrading scaffold materials, however, is cellular adaptation from a synthetic to a quiescent phenotype; in other words the cells become ‘lazy’ while sensing the presence of a supporting synthetic environment. Alternatively, the cells may attain contractile properties, resulting in scaffold compaction caused by traction forces (Tranquillo et al., 1992), necessitating mechanically robust cellular microenvironments to counteract these forces.

10.5.3 Scaffold fibre alignment, diameter and morphology The fibre alignment within the scaffold is important to obtain a scaffold that could mimic the fibres in the ECM, such as the collagen structure in the aortic



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heart valve leaflet. Fibres can guide cell growth and orientation (Riboldi et al., 2005; Stevens and George, 2005). Furthermore the fibre alignment in the scaffold directly influences its mechanical properties and functional behaviour (Courtney et al., 2006). Several electrospinning techniques are available to adjust and control fibre deposition and orientation. A simple method of achieving a high degree of oriented fibres is spinning towards two parallel arranged flat electrodes (Dersch et al., 2003). When expanding this method using more parallel arranged electrodes, an oriented cross- or star-shaped deposition can be achieved (Li et al., 2004). A more commonly used method is changing the fibre collection speed using a rotating target (Theron et al., 2001; Matthews et al., 2002). With this method more accurate control of fibre orientation is possible, allowing the creation of scaffolds that closely mimic the mechanical anisotropy of a heart valve leaflet (Courtney et al., 2006). Most tissue engineering studies depend on the reproducibility and control of the fibre geometry that electrospun architectures promise to provide. In the past, this has led to significant efforts in the field of electrospinning focusing on elucidating the influence of processing parameters on both the fibre diameter as well as the morphology. The properties of the polymer solution have a large influence in the electrospinning process and on the resulting fibre diameter and morphology. So the viscosity of the solution and its electrical properties will define the elongation of the jet, thereby influencing the diameter of the resulting electrospun fibre, whereas surface tension is crucial in the process for spinning either beaded or straight fibres. With increasing viscosity, the diameter of the electrospun fibre increases (Baumgarten, 1971; Deitzel et al., 2001; Demir et al., 2002). This may be attributed by the increased resistance of the solution to stretching in the jet (Jarusuwannapoom et al., 2005). Increased fibre diameter with a higher concentration of the polymer solution is attributed to the jet instability that occurs at increasing distance from the needle. As a result, the jet path is reduced, resulting in less stretching of the polymer solution in the jet. With increased conductivity of the solution, more charge can be carried by the jet. The conductivity can be increased by adding salt or polyelectrolyte or selecting solvents with a higher conductivity. Since the stretching of the jet is caused by repulsion of the charges at the surface, a higher conductivity will increase the stretching rate of the solution in the jet. As a result, beading in fibres is suppressed. In addition, the fibre diameter is reduced (Zong et al., 2002) as the higher conductivity causes the jet instability to shift towards the needle. This results in the opposite effect to higher viscosity; the jet path is increased, more stretching occurs and the fibres are spread over a bigger area on the target (Choi et al., 2004). The same occurs with a higher dielectric constant of the solution, resulting in a reduced fibre diameter and



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suppression of bead formation (Lee et al., 2003; Hsu and Shivkumar, 2004; Son et al., 2004). The fibre diameter can further be adjusted by external factors. Increasing the polymer feeding rate directly increases the diameter. Increasing the field strength, either by varying the needle to target distance or the applied voltage, creates greater columbic forces in the jet, resulting in increased stretching of the solution in the jet and thus, smaller fibre diameters (Buchko et al., 1999; Lee et al., 2004). This process occurs up to a point when the higher voltage increases the acceleration of the electrospinning jet, resulting in decreased flight time and hence, thicker fibres (Zhao et al., 2004). By using the various processing parameters described here, fibre diameters from ~4 nm up to about 20 mm can be electrospun. This gives ways of mimicking the heart valve ECM in a similar level of complexity. Micrometre-sized fibres give structure and mechanical stability and can mimic the collagen structure in the heart valve leaflet. Nanometre fibres, on the other hand, have a much bigger surface area for absorbing proteins and offer more binding sites to the receptors of cells. It is known that cells sense and respond to different length scales. On micrometre fibres, they attach in a similar fashion as on flat surfaces. The cells flatten and follow the orientation of the fibre, whereas a cell can bind to several nanofibres simultaneously and even spread in-between individual fibres of multiple nanofibres (Stevens and George, 2005). Therefore, to mimic the heart valve ECM structure more closely, multiple fibre diameters need to be integrated into one scaffold. Recently developed methods allow electrospinning of various controlled fibre diameters simultaneously and in an intermingled matter (Gentsch et al., 2010; Soliman et al., 2010). Next to fibre diameter, porosity (or void space) is important as this favours cell ingrowth, cell proliferation, cell motility and matrix synthesis, which are relevant for heart valve tissue formation. The addition of fibres with large diameters into nanofibrous scaffolds increases the pore size, since nanofibrous scaffolds lack pores and void spaces large enough for cells to grow in, resulting in a confluent cell layer on top of the scaffold (Lannutti et al., 2007). A theoretical model of the porosity in such electrospun architectures illustrates that the mean pore radius in an electrospun scaffold directly depends on fibre diameters since the architectures produced by electrospinning are based on stacks of layers of 2D sheets (Eichhorn and Sampson, 2005). Incorporating fibres of several micrometres directly increases pore and void spaces. A study on PCL meshes showed an optimum fibre diameter of 12 mm for homogenous 3D cell infiltration with human venous myofibroblasts (Balguid et al., 2009), a cell source commonly used for tissue engineering of heart valves. Natural nanofibres can be included into these meshes by using a fibrin gel as a cell carrier (Mol et al., 2005), representing an intriguing way of producing bimodal fibrous scaffolds, thereby allowing the



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cells to spread homogeneously along the fibres as well as within the pores. Scaffolds with multimodal fibre diameters are only one way to improve the generally limited cellular infiltration. Spinning simultaneous or sequential different materials followed by selectively dissolving one material (Pham et al., 2006; Holzmeister et al., 2007) or using salt leaching methods (Nam et al., 2007) are alternative approaches. These methods include additional steps and processes, which make the mesh production more complicated. Furthermore, these methods have the drawback that they heavily rely on the dependency of pore size on fibre diameter. A promising method independent of fibre diameter is low temperature electrospinning, where embedded ice crystals from the surrounding air act as a pore template resulting in a large increase in porosity with constant fibre diameter (Simonet et al., 2007). The first tissue engineered applications using such meshes showed complete and homogenous cell ingrowth (Henry et al., 2009; Leong et al., 2009). Next to controlling fibre alignment and diameter, it is also possible to structure the surface of the fibre itself. In general, electrospun fibres are smooth and have a circular cross-section (Fig. 10.5(a)). Flat (Fig. 10.5(b)) or ribbon shapes are possible, in particular when working with highly concentrated solutions (Ramakrishna et al., 2005). Inducing phase separations during the electrospinning process allows adaptation of the fibre surface topography, from smooth to porous. Selecting specific solvents or solvent mixtures (Bognitzki et al., 2001) can create polymer-rich and -poor regions, thereby forming pores within the fibres (Fig. 10.5(c)). Other approaches with similar topographical effects are electrospinning in a very humid environment (Megelski et al., 2002), electrospinning into liquid nitrogen (McCann et al., 2006) or by spinning immiscible polymers from the same solvent (You et al., 2006). Controlling the pores in the fibres might be of great interest as these pores could act as anchor points and affect cell behaviour. Various studies showed that surface topography, smooth as well as rough, can play a

(a)

(b)

(c)

15 µm

10.5 Scanning electron microscopy (SEM) pictures of PLA fibres electrospun with different parameters resulting in: (a) round fibres with a smooth surface, (b) flat fibres or (c) round fibres with a porous surface.



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prominent role in cell behaviour (Rea et al., 2004; Xu et al., 2004; Bakeine et al., 2009; Guillemette et al., 2009).

10.5.4 Bioactive scaffolds The biomimetic properties of the electrospun scaffolds can be improved by adding bioactive functions to the polymer. These functions can be chemically attached in a second labelling step with electrospun fibres (Casper et al., 2005; Ma et al., 2005) or directly co-electrospun with the synthetic polymer (Xie and Hsieh, 2003). The latter approach is easiest, but requires special attention with respect to the stability and activity of the biomolecules (Gunn and Zhang, 2010). In addition, the majority of the biomolecules might be embedded within the fibres and not on the surface, allowing their release over a prolonged period of time via diffusion or fibre degradation. Other reported ways to incorporate and add biofunctionality are, for example, by producing core-shell fibres or embedding the biomolecules in hollow or porous microspheres. This protects the biomolecules from the solvent during the electrospinning process and allows triggered release by external or internal factors (Stuart et al., 2010). Another promising option is to incorporate directly the biomolecules into the side chains or the backbone of the polymer (Dankers et al., 2005). As soon as relevant biomolecules are identified for heart valve tissue engineering, these can be incorporated into the heart valve scaffold using one of the described techniques.

10.6

Future trends

Most of the tissue requirements defined for heart valve tissue engineering can be obtained by electrospinning methods. Combining the various electrospinning methods and techniques is challenging, since they are not always compatible. For example, the addition of bioactive factors often reduces the electrospinning possibilities and results in less freedom in scaffold design. Nevertheless, to obtain a functional heart valve, these methods need to be integrated into one complex 3D heart valve scaffold, containing the three-layered structure. Each of these layers should ideally include tailored bioactivity and mechanical properties. Efforts are being undertaken to identify the best combination and concentration of the various bioactive factors and their implementation into the electrospun structure at the dedicated locations. In particular the fatigue and creep of the scaffold structures should be studied in more detail for in situ heart valve tissue engineering. To deal with patient to patient variations in neo-tissue formation, degradation should be controlled externally or triggered directly by the amount of cells and the produced ECM. Furthermore, builtin contrast agents allow monitoring of scaffold degradation and function of the heart valve in the patient after implantation. Only a multidisciplinary



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approach will finally result in a functional 3D valve scaffold, which will inspire the cellular microenvironment and guide cell growth, differentiation and functional tissue organisation towards an ideal heart valve replacement that will allow growth and remodelling.

10.7

Acknowledgment

This research forms part of the Project P1.01 iValve of the research programme of the BioMedical Materials institute, co-funded by the Dutch Ministry of Economic Affairs. The financial contribution of the Nederlandse Hartstichting is gratefully acknowledged.

10.8

References

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Bladder tissue regeneration

S. C. B a k e r and J. S o u t h g a t e, The University of York, UK

Abstract: Organ donor shortages have driven the search for tissue engineering and regenerative solutions in many tissue pathologies, but disorders of the urinary bladder have no allograft solution. Currently, end-stage bladder disease is addressed by reconstructing the bladder with vascularised segments of bowel, but serious long-term side-effects correspond to a clinical need to develop new approaches. The application of electrospun biomaterials, in combination with in vitro tissue engineering or in vivo regenerative engineering strategies, holds elusive promise for routine bladder augmentation or substitution. The major challenges include identification of suitable elastic/biodegradable polymers and scaffold architectures that match the physical and biological requirements of the bladder. Key words: bladder biomaterial, bladder regeneration, electrospun polystyrene, urinary tract tissue engineering, urothelium.

11.1

Structural/functional properties of the bladder

The ability of the bladder to both accommodate the storage of large (400– 500 ml) volumes of urine at low pressure and facilitate controlled voiding reflects the highly adapted structural and biological evolution of the bladder. The normal human bladder is composed of three main layers, comprising an inner urothelial lining supported on a basement membrane with an underlying collagen-rich, vascularised lamina propria and surrounded by interlaced layers of highly compliant detrusor smooth muscle (Fig. 11.1). Together, the biological and biomechanical properties of these individual tissue components contribute to the way the bladder functions as a compliant, distensible organ (reviewed by Korossis et al., 2006; Andersson and Arner, 2004). As the bladder fills, its contents are maintained by a striated (voluntary) and smooth muscle (autonomic) set of sphincter muscles around the bladder opening, the urethra. During filling, the sphincter muscles are contracted, keeping the urethra closed. Once full, the voluntary initiation of voiding triggers simultaneous relaxation of the sphincter and contraction of the detrusor by a complex interplay of the somatic and autonomic nerve system, which finally leads to a sharp rise in pressure. The precise nature of the sensor that triggers the voiding reflex is unknown, although mechanisms involving mechanosensitive receptors in the urothelium have been proposed (de Groat, 2006). 225 © Woodhead Publishing Limited, 2011


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Electrospinning for tissue regeneration Lumen Urothelium Basement membrane Lamina propria

Blood vessel

Stroma Muscle bundle

11.1 Haematoxylin- and eosin-stained section of the urinary bladder showing the multilayered urothelium, the lamina propria and the detrusor smooth muscle bundles (Scale bar = 500 µm).

During filling, the bladder distends to several times its starting volume and during voiding returns rapidly to its original size and shape. The distension is achieved through the extensive uncoiling of collagen fibres in the lamina propria, in conjunction with relaxation of the detrusor muscle, with the lamina propria acting as the capacitance layer and the detrusor functioning to prevent overdistension of the bladder wall (Chang et al., 1999). The postdeformation recovery of the collagen fibrils is attributed to the recoiling action of distensible elastin during successive cycles.

11.1.1 The urothelium The urothelium is a transitional epithelium, classified as such because its properties lie between stratified squamous and simple non-stratified epithelia. Although urothelium is mitotically quiescent, with a very low constitutive rate of cell turnover, it has a high regenerative capacity in response to damage (Varley et al., 2005). This regenerative property, which is important for conserving barrier function, also makes it an ideal target for tissue engineering



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and regenerative strategies. As discussed below, the urothelium not only plays an essential role in providing a urine-proof lining, but also plays an important role in tissue accommodation by contributing to the changes in luminal surface area. Histologically, the urothelium is stratified into basal and superficial cell zones, interposed by three to six intermediate cell layers that vary according to the degree of distension of the bladder. The superficial cells are highly specialised and provide the primary urinary barrier through expression of urothelium-restricted uroplakin proteins in the apical membrane (Hu et al., 2000; Olsburgh et al., 2003) and well-developed intercellular tight junctions (Varley et al., 2006). These molecular features contribute to transcellular and paracellular barrier functions, with the urothelium recognised as the ‘tightest’ epithelium in the body (Acharya et al., 2004, Slobodov et al., 2004). The uroplakins also contribute to a vesicular mechanism of apical membrane insertion and turnover that regulates changes in luminal surface area during bladder accommodation (Truschel et al., 2002).

11.2

Bladder disease and the need for bladder substitution

The need for bladder augmentation or reconstruction may arise as the consequence of an array of congenital and acquired conditions. With some 71 000 newly diagnosed cases of bladder cancer in the United States in 2009 (National Cancer Institute, 2010), muscle-invasive and recurrent highgrade bladder cancer is the most common indication for radical cystectomy, requiring major bladder reconstruction or replacement. Patients with voiding dysfunction that is due to therapy–refractory neuropathic bladder, or a low compliance bladder resulting from chronic inflammatory conditions may also require surgical intervention, as an abnormal increase in bladder pressure can result in vesicoureteric reflux and chronic renal failure. Reconstructive therapy, such as augmentation cystoplasty may be indicated in these patients to restore a low-pressure system within the urinary tract. Currently, the most commonly performed procedure to replace or augment the bladder involves surgical incorporation of a vascularised tissue graft, most usually a segment of ileum. This procedure of ‘enterocystoplasty’ is successful in a large proportion of patients, as it both improves continence and prevents the risk of further damage to the kidneys. Nevertheless, it is associated with a number of serious urinary diversion-related complications and a high perioperative morbidity (reviewed by Turner and Southgate, 2009). A significant number of patients undergoing cystectomy suffer from early complications of the surgery (Shabsigh et al., 2009; Froehner et al., 2009). Bowel tissue is unsuitable for long-term exposure to urine, as it can react by intense mucus production that acts as a nidus for stone formation



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and bacteriuria. In addition, metabolic disorders can occur owing to the absorptive nature of the bowel epithelium and, in the long term, there is an emergent risk of cancer (McDougal, 1992; Gough, 2001). In developing new tissue engineering or regenerative approaches, the clinical need is clear: to find an alternative to enterocystoplasty that delivers equivalent capacity and compliance in storing urine at safe pressures, but without the associated problems. The range of different tissue engineering/ regenerative strategies have been reviewed elsewhere (Bolland and Southgate, 2008; Turner and Southgate, 2009), but essentially involve either an adaptation of the surgical use of bowel to reconstruct the bladder in which the mucus-secreting absorptive bowel epithelium is substituted by in vitrogenerated autologous urothelium (‘composite cystoplasty’) (Fraser et al., 2004; Turner et al., 2011), or the use of biomaterials as a scaffold for full thickness bladder wall replacement. The latter strategy may use a bladder tissue-engineering approach where a scaffold is seeded with cells in vitro prior to transplantation, as described by Atala et al. (2006), or rely on cell recruitment and functional tissue integration in vivo. Clearly, the identification of a biomaterial that recruits a bladder tissue regenerative response in vivo is the most attractive idea, but to date this has not yet been achieved and when assessed in long-term large animal studies using clinically relevant graft sizes, there is a tendency for graft contraction (Brown et al., 2002). This highlights the need to develop more appropriate natural or synthetic biomaterials for the bladder.

11.3

Electrospun and other scaffolds for bladder tissue engineering

A general observation is that cells isolated from tissues and grown as adherent cultures commonly lose expression of differentiated characteristics. Regarding the bladder and associated urinary tract, a loss of differentiation is seen in both urothelial (Lobban et al., 1998, Southgate et al., 1994) and stromal/ smooth muscle cells (Baker and Southgate, 2008) in monolayer culture. Whereas the urothelium is a self-regenerating epithelial lining tissue that does not depend on a matrix to self-organise and develop a differentiated phenotype (Cross et al., 2005), the cells of the bladder wall are embedded within a collagenous matrix that is generated first during development and provides both structural and biological support. The development of biomaterial scaffolds is based on the premise that promoting cell:cell and cell:matrix interactions in three dimensions will enhance functional tissue biology. Although many scaffolds present a threedimensional (3D) appearance macroscopically, in terms of their surface area:volume ratios, they effectively deliver two-dimensional (2D) properties when viewed on the micro- or nanoscale (Fig. 11.2). Foams in particular, but



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(a)

(b)

11.2 (a) Scanning electron micrograph of urothelial cells on a dense woven polyglactin mesh and (b) micrograph of a phalloidin-FITC stained stromal cell in a highly porous electrospun polystyrene scaffold of lower diameter fibres (Scale bar = 50 µm).

also scaffolds composed of large diameter fibres, have open cellular-scale surface areas with which cells interact preferentially as if they were still in a culture flask, so that cell:biomaterial interactions dominate over desired cell:cell interactions. The major potential of electrospinning is to create sub-cellular scale fibres that mimic the extracellular matrix: for example, the collagen fibrils of different animals and different tissues assemble to different diameters but tend to range from 10–600 nm (Parry et al., 1978), creating a scaffold framework from which cell:cell interactions can emerge. The development of electrospun scaffolds can be classed into two types: those that use synthetically generated polymers versus those based on natural or biologically derived compounds.



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11.3.1 Synthetic polymer scaffolds Synthetic polymers have a number of advantages including low price, homogeneity of raw materials and avoiding the potential problems of using animal-derived products. In cell biology, polystyrene is ubiquitous as a twodimensional culture substrate and we previously reported the development of plasma-treated electrospun polystyrene to allow direct comparison between cell cultures grown on 2D versus 3D substrates (Baker and Southgate, 2008). The process of electrospinning affords tight control over scaffold architecture, enabling the effects of variables such as porosity and fibre alignment to be studied without confounding the study by changing polymers and the surface chemistry (Baker et al., 2006). In a proof-of-principle study, an electrospun polystyrene scaffold seeded with stromal cells (Baker and Southgate, 2008) was used to demonstrate the potential for creating a full thickness bladder tissue model in vitro (unpublished observations). An autologous, differentiated human urothelial cell sheet was generated (Cross et al., 2005), harvested by separation from its substrate using dispase II and transferred onto the stromal scaffold (Turner et al., 2011). The combined tissue was maintained at an air–liquid interface in culture for 12 days, before fixation and processing for immunohistology (Fig. 11.3). Whereas there are many advantages to using synthetic polymers, there are also other issues to consider. A polystyrene scaffold provides a permanent, non-biological structural support framework suitable for in vitro studies, but does not itself substitute for the dynamic, biologically interactive extracellular matrix. A further issue with synthetic polymer materials is their tendency

(a)

(b)

11.3 Development of an in vitro bladder tissue construct. (a) Haematoxylin- and eosin-stained section illustrating the maintenance of epithelial and stromal compartments within the integrated structure of the neo-tissue. (b) Immunoperoxidase labelling for human collagen type IV showing the deposition of collagen within the stroma. Note the intense deposition of collagen type IV at the interface of the urothelial and stromal cell compartments, marking the de novo basement membrane. The presence of visible electrospun scaffold fibres within the stroma is indicated by an arrow (Scale bar = 50 µm).



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for non-specific surface adsorption of proteins. Although such a feature may provide positive benefits, for example, the adsorption of serum proteins may facilitate cellular adherence to a polymer surface, nevertheless, this fouling of surfaces by serum proteins may dominate and serve to mask specific features of the biomaterial topography or chemistry. We have shown in in vitro studies that the phenotype of stromal cells on electrospun polystyrene scaffolds may be confounded by the adsorption of proteins, such as serum albumin, from the culture medium (Baker and Southgate, 2008). Albumin represents a major serum component, where it acts as a carrier for a wide range of bioactive factors. Thus, when albumin is adsorbed onto a scaffold surface, associated growth factors may play a significant role in modulating cell phenotype. In our study, we showed that whereas smooth muscle cell differentiation was promoted in the absence of serum by exogenous TGFb1 in 2D culture, this effect was masked in equivalent 3D scaffolds, where serum proteins used to promote initial cell adherence remained adsorbed on the scaffold, even following withdrawal of serum from the medium (Baker and Southgate, 2008). These observations highlight the difficulties in defining clear relationships between biomaterial and biological properties, as required for the rationalisation of biomaterial design. One proposed approach to introduce better control over synthetic biomaterial properties is surface modification by PEGylation. In addition to preventing non-specific protein adsorption, this also provides a route to surface functionalisation by heparinisation, which can act as an ECM-mimetic route for the sequestration and presentation of cytokines (Rohman et al., 2009). A further hindrance to the progress of synthetic material development for tissue engineering has been a lack of investment in new polymer technologies. Companies involved in tissue engineering have been hesitant to develop new polymers for biomaterials owing to the time and expense associated with regulatory approval, preferring instead to pursue iterations of already approved polyesters (including poly-lactides, poly-glycolides and poly-caprolactones). This has been of particular detriment to bladder tissue engineering, where the remarkable compliance of the bladder wall is poorly matched by a lack of elasticity in the approved polyesters. There are several synthetic polymers which could approach the elasticity of the bladder extracellular matrix, whilst retaining suitable biodegradability and strength. So far however, none of these polymers, which can all deliver 500% strain or greater, have been employed in electrospun materials for urological tissue engineering. Examples include poly(ester)urethanes (Gorna and Gogolewski, 2000), poly(diol citrates) (Yang et al., 2006), poly(ethylene glycol)/poly(butylene terephthalate) (Deschamps et al., 2002; Webb et al., 2004) and poly(1,3-trimethylene carbonate)/d,l-lactide (Pego et al., 2003).



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11.3.2 Natural/biologically derived polymer scaffolds Perhaps the most biologically relevant natural materials developed to date are the decellularised porcine bladder materials (e.g. Bolland et al., 2007). These scaffolds are attractive because they retain tissue-specific architectures and, depending on processing, may retain sequestered growth or other bioactive factors. However, retention of mature tissue architecture does not necessarily provide the best environment for cell colonisation and retaining the cytokine profile of a mature tissue may not support the developmental processes required to create essentially new tissue (Turner and Southgate, 2009). To address the problem of restricted pore sizes found in natural matrices, one group lyophilised, comminuted and pepsin-digested the decellularised extracellular matrix extracted from porcine urinary bladder to allow electrospinning (Stankus et al., 2008). The bladder matrix was dissolved in hexafluoro-1,1,1,3,3,3-isopropanol (HFIP), at concentrations from 6–15 wt%; however, instabilities in the Taylor cone meant that stable fibre generation was not possible without adding poly(ester-urethane)urea as a more viscous stabilising polymer (Stankus et al., 2008). The instabilities observed when electrospinning extracellular matrix protein alone manifested as bead-on-astring aberrations, considered unsightly and often considered to have negative implications for subsequent application, although their contribution to cell biology has yet to be adequately qualified. Individual collagen isotypes have been successfully electrospun previously both singly and as copolymers (Matthews et al., 2002, 2003; Boland et al., 2004; Huang et al., 2001), but one approach to developing a more tissuespecific electrospun matrix is to extract and use collagens in their naturally occurring ratios. In unpublished work, we extracted collagens from fresh porcine bladder tissue, using a pepsin digestion and acetic acid solubilisation approach. This approach yielded soluble collagens, which are more applicable to electrospinning than digested decellularised bladder matrix which retains a greater degree of cross-linking. The porcine bladder collagen extract was solubilised by overnight stirring in HFIP at a concentration of 0.08 g ml–1, based on the method developed by Matthews et al., (2002). The collected electrospun mat was highly homogeneous, with a mean fibre diameter of 200 nm (Fig. 11.4). An unfortunate complication of the process was that the resultant scaffolds were highly soluble in cell growth medium and required cross-linking to stabilise them. Glutaraldehyde cross-linking (performed in 20% glutaraldehyde vapour over 24 hours) resulted in marked scaffold contraction and reduced elasticity, limiting the applicability of these scaffolds. The creation of natural electrospun collagen scaffolds remains an attractive idea to generate tissue-specific scaffolds, but requires significant further development to produce usable bladder matrix scaffolds. The development of electrospun elastin scaffolds was reported to have similar problems that



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5 KV

X2, 500

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16 MM

11.4 Scanning electron micrograph of electrospun scaffold generated from porcine bladder matrix extract using the following conditions: solution concentration, 0.08 g ml–1; applied voltage, 20 kV; flow rate, 0.03 ml min–1; working distance, 10 cm. (Scale bar = 10 µm).

required cross-linking to resolve, with glutaraldehyde cytotoxicity noted as a potential problem (McClure et al., 2008). Future work in this area might consider carbodiimide (Park et al., 2002) or genipin (Tsai et al., 2000) as alternative and less-toxic cross-linkers. Fibrinogen is a soluble blood protein involved in clotting and wound healing which could form a suitable provisional matrix for supporting the development of new tissues. A recent study described the evaluation of electrospun fibrinogen, which was not cross-linked, but cultures were supplemented with aprotinin to prevent rapid cellular breakdown (McManus et al., 2007). Although there were a number of limitations to the study, the results suggest that rather than being constrained on the surface, as is common with many biomaterials, human bladder smooth muscle cells (HBSMC) were able to infiltrate the matrix. The authors had previously assessed the mechanical properties of electrospun fibrinogen (110 mg ml–1 in HFIP) and found a favourable increase in strain at failure from ~22% to ~120% when the scaffolds were wetted with serum-containing cell culture medium (McManus et al., 2006); this compares reasonably with strain values for porcine bladder tissue reported separately (Baker et al., 2006). However, as noted above (Baker and Southgate, 2008), the presence of serum may have negative effects on differentiated HBSMC phenotype, which were not explored in the McManus study.



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11.4


Electrospinning fit for purpose

To date, electrospinning has not made the contribution to urology that it has in cardiovascular engineering or wound management and it seems most likely that this is due to the mechanical properties of electrospun polymers being a poor match for bladder tissue. However, there is a growing body of work suggesting electrospun biomaterials could in the future make a significant contribution to urinary tract reconstruction.

11.4.1 Control over architecture One of the great attractions of electrospinning is its ability to offer fine control over the fibre diameter and scaffold architecture generated, with at least two predictive models described (Fridrikh et al., 2003; Baker et al., 2006). In fact the level of architectural control afforded by electrospinning is not currently matched by a clear understanding of the scaffold requirements of tissues for regeneration. Relatively few studies have systematically compared different architectures in controlled biological studies and as such, most scaffold specifications remain based on assumed requirements, rather than supported by objective evidence. For instance, the approach taken by some researchers has been to mimic the mature tissue architecture directly when generating electrospun scaffolds. In a study using electrospun cellulose acetate, a multilayered scaffold was developed using a cross-sectional scanning electron micrograph of urinary bladder matrix as the blueprint, which was described as comprising three distinct regions designated fibrous, cellular and dense (Han and Gouma, 2006). The study demonstrated the versatility of electrospinning in generating three architecturally distinct cellulose acetate scaffolds by creating a three-layered composite tissue mimetic (Han and Gouma, 2006). However, it might be argued that rather than recreating mature tissue architectures, the need is to create developmental frameworks within which cells mature and develop their own matrix. Furthermore, the creation of a dense fibrous layer (with significant aberration content) at the base of the scaffold (Han and Gouma, 2006) appears likely to interfere with the processes of cell infiltration and angiogenesis. Unfortunately, no cell biological analysis was performed and as a result, the differential contribution of architecture to cell phenotype cannot be qualified from this study.

11.4.2 Promoting cellular alignment Although randomly oriented scaffolds may be useful, a majority of naturally occurring tissues, including the bladder (see Fig. 11.5), exhibit preferential collagen fibre alignment (Baker et al., 2006). This fibre alignment endows the tissue with unique functional material properties that vary depending on the direction of the applied load. © Woodhead Publishing Limited, 2011


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Stroma

Urothelium

11.5 Optical micrograph of a transverse section of porcine bladder tissue stained by the van Gieson method, illustrating the aligned collagen fibrils, (Scale bar = 50 µm).

Electrospinning of aligned fibres is usually achieved by the relatively simple process of increasing the rotational speed of the fibre-collecting mandrel so that it effectively winds the fibres onto it as they are collected. The speed required is dependent on the polymer used and other process parameters: for example, increasing the mandrel rotation rate so the collecting surface was moving at 5 m s–1 was sufficient to create well-aligned fibres of polystyrene (Baker et al., 2006). Aligned electrospun materials have also been created using polyglycolic acid (Boland et al., 2001) and collagen type I (Matthews et al., 2002), although both examples retained substantial numbers of misaligned fibres indicating a need for further optimisation of the collection parameters. If high enough speeds are used, the mandrel rotation rate can even be used to draw the fibres into thinner diameters as they are collected. The key problem faced by scientists using aligned electrospun fibre materials is the reduction in porosity, which occurs because aligned fibres pack together much more densely than randomly oriented fibres of a similar diameter. This may not prove to be a problem where cells are concerned, so long as fibres are collected dry and do not anneal to each other, as they will remain free to be moved aside by invading cells. Vascular smooth muscle cells grown on aligned electrospun poly(llactide-co-e-caprolactone) were reported to develop aligned cell bodies and cytoskeletal (actin/myosin) networks, although the extent of alignment was not quantified (Xu, 2004). In a study of urinary tract stromal cells grown on random and aligned electrospun polystyrene scaffolds, the alignment of electrospun fibres and F-actin microfilaments were analysed using image analysis software (Image Pro Plus® from Media Cybernetics) (Baker et al., 2006). In each case, the scaffold or cellular ‘fibres’ were grouped into subsets of 20° based on their alignment and plotted on a distribution histogram to derive an alignment factor, based on the height of the main histogram peak. An alignment factor of 1 corresponded to 100% fibre alignment (within



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20°) and an alignment factor of 0.11 corresponded to totally random fibre orientations. The alignment of cellular actin microfilaments (0.19 and 0.74), closely matched the alignment of the scaffold fibres (0.15 and 0.66), for random and aligned scaffolds, respectively (Baker et al., 2006). Furthermore, aligned scaffolds and cells more closely matched the alignment of native bladder tissue (0.51) than did random scaffolds (Baker et al., 2006). These observations are additionally supported by studies of human ligament fibroblasts (Lee et al., 2005), cardiomyocytes (Zong et al., 2005) and neural stem cells (Yang et al., 2005) cultured on aligned electrospun scaffolds, although none of these studies quantified cellular alignment. Cellular alignment occurs embryologically owing to strain patterns and growth factor gradients, which have yet to be effectively recapitulated in tissue engineering, although the potential for using aligned biomaterials to promote function and anisotropic strength could prove to be important. In the case of bladder stroma, the detrusor muscle is aligned in layers, which run perpendicular to each other (Fig. 11.1) and, in theory, this might be mimicked by forming a composite scaffold of several aligned scaffolds combined in perpendicular orientations. These findings could be beneficial in tissue engineering applications where the alignment of cells on a biodegradable biomaterial could lead to the deposition of an aligned collagen matrix. The effects of scaffold alignment on cell phenotype go beyond mere morphology as shown by Lee et al. (2005), who observed an increase in total collagen production by human ligament fibroblast cells cultured on aligned versus randomly orientated scaffolds of electrospun polyurethane (PU). Another study found that aligned actin networks were associated with the deposition of aligned ECM molecules, in particular fibronectin, which could be important in mimicking the mechanical properties of tissues (Halliday and Tomasek, 1995). Studies of micro-patterned surfaces have also demonstrated the importance of cell shape in regulating smooth muscle cell differentiation. Elongated embryonic mesenchymal cells were found to differentiate into smooth muscle cells, whereas rounded cells remained negative for markers of smooth muscle differentiation (Yang et al., 1999).

11.4.3 Surface modification and functionalisation Plasma treatments are commonly used to modify the surface chemistry of hydrophobic polymer scaffolds to enhance ‘wettability’ and this also affects changes in the spectrum of adsorbed serum proteins. The high energy of gas plasma creates new chemical groups on the scaffold surface and increases the surface roughness of scaffolds (Qu et al., 2007). It has been suggested that such nano-scale features induced on PU or poly(lactic acid-co-glycolic acid) (PLGA) films enhanced bladder smooth muscle cell proliferation (Thapa et al., 2003a, 2003b), although whether changes in serum adsorption played a role was not assessed. © Woodhead Publishing Limited, 2011


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As discussed elsewhere (Bolland and Southgate, 2008), one of the challenges of functional bladder tissue development is the need for rapid and complete graft vascularisation to prevent fibrosis and shrinkage, which is a recognised problem for thick tissues (Ko et al., 2007). In the case of composite cystoplasty, the vascularisation requirement is catered for by the use of a prevascularised host tissue (Fraser et al., 2004; Turner et al., 2011). In one report, the omentum, which is a large and heavily vascularised fold of peritoneum, was wrapped around the construct providing an immediate and effective vascularisation bed (Atala et al., 2006). Others have explored the delivery of vascular endothelial growth factor (VEGF) by genetic modification (Guan et al., 2007) or by incorporation into a PLGA sponge (Elcin and Elcin, 2006). As indicated above (Section 11.3.1), surface modification of a scaffold by the covalent attachment of heparin offers a route to sequestering matrix-binding bioactive factors, such as VEGF, in a natural configuration for presentation to cells (Rohman et al., 2009) and this surface modification process could be applied to electrospun scaffolds.

11.5

Future trends

Electrospinning has unmet potential for bladder regeneration, particularly if a new generation of more elastomeric biodegradable polymers is developed for this field. Co-spinning of synthetic elastomeric polymers with natural bladder extracellular matrix proteins could deliver novel materials with defined mechanical/architectural properties. The potential exists for these to be examined in terms of bladder tissue generation in vitro, prior to testing in vivo. The delivery of bioactive growth factors or other cytokines from the scaffolds is another area of potential development. In theory it would be possible to incorporate biphasic factor release, with an immediate effect factor incorporated by surface immobilisation on the finished scaffold and with a second phase factor incorporated into the bulk of the polymer prior to spinning. For example, this could allow an initial phase of cell recruitment and proliferation into the scaffold, followed by tissue maturation owing to the release of differentiation-inducing factors. Alternatively, an angiogenic factor could be tagged onto the scaffold surface to promote early vascularisation, with nerve growth factor incorporated into the bulk being released slowly over the following months as the material biodegrades.

11.6

Conclusions

In summary, control over scaffold architecture afforded by electrospinning has the potential to promote the morphogenesis of functional tissue constructs, including the bladder. However, there remains a need for systematic study



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and rationalisation, as the field of three-dimensional tissue biology is in its infancy and cannot yet be used to inform how biomaterial design can harness the regenerative and developmental processes required for successful tissue engineering. Now, more than ever, is the time for materials scientists to work in collaboration with cell biologists to drive forward an understanding and translation of biomaterial design into tissue engineering and regenerative medicine applications.

11.7

Acknowledgement

The authors thank Felix Wezel for critical review of the manuscript. Jennifer Southgate is supported by York Against Cancer.

11.8

References

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Thapa, A., Webster, T. J. and Haberstroh, K. M. (2003b) ‘Polymers with nano-dimensional surface features enhance bladder smooth muscle cell adhesion’. J Biomed Mater Res A, 67, 1374–83. Truschel, S. T., Wang, E., Ruiz, W. G., Leung, S. M., Rojas, R., Lavelle, J., Zeidel, M., Stoffer, D. and Apodaca, G. (2002) ‘Stretch-regulated exocytosis/endocytosis in bladder umbrella cells’. Mol Biol Cell, 13, 830–46. Tsai, C. C., Huang, R. N., Sung, H. W. and Liang, H. C. (2000) ‘In vitro evaluation of the genotoxicity of a naturally occurring crosslinking agent (genipin) for biologic tissue fixation’. J Biomed Mater Res, 52, 58–65. Turner, A. M. and Southgate, J. (2009) ‘The tissue-engineered urinary bladder’. In Fundamentals of Tissue Engineering and Regenerative Medicine, Meyer, U., Meyer, T., Handschel, J. and Wiesmann, H. P. (eds), Springer-Verlag, Berlin/Heidelberg, 429–40. Turner, A., Subramanian, R., Thomas, D.F.M., Hinley, J., Abbas, S.K., Stahlschmidt, J. and Southgate, J. (2011) ‘Transplantation of autologous differentiated urothelium in an experimental model of composite cystoplasty’. Eur Urol, 59, 447–54. Varley, C., Hill, G., Pellegrin, S., Shaw, N. J., Selby, P. J., Trejdosiewicz, L. K. and Southgate, J. (2005) ‘Autocrine regulation of human urothelial cell proliferation and migration during regenerative responses in vitro’. Exp Cell Res, 306, 216–29. Varley, C. L., Garthwaite, M. A., Cross, W., Hinley, J., Trejdosiewicz, L. K. and Southgate, J. (2006) ‘PPARgamma-regulated tight junction development during human urothelial cytodifferentiation’. J Cell Physiol, 208, 407–17. Webb, A. R., Yang, J. and Ameer, G. A. (2004) ‘Biodegradable polyester elastomers in tissue engineering’. Expert Opin Biol Ther, 4, 801–12. Xu, C. Y. (2004) ‘Aligned biodegradable nanofibrous structure: a potential scaffold for blood vessel engineering’. Biomaterials, 25, 877–86. Yang, F., Murugan, R., Wang, S. and Ramakrishna, S. (2005) ‘Electrospinning of nano/ micro scale poly(l-lactic acid) aligned fibers and their potential in neural tissue engineering’. Biomaterials, 26, 2603–10. Yang, J., Webb, A. R., Pickerill, S. J., Hageman, G. and Ameer, G. A. (2006) ‘Synthesis and evaluation of poly(diol citrate) biodegradable elastomers’. Biomaterials, 27, 1889–98. Yang, Y., Relan, N. K., Przywara, D. A. and Schuger, L. (1999) ‘Embryonic mesenchymal cells share the potential for smooth muscle differentiation: myogenesis is controlled by the cell’s shape’. Development, 126, 3027–33. Zong, X., Bien, H., Chung, C. Y., Yin, L., Fang, D., Hsiao, B. S., Chu, B. and Entcheva, E. (2005) ‘Electrospun fine-textured scaffolds for heart tissue constructs’. Biomaterials, 26, 5330–8.



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12


F. A c o c e l l a and S. B r i z z o l a, Università degli Studi di Milano, Italy

Abstract: Tracheal circumferential defects involving more than half of the tracheal wall still represent an unsolved problem. Several studies have developed different methods to help repair cartilage and improve healing but a suitable tracheal reconstruction or replacement has not been achieved yet. Novel bioengineering technologies seem to be the new answer to this serious problem. This chapter briefly describes the fundamentals of the anatomical and physiological tracheal functions and provides a review of trachea tissue engineering. Then it describes the project and development by means of electrospinning of a biodegradable tubular tracheal scaffold with an in vitro and in vivo preliminary experimental approach. Key words: animal model, Degrapol®, electrospinning, neovascularisation, tissue engineered trachea (TET).

12.1

Anatomy of the trachea and main pathologies of surgical concern

12.1.1 Tracheal anatomy The trachea is a single cartilaginous and membranous conduit crucial for the ventilation and the clearance of upper air-way secretions. Part of the trachea lies in the cervical region while a small part of it runs under the sternal notch through the thoracic cavity. The structure of the trachea in the adult human male averages 11.8 cm in length (range 10–13 cm) and 1.6–2.4 cm in width from the lower border of the cricoid cartilage to the top of the carinal spur. The trachea is nearly cylindrical, being flattened posteriorly and composed of 18–22 C-shaped rings of hyaline cartilage (Fig. 12.1). The posterior part of the trachea consists of a muscular structure (pars membranacea) that stretches under inflow and outflow pressure. At the spur level the trachea divides into two main bronchi (right and left) that provide continuity to the respiratory system. The right main bronchus continues more vertically, whereas the left is always more horizontal with respect to the trachea, while in infants the two bronchi lie more transversely. As a base rule in infants more than half of the trachea is found within the neck, whilst in an adult half of the organ is positioned above the sternal notch. This proportion varies during flexion and extensions of the neck caused by the tracheal movement. During childhood the cartilage is thinner and eventually more compressible laterally than in 242 © Woodhead Publishing Limited, 2011


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Tracheal tissue regeneration Oesophagus

Tracheal ring

243

Larynx

Trachea

Bronchi

Left

Right

12.1 Drawing of the human trachea.

adults. The cross-sectional configuration of the trachea may be markedly altered with increasing age, particularly in the presence of chronic obstructive lung disease which could cause different shape deformities with potential clinical relevance. The blood supply to the organ is crucial for tracheal surgical therapy and it is essentially based on three tracheoesophageal branches coming from the inferior thyroid artery. The first branch supplies the lower cervical trachea. The second and the third supply the middle and the upper sections respectively. The bronchial arteries originating from the aorta supply the carina and the lowest part of the trachea. Starting from the vessels reaching the trachea a lateral longitudinal anastomosis is seen. This is crucial for the entire blood supply of the organ. From here several anterior transverse intercartilaginous arteries run deeply in the trachea anastomosing with the contralateral arteries at the midline. This vascular network reaches the inner part of the trachea and submucosa, and forms the sub-mucosal plexus that provides sole nourishment to the cartilage and is an important source of blood for the membranous wall and mucosa. Ultrastructurally the trachea is formed by hyaline cartilage (forming the tracheal ring), smooth muscle cells (forming the pars membranacea) and, going deeper, the sub-mucosal layer which houses the vascular plexus, as discussed above. The luminal surface is lined by the respiratory mucosa made by a pseudostratified columnar epithelium with cilia and goblet cells. These two different kinds of cells are essential to produce mucus for moisturising the airway and excluding foreign particles using vibratile cilia (Fig. 12.2).



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Electrospinning for tissue regeneration 147.9 µm

A 192 µm

211.9 µm

B

C

12.2 Ultrastructural section of the tracheal wall showing hyaline cartilage (A), sub-mucosal vascular plexus (B) and mucosal layer (C)

12.1.2 Pathologies of surgical concern About half of the trachea could be successfully resected and reconstructed by end-to-end anastomosis. Grillo (2002) This is the basic concept of tracheal surgery. On the other hand pathologies involving more than half of the trachea need tracheal substitutes, biological or synthetic, but several clinical situations lead to organ amputation and creation of a tracheostomy with consequential aphonic voicelessness and potential airway infection. In general it is possible to declare that all the pathologies involving the trachea could result in tracheal resection and substitution owing to their extension of disease. As described by Grillo (2003) tracheal diseases can be split into two main categories: malignant and benign. Tracheal tumours are mainly malignant. Primary tumors are squamous cell carcinomas and ‘cylindromas’ (adenoid cystic carcinomas), while secondary tracheal neoplasia are bronchogenic carcinomas of the carina and thyroid carcinomas (owing to their anatomical relationship). Among benign lesions we can distinguish congenital disease (tracheal agenesia, stenosis, esophago-tracheal fistula and laringo-tracheal cleft) and acquired disease (post-intubation lesions and strictures, subglottic laryngeal stenosis, tracheoesophageal strictures and tracheal trauma).



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Tissue engineered trachea (TET)

12.2.1 Tissue engineering and tracheal replacement The management of tracheal pathology, such as stenosis or cancer, often requires a tracheal reconstruction. Primary end-to-end anastomosis after resection is the method of choice and can usually be performed successfully for defects of up to 50% in length (Kon, 1983). Unfortunately, there are cases in which primary anastomosis is not possible, such as after extensive burns, trauma, tumour resection, or post-intubation injuries. Additionally, the treatment of congenital tracheal atresia or stenosis can be hindered by the lack of sufficient tissue for surgical reconstruction, as the length of trachea involved may be extensive. Over the last 60 years, multiple approaches to tracheal reconstruction have been attempted both clinically and experimentally, including the use of prosthetics, autografts, homografts and allografts. Prostheses such as Dacron polyurethane mesh, polytetrafluoroethylene, polypropylene mesh, silicone rubber, and even glass tubes (Jacobs, 1988) have been tried and have often been fraught with complications such as infection, extrusion and stenosis (Table 12.1). Autogenous and alloplastic tissues have also been used from sources such as fascia, skin, bone, periosteum, cartilage, perichondrium, tracheal allografts, muscle, oesophagus, pericardium, dura mater and the small bowel (Fonkalsrud, 1971; Cohen, 1985; Har-El, 1989; Letang, 1990). In any case the main problem for all these grafts remains the development of late stenosis of the prosthesis. Because of these difficulties, investigators are still in search of the ideal tracheal replacement material, which is able to satisfy the requirements of a tracheal prosthesis model and is safe and durable in practical use. First, as Belsey (1950) iterated in the 1950s, the tracheal substitute should be a laterally rigid and longitudinally flexible tube even if the latter requirement has proved to be desirable, but not essential (Grillo, 1965). The tracheal conduit must further be initially airtight and become integrated into adjacent tissues, so that chronic inflammation, granulation tissue, infection and erosion do not occur. Also we cannot forget that the need for an immunosuppressive regime is unfavourable for different reasons but especially in the frequently extensive tracheal cancer that requires a transplant. Furthermore, for a method to be practically considered, the technique of construction or insertion of the conduit must be surgically straightforward and the results must be predictably successful. Finally the materials for tracheal replacement must be biocompatible, non-toxic, nonimmunogenic, non-carcinogenic, must not dislocate or erode over time, should ideally provide or facilitate epithelial resurfacing, should avoid stenosis or late buckling, resist bacterial colonisation, avoid accumulation



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Table 12.1 General approaches used for tracheal replacement

Type

Complications

Synthetic materials Solid

Steel Glass Polyethylene Silicone Teflon Polyurethane

No incorporation and epithelisation

Porous

Steel wire Titanium Collagen PTFE Polyurethane Dacron Teflon

Scar tissue formation Lack of epithelisation

Bioprosthesis Etherologous tissues Cadaver trachea Aortic homograft Autogenous tissues

Fascia Aorta Tracheal wall Cartilage Derma Pericardium Periosteum Buccal mucosa Dura mater Bone strips

Vascularised Intercostal muscle autogenous tissues Diaphragm Oesophagus

Scar tissue formation Deformation, contraction, calcification Malacia and degeneraton Need for foreign material support

Need for foreign material support Requirement for major surgery

of secretions, and must be permanent constructions (Jackson, 1950; Scherer, 1986). Since 1988, tissue engineering of cartilage has been successfully pursued by a number of groups. Osada (1994) and his group carried out a tracheal replacement in ten mongrel dogs using a long knitted Dacron® tube with inner silicone rubber coating seeded with autologous fibroblasts after static culture. The animals long term survival was satisfactory and the results encouraging. During the same year, Vacanti (1994) and associates were able to create new hyaline cartilage in athymic mice by using isolated bovine chondrocytes and biodegradable suture material (polyglactin 910 and polyglycolic acid, (PGA)), which served as a temporary scaffold to which cells attached until they created their own matrix.



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One of the main limitations in this approach is the specific mechanical requirement placed on the trachea. The trachea must maintain flexibility in the longitudinal direction to allow for free movement of the head, whilst maintaining the rigidity necessary to prevent collapse of the trachea during breathing. This is accomplished in native tissue by presence of cartilaginous rings, and is not adequately modelled by a cartilaginous tube. For these reasons Kojima (2002), designed another study to evaluate the ability of autologous tissue-engineered cartilage with a helical shape to provide the structural component of a functional tracheal replacement (Fig. 12.3(a)).

(a)

(b)

12.3 (a) Helical template fabricated with a silicone mold-making kit. (b) The chondrocyte-seeded matrix was placed in the grooves of the template (arrow) and the entire template was wrapped with the fibroblast-seeded mesh.



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Chondrocytes and fibroblasts were seeded onto separate nonwoven meshes of PGA fibres, 14 mm in diameter (Davis and Geck, Danbury, CT, USA). The chondrocyte-seeded mesh was placed in the grooves of a helical template made from Silastic ERTV mold-making kit (Dow Corning) and then covered with the fibroblast-seeded mesh (Fig. 12.3(b)). These implants were placed either in a subcutaneous pocket in the nude rat or in the neck of a sheep. Sheep tissue-engineered tracheas were harvested from the neck at 8 weeks and anastomosed into a 5 cm defect in the sheep trachea. Gross morphology and tissue morphology were similar to that of native tracheas and histology revealed the presence of mature cartilage surrounded by connective tissue. Other biomaterials have also been used in combination with chondrocytes to produce high quality tissue-engineered cartilage. These biomaterials include calcium alginate gels, collagen gels, fibrin glue, and agarose gels, among others. Another study by Kojima (2003a) evaluated the feasibility of creating engineered tracheal equivalents grown in the shape of cylindrical cartilaginous structures using sheep tracheal and nasal septum cartilage-derived chondrocytes. Tracheal and nasal chondrocytes were separately seeded onto PGA matrices and these cell–polymer constructs were then implanted subcutaneously in nude mice for 8 weeks. This approach demonstrated that the cell yield and the property of the resulting engineered cartilages were very similar for both tracheal and nasal chondrocytes (Fig. 12.4).

(a)

(b)

(c)

(d)

12.4 Appearances of tracheal chondrocytes (a) and (c) and nasal chondrocytes (b) and (d) seeded on non-woven mesh of polyglycolic acid (PGA) fibres (Davis and Geck, Danbury, CT) for 8 weeks to create tissue engineered trachea (TET).



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In a follow-up study (Kojima, 2003b), the same group evaluated the feasibility of producing a composite engineered tracheal equivalent lining its lumen with nasal epithelial cells. At first chondrocyte suspensions were seeded onto a matrix of PGA. Cell–polymer constructs were wrapped around silicon tubes and cultured in vitro for one week, followed by implanting into subcutaneous pockets on the backs of nude mice. After 6 weeks the epithelial cells were suspended in a hydrogel and injected into the embedded cartilaginous cylinders following removal of the silicon tube which until then had been used as an internal support for the engineered construct. Implants were harvested after 4 weeks and the morphology of implants resembled that of native sheep trachea while the histology showed the presence of mature cartilage and formation of a pseudo-stratified columnar epithelium. Other polymers were used in subsequent studies with the same success, including a nonwoven mesh of PGA and PGA/poly-l-lactic acid copolymers. Recently, Ruszymah (2005) and colleagues undertook a study to reconstruct the trachea with human nasal septum chondrocytes by using a combination of biodegradable hydrogel and non-biodegradable high-density polyethylene (HDP) as the internal predetermined shape scaffold. After 8 weeks of in vivo implantation, the TET constructs were harvested. The macroscopic appearance of the TET constructs demonstrated that the HDP constructs were 80–90% covered with yellowish glistening cartilage-like tissue without any sign of inflammation. Okumus (2005) and colleagues designed an axial biosynthetic prefabricated flap to reconstruct the circumferential tracheal defects in ten rabbits. The inner mucosal lining was substituted by hairless epithelium obtained from the proximal ear, the tracheal cartilage by polypropylene mesh and the tracheal adventitia by lateral thoracic fascia to ensure a good vascular supply. Then the epithelial graft, polypropylene mesh and lateral thoracic fascia were put into a tube around a silicone catheter and placed into the cervical subcutaneous area for 2 weeks. A silicone catheter was taken out and prefabricated neotrachea adapted to the defect formed in native trachea and anastomised. The constructs were evaluated after 4 weeks by radiological, macroscopical and histological examination. In high level active research there are essentially three emerging concepts: the use and development of dynamic in vitro culture systems (bioreactors), the use and development of microfabrication technologies to create vascularised tissues and organs, and the search for and use of an appropriate multipotent, undifferentiated stem cell in tissue engineering. Producing a dynamic in vitro microenvironment for tissues may be an important aspect in guiding the formation of tissue with certain structural and functional characteristics. The in vitro modulation of chondrogenesis by dynamic cell seeding and bioreactors has been investigated. Vunjak-Novakovic (1999) and colleagues found that hydrodynamic conditions in convective-flow tissue culture bioreactors can



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modulate the composition (matrix components), morphology, mechanical properties and electromechanical function of tissue engineered cartilage. These initial results are encouraging for the development of a TET replacement in humans and tissue engineering has emerged as a rapidly expanding approach for addressing the organ shortage problem.

12.2.2 First transplant of a human tissue engineered tracheal prosthesis Moving from their own previous and successful research Macchiarini (2008) have made a tubular tracheal human transplant. A 30-year-old female patient with end-stage airway disease underwent the complete resection of the left main bronchus with a bioengineered human trachea replacement. This was achieved by using autologous epithelial cells and chondrocytes that were isolated from biopsies and subsequently cultivated in vitro. The confluent cells were then seeded onto a matrix consisting of decellularised tracheal segment retrieved from a 51-year-old female transplant donor. The chondrocytes were seeded onto the external surface whilst the epithelial cells were seeded onto the internal side with the same density of 1 ¥ 106 cells per ml. The newly seeded tracheal segment was incubated in a bioreactor for 96 hours to control cell proliferation and matrix deposition. After this time, the patient underwent reconstructive surgery with positive results. The post-operative course was uneventful and at 14 days, 1 month, 2 months and 3 months the graft appeared to be healthy. This is a highly important step for the production of a tissue engineered tracheal substitute and gives great hope for the future of the airway reconstruction.

12.2.3 Tracheal cartilage, structure and biomechanics Owing to the structural and functional complexity of the organ, the various tissue components must be considered independently to realise a feasible tracheal analogous substitute. It is clear that the cartilaginous component of the tracheal rings will play a central role in the development of a functional tissue, being resistant to the inspiratory collapse and, at the same time, characterised by adequate flexibility. Tracheal cartilage, a type of hyaline cartilage, is a composite material in which the collagen fibres, essentially type II, are immersed in a hydrated proteoglycan matrix (aggrecan, versican, hyaluronane). Its main function is to keep the airway wall open despite intrathoracic pressure differences during breathing that would otherwise cause it to collapse and limit air flow. Moreover the cartilaginous rings are crucial in determining the dimension and compliance of the trachea; indeed, the rings limit prolapse of the pars membranacea, when the transmural pressure is positive, or its invagination,



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when the transmural pressure is negative. Human tracheal cartilage consists of a collagen-rich ablumenal superficial zone, a proteoglycan-rich core and a collagen-rich lumenal superficial zone. At each surface there is a gradual transition between the superficial zone of the cartilage and a collagenous perichondrium. Chondrocytes are in lacunae surrounded by extracellular matrix and in the superficial, collagen-rich zones are closer together and flattened parallel to the plane of the nearest cartilage surface. Collagen fibrils in the superficial zones are oriented in the plane of the cartilage surface whilst in deeper layers of the cartilage, collagen fibrils are oriented less regularly and the proteoglycan content decreases (Fig. 12.5). The biomechanical properties of the tracheal cartilage are a consequence of its structural features. Any biomechanical change in airway cartilage could influence the mechanics of maximal expiratory flow and cough. Age-related changes in biomechanical properties and biochemical composition of airway cartilage could influence the airway dynamics owing to the shortening of airway smooth muscle. (a)

(b)

(c)

(d)

12.5 Scanning electron microscopy images (SEM) of the different tracheal cartilage zones: (a) Luminal surface is to the left with chondrocyte lacunae visible in an abundant regular layering of the matrix (Scale bar: 40 µm). (b) Higher power view of the boxed area from (a) showing the luminal superficial zone. Rows of collagen fibrils (arrows) are oriented perpendicular to one another (Scale bar: 10 µm). (c) SEM of the peripheral abluminal side with perichondrium on the left (Scale bar: 100 µm). (d) Higher power view of the boxed area from (c) showing perpendicular fibres deposited on this side (Scale bar: 10 µm).



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The equilibrium tensile modulus of airway cartilage may be most directly relevant to normal breathing, where cartilage provides a relatively constant load to the trachealis muscle and maintains airway wall stiffness over a range of airway diameters that are produced by changes in the tone of the trachealis muscle occurring over minutes to hours. Furthermore, the equilibrium tensile modulus is likely to be independent of proteoglycans in this tissue, as has been shown for articular cartilage (Schimdt, 1990). In contrast, in a forced expiration or cough, transmural pressure changes occur in less than a second and the viscoelastic properties of the airway cartilage may be most important in maintaining airway calibre. Under either a sustained force or a transient force, the resistance of tracheal cartilage to bending depends on the tensile modulus in layers furthest from the neutral axis (i.e. the ablumenal superficial zone) and the compressive modulus of the central zone of the cartilage. The tensile modulus is primarily a function of collagen fibril organisation and the compressive modulus is a function of aggregating proteoglycans that are abundant in the central zone. There is substantial experimental data in the literature (Rains, 1992) regarding the Young’s modulus value (E) of the tracheal cartilage which differs widely depending on age-related individual changes or cartilage portion and layer creating structural inhomogeneity, whilst the Poisson modulus value (n) is generally constant (Table 12.2).

12.3

Electrospun biodegradable tubular tracheal scaffold

12.3.1 Scaffolds and tissue engineering This section describes the development of a biodegradable microstructured tubular scaffold created by electrospinning to mimic the rabbit native trachea shape and functional characteristics and its experimental evaluation both in vitro and in vivo. The architecture of an engineered tissue substitute plays an important role in modulating tissue growth. The primary function of a scaffold is tissue conduction and, therefore, it must allow cell attachment, migration onto or within the scaffold, cell proliferation and cell differentiation. It must also provide an environment where the cells can maintain their phenotype Table 12.2 Mechanical features of the human tracheal cartilage (tensile properties). E is Young’s modulus value; n is Poisson modulus value Age(years)

E (MPa)

n

References

80–81 36–74 17–81 17–81

16.67 2.5–7.7 1.8–15 4.6–13.6

0.3 0.3 0.3 0.3

Begis (1988) Lambert (1991) Rains (1992) Roberts (1991)



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and synthesise required proteins and molecules. Scaffold characteristics include high porosity, high surface area, structural strength and specific three-dimensional (3D) shape. Although various structures of engineered tissue scaffolds have been developed for tracheal replacement, the goal of producing a clinically useful tissue scaffold is still far from being realised. An ideal tissue engineered scaffold should be mechanically stable and able to function biologically in the implant site (Thomson, 1997). Mechanical stability is dependent primarily on the selection of the biomaterial, the architectural design of the scaffold and the cell–material interactions. The ultimate goal of the scaffold design is the production of an ideal structure that can replace the natural extracellular matrix (ECM) until host cells can repopulate and synthesise a new natural matrix. To achieve this goal the scaffold material must be carefully selected and the scaffold architecture must be designed to ensure that the seeded cells are biocompatible with the engineered scaffold. The surface chemistry of a tissue engineered scaffold is dependent upon the type of the material, ranging from natural biopolymers to synthetic polymers. The most commonly used natural biopolymers include demineralised bone matrix (Dahlberg, 1991), agarose and collagen (Watt, 1988), hyaluronan (Allemann, 2001), basement membrane (Ponticiello, 2000) or alginate (Bonaventure, 1994). Synthetic polymers that are used include degradable polyesters, such as PGA and polylactic acid (PLA) and their copolymers, poly(d,l-lactide-co-glycolide) (PLGA) (Agrawal, 2001). These biodegradable polymers have a long history of clinical use and currently are used in various tissue engineering applications. Furthermore the tendency of host tissue to form a fibrous capsule around an implant is an important challenge for scaffold biocompatibility. The encapsulation of porous implants depends partially on their material architectural features. There are in general two architectures of porous biomaterial implants: those made of interconnected fibres (fibro-porous meshes) and those made of interconnected pores (porous meshes). Softtissue response is sensitive to the geometric features of them both (Davilla, 1968; Jansen, 1992). Basically materials with small pores sizes ( 80–100 mm). Sanders (2000, 2005) and his group demonstrated the fundamental role of the fibre diameter and pore dimension on cell adhesion and proliferation on an electrospun fibro-porous mesh (Fig. 12.6). It is clear, however, that tight control of fibre diameter is important. Some studies reported in the literature (Paqu ay, 1996) suggest that the reason why large-diameter fibre meshes become encapsulated is related to capsule formation around individual fibres within the mesh. Capsules from adjacent fibres may eventually interact to form a single capsule around the entire perimeter of the implant. It is only with large fibre spacing that interaction between



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Implant with a well-defined capsule at the perimeter (%) Encapsulated fibres in implant cross-sections (%)

Encapsulation (%)

60 50 40 30 20 10 0

0

10

20

30 40 50 Fibre spacing (µm)

60

70

12.6 Encapsulation dependence on fibres spacing. Only the 6 µm spacing meshes showed encapsulation around the perimeter. All meshes showed some encapsulation of individual fibres (Sanders, 2005).

70 µm

25 µm

Fibre Chondrocyte

12.7 Bidimensional model showing ideal space in intra-fibres (30/80 µm) and ideal fibre diameters (6/25 µm) for fibre-porous meshes to be more suitable for adhesion, proliferation and differentation of cells.

capsules around adjacent fibres and the consequent perimetric encapsulation would be avoided. With small-diameter fibres, where most of the fibres are not encapsulated, interaction between adjacent capsules is reduced and the minimum spacing between fibres necessary to avoid encapsulation at the perimeter is reduced. Thus, according to what is reported in the literature the ideal fibre-porous mesh with 6/25 mm fibres diameter and an inter-fibre space of 30/80 mm could promote cells adhesion, viability, proliferation and also exchange of nutrients (Fig. 12.7). The electrospinning technique is effective in generating very small diameter fibres ranging from 0.05 mm to 5 mm which are useful in different



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applications such as separation membranes, wound dressing materials or artificial blood vessels (Doshi, 1995; Hohman, 2001). In our studies, we chose the DegraPol® polymer because it combines the advantages of traditional polyesters with high processability and marked elasticity properties; its use has been investigated for a long time, showing promising results especially in the fields of cartilage tissue engineering, with appropriate cellular attachment, growth and proliferation. As described by Saad (1996–1999) Degrapol® has been successfully tested under different forms (i.e. foam) both in vitro and in vivo using different kinds of cells.

12.3.2 DegraPol®: a degradable block polyesterurethane for tissue engineering DegraPol® is a degradable block polyesterurethane, consisting of crystallisable blocks of poly((R)-3-hydroxybutyric acid)-diol and blocks of poly(ecaprolactone-co-glycolide)-diol linked by a diisocyanate. Being a block copolymer, DegraPol® combines the advantages of traditional polyesters with high ductility and marked elasticity properties. The use of DegraPol® foams and porous membranes as scaffolds for tissue engineering has been investigated for a long time, showing promising results especially in the fields of cartilage (Raimondi, 2004), trachea (Yang, 2003) and smooth muscle (Danielsson, 2006) tissue engineering. There is significant evidence of DegraPol® demonstrating in vitro and in vivo biocompatibility properties. Saad described in his studies the response of in vitro cultured fibroblasts, macrophages and osteoblasts to the crystalline domain of the material. He reported that phagocytosis of the crystalline segments cause dose-dependent cell activation, cell damage and cell death in macrophages but not in fibroblasts. These studies moreover indicated that DegraPol® exhibits good cell compatibility and does not induce cytotoxic effects in osteoblasts and chondrocytes. In another study, Saad (1997) also performed subcutaneous implantations in rats of polyesterurethanes samples processed into films (5 mm in diameter and 150 mm in thickness) by compression molding; the results of these in vivo experiments showed that two months after implantation the thickness of the capsule was lower than 30 mm. Borkenhagen (1998) described in vivo performances of DegraPol® as a tubular structure used as a nerve guidance channel, showing that the inflammatory reaction associated with polymer degradation does not interfere with the nerve regeneration process. Recently (Riboldi, 2005), DegraPol® scaffolds were manufactured by electrospinning in the novel form of microfibrous membranes with fibres about 10 mm in diameter and a fibre-to-fibre distance of about 10 mm. This advance, together with the above cited findings about the polymer biocompatibility, suggests that DegraPol® holds promise for use as a scaffold for tracheal cartilage tissue engineering.



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12.3.3 DegraPol® characteristics The molecular weight of the electrospun polymer used was 60 765 Da measured by gel permeation chromatography (GPC) and it was chemically composed of hard segments (40%) and soft segments (60%), providing the material with specific long-term degradation behaviour in water at 37.5°C, as shown in Fig. 12.8.

12.3.4 DegraPol® electrospun tubular scaffold A small tubular structure was obtained by collecting fibres on a solid, cylindrical, rotating iron collector (~300 rpm) with a 5 mm diameter, fixed to a drill and positioned 20 cm from the pipette (Fig. 12.9). A chloroform Degrapol® solution (Merk KGaA) was initially heated to a maximum temperature of 40°C preventing the polymeric solid particles from degrading during melting. Remarkable changes in the collected fibres were observed when varying the electrospinning and/or solution parameters: ∑

∑

Solution flow-rate: larger fibre diameters were observed with increasing flow rate as shown in Fig. 12.10. The mean fibre diameter obtained with a flow rate of 1.19 ml h–1, 2.38 ml h–1 and 3.57 ml h–1 were 10 mm, 20 mm and 25 mm, respectively. Temperature: Fig. 12.11 shows scanning electron microscope (SEM) 100000 90000 Molecular weight (Da)

80000 70000 60000 Mp

50000 40000

Mw

30000 20000 10000 0 0

20

40

60

80 100 Time (days)

120

140

160

180

12.8 DegraPol® degradation in buffered aqueous solution at 37°C. After 4 weeks of hydration, the average Mp of the polymer has dropped to approximately 45% of the initial value. At 12 weeks, this value has decreased further, approaching 25% of the starting value. Within 25 weeks, the average Mp value is below 10 000 Da (Mp is the peak average molecular weight, Mw is the weight average molecular weight).



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1 cm

12.9 Collector spindle, consisting of a 5 mm diameter metal cylinder, fixed to a drill. A thin layer (white) of fibres can be seen on the spindle. 50 µm

(a)

100 µm

(b)

50 µm

(c) ®

12.10 Three chloroform Degrapol solutions (Merk KGaA) were electrospun at different flow rates: the fibre diameter increased with increasing flow rate. (a) 1.19 ml h–1, (b) 2.38 ml h–1, (c) 3.57 ml h–1 (Scale bar = 50 µm). 50 µm

(a)

50 µm

50 µm

(b)

(c)

12.11 Three scaffold surfaces obtained under different thermic conditions, resulted in increasing interconnected fibres at higher temperatures. (a) 25°C, (b) 20°C, (c) pre-dipping in liquid nitrogen (Scale bars = 50 µm (a, c) and 100 µm (b)).

images of reticulate fibres produced under different environmental thermic conditions. The electrospun scaffold porosity, resulting from interconnections and melting of fibres, appears to vary according to the



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∑


environmental temperature (T), that is porosity decreases with higher temperatures. The first photograph (SEM) was taken in the summer, at an ambient temperature of approx. 25°C and 60% humidity; the middle one was taken in the winter, at an ambient temperature of approx. 20°C and 40% humidity. In the third image, the collector spindle had been pre-dipped in liquid nitrogen (–170°C), so that the fibres are fully disconnected and porosity is remarkably increased. Concentration: the solution concentration plays a major role in the fulfilment of reticulate fibres characterised by specific size, intrafibre spaces and porosity. The solution viscosity increases with higher polymer content, resulting in an increased mean fibre diameter and porosity and decreased fibre interconnections as described below. A solution concentration of 25–27% meets our requirements perfectly.

Various tests led to reticulate fibres with a diameter ranging from 10 to 25 mm, intrafibre spaces of 30–70 mm and good porosity, resulting from poor fibre interconnections. Figure 12.12 indicates that the reticulate fibre obtained under the conditions listed in Table 12.3 is similar to the model described in Fig. 12.7.

12.4

Scaffold fulfilment

12.4.1 Selection of the scaffold profile In mammals, the tracheal architecture and profile create a fixed cross-sectional area of the trachea during extreme bending caused by body, head and neck

22 µm

80 µm

30 µm

12.12 SEM Image of an electrospun scaffold section with fibre spacings measured according to the standards specified in literature.



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Table 12.3 Operational conditions and solution properties utilised in the production of electrospun reticulate fibres Variables

Values

Solution flow rate Temperature Solution concentration

3.57 ml h–1 –170°C 27% (w/w)

The collector spindle was pre-dipped in liquid nitrogen prior to covering with fibres.

movements, preventing the obstruction of its inner lumen and the resulting respiratory arrest. This particular physiological behaviour derives from the regular distribution of cartilage rings along the trachea and the bronchi and prompted us to create a scaffold with this specific structural characteristic. Therefore, we studied three different profile models: spiral, ringed and toothed, with identical functional properties mimicking as near as possible the anatomical peculiarities and functional necessities of the mammalian trachea, as shown in Fig. 12.13. The goal was to identify which of these three profile models resulted in the best shape and function and could then be tested in a preclinical animal model.

12.4.2 Development of the spiral profile In the case of bending stress, the spiral profile prevents kinking of the structure, as the coils collide filling the intra-ring spaces, whilst leaving the inner lumen unchanged. First, a very thin tubular structure with a diameter of 0.8 mm was spiral-coiled around the collector wire (0.5 mm diameter) and then covered with a further layer of fibres (6/25 mm diameter) providing stability and consistency for the structure. The spiral scaffold (Fig. 12.14) obtained, shows all the design characteristics depicted in Fig. 12.13. The real issue for this type of electrospun structure was represented by nonhomogeneous fibre deposition, mainly in the proximity of the coils, resulting in a local high-porosity size (> 80–100 mm) structure that is unfavourable to the survival of cells. The same problem occurred to the internal spiral-rolled pipe, although it would have favoured the passage of growth factors during cell culturing. Finally, assembly of the structure caused many problems, such as manually extracting the small pipe from the wire or coiling the pipe around the spindle at a specific distance.

12.4.3 Development of the ringed profile A 1 mm thick electrospun tubular scaffold with a 5 mm diameter was initially utilised for the development of the ringed scaffold. After this scaffold had been covered with cyclohexane and slowly dipped in dry ice for refrigeration,



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8

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12.13 Three scaffold models and relevant measurements (mm) investigated for tracheal scaffolds. The diameter of the inner lumen is 5 mm for all models.

it was fixed perpendicularly to the blade for the next process. A cryostat (Shandom ‘Microtome’ 5030, Great Britain) operating under low temperature ambient conditions (–20°C), has been used to dissect the tubular scaffold into rings 2 mm in length. The same 5 mm collector spindle was slipped into the rings spaced at 1 mm intervals and coated with an extra layer of electrospun fibres (6/25 mm diameter). The problems reported in this case coincide with those of the spiral scaffold. The inconsistency of porosity near the rings and assembly problems were major issues. The ringed scaffold is shown in Fig. 12.15 and Fig. 12.16

12.4.4 Development of the toothed profile The starting conditions of the electrospinning equipment and the solution (Table 12.3) are unchanged, including the type of collector spindle with a



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5 mm

(a)

5 mm

(b)

12.14 Trachea-scaffold type 1 made by electrospinning: spiral profile. External structure (a) and manual bending (b) demonstrating the structure kink resistance.

12.15 Ringed profile. Image showing the longitudinal section of the scaffold level with the ring.



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Electrospinning for tissue regeneration 5 mm

(a)

5 mm

(b)

12.16 Trachea-scaffold type 2 made by electrospinning: ringed profile. External structure (a) and manual bending (b) demonstrating the structure kink resistance.

5 mm diameter. First, the collector spindle was dipped in liquid nitrogen for a few seconds and immediately processed for electrospun fibres deposition. After approximately 30 min a layer of fibres 2 mm thick (6/25 mm diameter) has been obtained and we proceeded with the mechanical lathe processing of the scaffold. In this case, a 1-mm thick blade for metal cutting has been used running for 2 mm along the longitudinal section of the scaffold (Fig. 12.13) and leaving a 0.5 mm layer of fibres on the spindle. All the problems that related the inconsistency of the porosity in both previous scaffolds, were overcome by means of a layer of fibres obtained in a single process. Moreover, this method was carried out in controlled automated operational steps. Figure 12.17 shows the toothed scaffold.

12.4.5 Inner layer In order to obtain an inner homogeneous surface that would be free of tissue overproductions, inclinations and narrowing of the surface, we decided to



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5 mm

(a)

5 mm

(b)

12.17 Trachea-scaffold type 3 made by electrospinning: toothed profile. External structure (a) and manual bending (b) demonstrating the structure kink resistance.

install a very thin electrospun microporous layer (approximately 100 mm thick with fibres of 20/25 mm diameter) in the inner lumen, enabling permeability of gas, growth factors and cells nutrients, and, at the same time, avoiding cell proliferation and cartilage tissue growth, which would otherwise result in the occlusion of the inner lumen. This microporous layer was obtained by electrospinning a different solution composed of Degrapol® (16% w/w) dissolved in chloroform (90%), with the addition of methanol (10%). This particular layer has been obtained by processing the collector spindle for electrospun fibres deposition for approx. 1 min (ambient temperature of approx. 25˚C) prior to refrigeration and coating it with the final layer of fibres (Section 12.4.4). Figure 12.18 shows the SEM image of the obtained inner layer.

12.5

In vitro and in vivo evaluation of the cell and tissue response

As previously described, one of the major approaches in tissue engineering includes the seeding and culture of cells on a 3D scaffold prior to its



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100 mm

12.18 SEM image of the inner degrapol® electrospun microporous layer made from a solution of polymer dissolved in chloroform and methanol (90:10 mix).

implantation. Other fundamental requirements concerning the scaffold geometry, such as porosity, pore interconnectivity or pore size to ensure development of an in vitro functional tissue after cell seeding attachment and migration, are of great importance. Moreover it has been demonstrated that during expansion in the monolayer, chondrocytes de-differentiate, assuming a more flattened appearance and producing type I instead of type II collagen; however, the chondrocytic phenotype may be rescued by transferring cells to a 3D culture system (Benya, 1982; Bonaventure, 1994; Jakob, 2001). Our initial in vitro experiments investigated the static culture of chondrocytes on the electrospun microstructure scaffold, in order to determine whether our scaffold provided a good 3D environment for cell adhesion and proliferation and whether the chondrocytic phenotype was preserved. Later we investigated the effect and reliability of dynamic conditioning on the development of in vitro tracheal engineered tissue for a period of 14 days. Lastly, we determined which tissue harvesting technique was the easiest, safest and most minimally invasive and what was the optimal source of cartilage cells. For this reason we investigated the isolation and expansion of rabbit chondrocytes from different anatomical sites to determine an ideal cell concentration for cell seeding.



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12.5.1 In vitro static culture Briefly autologous chondrocytes were harvested from xifoid, chondrocostal and femural cartilage of a New Zealand White rabbit, digested in 0.3% collagenase II, and the cells serially passed. After the expansion phase (6–8 weeks) the chondrocytes were seeded onto the external surface of a half-cut scaffold with a seeding density of 25 ¥ 106 cell/ml, and the cells–polymer construct maintained during in vitro static culture (Fig. 12.19). Five days after static cell seeding, SEM images showed an even distribution of chondrocytes throughout the scaffold, suggesting appropriate exchange of gas and nutrients (Fig. 12.20(a) and (b)).

12.5.2 In vitro dynamic culture We designed a culture model system with an ad hoc bioreactor to allow better colonisation of the electrospun tracheal scaffold that favours the transport of nutrients from the medium to the adhering cells through strict control of culture parameters. Briefly, scaffolds were installed on the bioreactor and chondrocyte suspensions were seeded onto the external scaffold surface with gentle rotation in order to cover the whole circumference of the scaffold. The seeded scaffolds were transferred to the incubator (37°C, 5% CO2) and maintained in culture medium for about 2 hours to enhance cell attachment. After this time, the bioreactor was turned on gradually to reach a rotation speed of 5 rpm, moving up to 7 rpm after a further two hours until the end of the culture period. The scaffold was imaged by environmental scanning electron microscopy

1 mm

Seeding area

12.19 SEM micrograph showing 5 days of static culture for chondrocytes on the seeding area surface of the Degrapol® toothed profile matrix (magnification 12x).



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20 µm

(a)

20 µm

(b)

12.20 SEM micrographs demonstrating adhesion and infiltration of chondrocytes on the scaffold surface (a) and within intrafibre spaces in the 3D scaffold (b) after 5 days of static culture.

(ESEM) following culture for 14 days in the bioreactor. The images demonstrated the adherence of the cell on the scaffold, suggested by their flattened shape, cell spreading and the fading edges of the cells. The images also demonstrated that cells cling to the material’s fibres through numerous filopodia that were clearly visible within the interfibres spaces (Fig. 12.21). The validity of the scaffold’s design and processing was confirmed by the presence of adhering cells that were restricted to the external rings of the scaffold: cellular adherence was effectively hampered by the occlusion of porosity established at the bottom of the grooves and along the rings wall (Fig. 12.22). Examination of the longitudinal section of the scaffold indicates satisfactory cellular colonisation of the matrix thickness: a remarkable number of cells



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20 µm (a)

2 µm (b)

12.21 ESEM images of the scaffold following a 14-day period of dynamic culture in a bioreactor. Evidence of the good adherence of the cellular component to the scaffold fibres was observed by (a) cell spreading and interfibre colonisation. (b) Magnification of the rectangular white box in (a).

penetrated in depth, so that a considerable number of cell aggregates were observed at approx. 570 mm from the external surface. Immunocytochemical examination of the dynamised scaffold samples, using a monoclonal anticollagen type II antibody, showed the presence of chondrocytes both on the surface and throughout the Degrapol® fibres (Fig. 12.23). The number



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20 µm

12.22 ESEM image of the scaffold microstructure at the bottom of the grooves: cellular adherence has been effectively hampered by the occlusion of porosity established during processing.

(a)

(b)

(c)

12.23 Immunocytochemical images showing a clear amount of positive cells on the scaffold surface (white arrows) and their tendency to penetrate deep inside the fibres. The number of collagen type II positive cells was higher with evidence of a thin layer on the external surface (a) and (c) while their numbers decrease thoughout the scaffold thickness (b) (Scale bars = 50 µm).

of collagen type II positive cells was higher, with evidence of a thin layer on the external surface and a decreased number throughout the scaffold thickness. In order to investigate the potential effects of mechanical stimulation on the cells, gene expression for type II collagen, type I collagen, aggrecan and glyceraldehyde 3-phosphate dehydrogenase (GADPH) (control) was determined by a polymerase chain reaction (PCR). Semi-quantitative results of the preliminary tests on gene expression conducted on chondrocytes at



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two different time points, preseeding and post-dynamic culturing in the bioreactor, are reported in Fig. 12.24. The results obtained suggest constant expression of type II collagen in both samples with a marked decrease of type I collagen in the post-dynamic samples. Aggrecan gene expression showed the same aspect in the pre- and post-bioreactor samples. The weak signal in the post-bioreactor samples can reasonably be related to the presence of a poor extracellular matrix produced by the chondrocytes. Data obtained comparing the native cartilage tissue, the pre- and the post-bioreactor samples showed a significant difference between the three tissue types, but only for the expression of type I collagen (p < 0.001). Type II collagen expression was statistically significant between the native and the TET but not between the pre- and post-dynamic culture tissues, which contrasted to the expected results following dynamic stimulation.

12.5.3 In vivo experimental procedures Data obtained after in vivo implantation of the ad hoc designed toothed profile tracheal scaffold in rabbits in an isolated vascular flap using the common carotid artery and the external jugular vein as blood carriers (Fig. 12.25) showed the behaviour of the electrospun polymer in the natural, biological environment with emphasis on degradability, biocompatibility and ability to induce angiogenesis (Brizzola, 2009). The scaffold was made by three 360° rings 2.0 mm in length interposed by 0.5 mm of inter-ring space. The tubular Pre bioreactor

Post bioreactor

GADPH

Collagen II

Collagen I

Aggrecan

12.24 Gene expression for chondrocytes at two different time points: preseeding and post-dynamic culturing in the bioreactor. The presence of GADPH-related frequency bands indicates a satisfactory outcome of the amplification reaction, confirming the expected results for collagen types I and II.



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12.25 Vascular axis formed by the common carotid artery (arrowhead) and the internal jugular vein (arrow) of rabbit. The integrity of the vago-sympathetic trunk was carefully maintained.

skeleton measured 6 mm outside and 5 mm inside diameter, respectively, and the entire length was made up by 1 cm of electrospun Degrapol®, with 6/25 mm fibre diameter and 30/80 mm interfibre space and a molecular weight 60.765 Da. A Teflon cylinder was introduced inside the lumen to prevent internal tissue overgrowth. Rabbits were divided in 3 groups: Group 1, Group 2 and Group 3 refers to two, six and eight weeks of implantation, respectively. The tissue invasion started from the vascular axis and proceeded, time depending, towards the anti-pedicle zone. A thin fibroconnective tissue was clearly present around the tracheal scaffold in all animals (Fig. 12.26). Full polymer resorption during the 8 weeks of implantation was not complete. In all samples examined, fibroblasts and matrix in the form of fibrils were in direct contact with the Degrapol® scaffolds (Fig. 12.27(a)–(h)). Initially (Group 1) few cells and fibrillar matrix adhered to the scaffold and spanned within it (Fig. 12.27(a)–(d)). The number and quantity of cells and fibrillar matrix strongly increased with time and after 8 weeks (Group 3) the scaffolds’ external surface was partially concealed (Fig. 12.27(e)) while the polymer fibres were incorporated in a dense fibrous tissue with interposed cells (Fig. 12.27(f), (h)). Histological examination showed a clear presence of cells and extracellular fibrillar matrix, starting from the external side of the scaffolds and moving towards the inner lumen (Fig. 12.28). The cellular invasion proceeded throughout the scaffold thickness (Fig. 12.29(a)–(c)) and visible new blood vessels were evident mostly in the area of maximum tissue deposition (Fig. 12.29(c)). As predicted, the zones of the



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12.26 Macroscopic evaluation of the complex vascular pedicle scaffold for angiogenesis (asterisks) and neo-tissue formation (arrowheads). Rabbits were divided into three groups: G1, G2 and G3, which refers to implantation at two, six and eight weeks, respectively. The images clearly show a progressive time-dependent neovascular tissue growth onto and through the scaffold starting from the vascular axis.

scaffold closest to the vascular carrier were reached by fibroblast-like cells before the scaffold zones located at a greater distance (Acocella, 2007). The pedicle region showed early signs of neo-angiogenesis by means of tissue deposition accompanied by new formation of blood vessels. In all implanted scaffolds, large vascular spaces filled with erythrocytes were observed in close proximity to the vascular pedicle. A few macrophages (inflammatory cells) were involved in the reactive process and were mainly localised in proximity to the blood vessels. There was no evidence of other cells involved in the immune process being present, confirming the good biocompatibility of the polymer and its fibrous structure. Transmission electron microscopy (TEM) evaluation showed cells to be surrounded by collagen fibrils (Fig. 12.29(d) and (e)). The newly synthesised and secreted collagen was organised into tightly packed fibrils



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(a)

(b)

(c)

(d)

(e)

(f)

(g)

(h)

12.27 SEM images showing (a)–(d) that the Degrapol® scaffold, after 2 weeks implantation, was partially covered by cells and fibrillar matrix, with cells lying in close contact to the scaffold fibres (arrowheads) and bridging between fibres (arrows) as demonstrated by their visible filopodia. After 8 weeks implantation (e)–(h) cells and fibrillar material have increased in number and quantity and conceal a larger area of the Degrapol® scaffold (Scale bars: (a) 50 µm, (b) 5 µm, (c)–(h), 20 µm).

with a regular and parallel arrangement to form bundles (Fig. 12.29(d) and (e)). Fibroblast-like cells, characterised by a spreading phenotype, penetrated into the surrounding extracellular matrix (Fig. 12.29(e)). As the cells proliferated and migrated centripetally filling the pores of the scaffold, small muscle fibres (Fig. 12.29(f)) and endothelial cells defining capillary structure became visible (Fig. 12.29(f) and (g)). Optical immunofluorescence staining showed a regular and increased deposition of collagen starting from



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(b)

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12.28 Histological analysis. Scattered early evidence of fibrovascular invasion (a). The presence of large vascular spaces (white asterisks), provides evidence of a neo-angiogenetic process (b,c), filled with a variable number of erythrocytes (black asterisks) (hematoxylin-eosin).

the surface and filling the biopolymer network (Fig. 12.29(h)). At the end of the invasion process the collagen was also localised around the neo-vessels (Fig. 12.29(i)). Optical and electron microscopy, and immunohistochemistry and immunocytochemistry have been used for this purpose. The monoclonal CD31 antibody analysis, detecting progenitor cells (activate endothelial cells-ECs), showed that positive cells were visibly crowded especially on the external surface of the scaffold (Fig. 12.29(j)) while a short time after cell invasion, the expression of CD31 was localised to the endothelial cells forming the blood vessel walls (Fig. 12.29(k)). Cell numbers increased over time, with many of the cells differentiating into muscle fibres, validated by the expression of a-SMA (alpha smooth muscle antibody) (Fig. 12.29(l)). In all samples few inflammatory cells expressing CD14 were visible (Fig. 12.29(m)), while most cells adopting a pronounced migratory and spreading phenotype showed a strong positivity for cathepsin B, a proteolytic enzyme involved in different biological mechanisms, in other words neo-angiogenesis (Fig. 12.29(n)). Our analysis showed that over time there was a migration of CD31 positive cells, which is generally expressed from endothelial progenitor cells derived by an earlier common myeloid progenitor or hematopoietic stem cell. Based on the hypothesis that CD31 is an early indicator during endothelial differentiation, CD31-bright cells were identified as precursor cells moving from the vascular pedicle and then colonising the full thickness of the Degrapol® scaffold. The scaffolds revealed a noticeable amount of neo-tissue formation that progressively substituted the biopolymer fibres by filling the interfibre spaces with cells and extracellular fibrillar matrix made of collagen. Morphological and immunocytochemical characterisation was consistent with endothelial cell migration, spatial disposition and differentiation of migrating stem cells in various cell types. The positivity of cathepsin B in the present cells advocates its properties in promoting angiogenesis through a process of migration and invasion, already described by others in both in vitro and in vivo studies (Lutgens, 2007; Premzl, 2006). Cathepsin B seems to be implicated in different pathological



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(a)

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12.29 Semi-thin cross-section of excised tracheal scaffold (a)–(c). Progressively, cells and fibrillar material invade centripetally the Degrapol® scaffold. Transmission electron microscope images showing a detail of the Degrapol® scaffold covered by cells and fibrillar material. A large number of collagen bundles longitudinally sectioned (d), with fibroblast-like cells (e) surrounded by collagen bundles cross-sectioned (arrowheads), small muscle fibres (f) and minute blood vessels (g) were visible. Prosthesis crosssectioned (h, i) showing an increased deposition of collagen filling the biopolymer network (h) and surrounding the neovessels (i, v). Immunofluorescence (j, k) with monoclonal anti-CD 31, antiSMA (l), anti CD14 (m), and anti-cathepsin B (n) muscle fibres (l), macrophages (m) and cells characterised by cathepsin B production (n) are visible (white arrowheads). Cryosections of Degrapol® scaffold before (o) and after (p) incubation with cathepsin B enzyme for 1 h at 37°C. The fibres of biopolymer were broken and reduced in dimension (p) owing to enzyme degradation. (Scale bars: (a)–(c), (i)–(p) 100 µm, (d)–(f) 2 µm, (g) 4 µm, (h) 50 µm).



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processes such as tumour angiogenesis, neurodegeneration or abdominal aorta aneurysm (Abisi, 2007) but how it regulates angiogenesis, cell proliferation, invasion and apoptosis is still poorly understood. Moreover, our simple enzymatic activity analysis showed that cathepsin B could be responsible for the polymer degradation similar to degradation of the biological matrix. Our data showed the expression of cathepsin B in progenitor cells and expression of CD31 and SMA, which is consistent with a migratory and angiogenetic cellular phase. To our knowledge, it is the first time that cathepsin B has been found during an in vivo evaluation of a polymer scaffold undergoing degradation (Santerre, 2005). Further studies on the interaction between cathepsin B and Degrapol®, in terms of chemical structure, would improve our knowledge of which part of the polymer is subjected to the degradation process.

12.6

Conclusions

A new tubular tracheal construct, basing on the anatomy of the native trachea, has been designed, using the electrospinning technique. This method allowed us to realise a microstructured scaffold with an optimal geometry for cell adhesion, growth and proliferation. The toothed profile exhibited theoretical functional characteristics that prevent kinking or collapse during eventual physiological movements of the implanted scaffold. The electrospinning process resulted in a 3D microfibrillar network with an optimal fibre diameter (6/25 mm), an interfibre space of 30/80 mm and open porosity on the surface defined as the gold standard in allowing appropriate gas flow and exchange of growth factors and cell nutrients (Sanders, 2005). Furthermore a very thin microporous inner layer was added to the scaffold to prevent tissue ingrowth and the subsequent reduction in the scaffold’s lumen. The absence of a foreign body reaction after implantation in animals confirmed that the architecture of the structure was similar to a natural extracellular matrix, indicating that the electrospun scaffold was suitable as a tissue substitute. This tracheal engineered scaffold showed a specific profile with open porosity (seeding area) on the rings’ surface and closed porosity over the intra-annular walls, controlling and guiding the cellular attachment and proliferation, which then allowed the residing cells to build up their own ECM as the polymer fibres were hydrolysed and degraded over time. The very thin microporous inner layer proved to be very useful in enabling gas flow and exchange of growth factors and cellular nutrients, whilst at the same time preventing tissue ingrowth and thus avoiding reduction of the scaffold’s lumen. The use of a dynamic culture bioreactor system documented a very high efficiency in the cell-seeding procedure. Scaffold movement did not hamper cell adhesion to the fibre surface and seemed to promote cell migration



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throughout the thickness of the scaffold fibres, which favourably supported cell phenotype. The results of the in vivo studies demonstrated that this bioresorbable polymer provided a good substrate for fibrous tissue deposition and formation of neo-angiogenesis throughout the tubular thickness of the scaffold. This is essential for scaffold stability and epithelial growth. Looking forward it is possible to theorise about a new scaffold design that will take advantage from the host biological process to guide more sophisticated and physiological tissue regeneration.

12.7

Acknowledgements

The authors thank abmedica s.p.a., Lainate, Italy for providing Degrapol®.

12.8

References

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reconstruction of the canine trachea with a free revascularized small bowel graft’, Ann Thorac Surg, 49, 955–8. Lutgens SP, Cleutjens KB, Daemen MJ and Heeneman S (2007), ‘Cathepsin cysteine proteases in cardiovascular disease’, FASEB J, 21(12), 3029–41. (Review). Macchiarini P, Jungebluth P, Go T, Asnaghi MA, Rees LE, Cogan TA, Dodson A, Martorell J, Bellini S, Parnigotto PP, Dickinson SC, Hollander AP, Mantero S, Conconi MT and Birchall MA (2008), ‘Clinical transplantation of a tissue-engineered airway’, Lancet, 372, 2023–30. Okumus A, Cizmeci O, Kabakas F, Kuvat SV, Bilir A and Aydin A (2005), ‘Circumferential trachea reconstruction with a prefabricated axial biosynthetic flap: experimental study’, Int J Pediatr Otorhinolarynol, 69, 335–44. Osada H, Takeuchi S, Kojima K and Yamate N (1994), ‘The first step of experimental study on hybrid trachea: use of cultured fibroblasts with artificial matrix’, J Cardiovasc Surg, 35(Suppl. 1–6), 165–8. Paquay YC, de Ruijter JE, van der Waerden JP and Jansen JA (1996), ‘A one stage versus two stage surgical technique: Tissue reaction to a percutaneous device provided with titanium fiber mesh applicable for peritoneal dialysis’, ASAIO J, 42, 961–7. Ponticiello MS, Schinagl RM, Kadiyala S and Barry FP (2000), ‘Gelatin based resorbable sponge as a carrier matrix for human mesenchymal stem cells in cartilage regeneration therapy’, J Biomed Mater Res, 52, 246–55. Premzl A, Turk V and Kos J (2006), ‘Intracellular proteolytic activity of cathepsin B is associated with capillary-like tube formation by endothelial cells in vitro’, J Cell Biochem, 97(6), 1230–40. Raimondi MT, Boschetti F, Falcone L, Migliavacca F, Remuzzi A and Dubini G (2004), ‘The effect of media perfusion on three-dimensional cultures of human chondrocytes: integration of experimental and computational approaches’, Biorheology, 41(3), 401–10. Rains JK, Bert JL, Roberts CR and Paré PD (1992), ‘Mechanical properties of human tracheal cartilage’, J Appl Physiol, 72, 219–25. Riboldi SA, Sampaolesi M, Neuenschwander P, Cossu G and Mantero S (2005), ‘Electrospun degradable polyesterurethane membranes: potential scaffolds for skeletal muscle tissue engineering’, Biomaterials, 26(22), 4606–15. Roberts CR and Parè PD (1991), ‘Composition changes in human tracheal cartilage in growth and aging, including changes in proteoglycan structure’, Am J Physiol, 261(2 Pt 1), L92–101. Ruszymah BHI, Chua K, Latif MA, Hussein FN and Saim AB (2005), ‘Formation of in vivo tissue engineered human hyaline cartilage in the shape of a trachea with internal support’, Int J Pediatr Otholaryngol, 69, 1489–95. Saad B, Ciardelli G, Matter S, Welti M, Uhlschmid GK, Neuenschwander P and Suter UW (1996), ‘Characterization of the cell response of cultured macrophages and fibroblasts to particles of short-chain poly[(R)-3-hydroxybutyric acid]’, J Biomed Mater Res, 30(4), 429–39. Saad B, Hirt TD, Welti M, Uhlschmid GK, Neuenschwander P and Suter UW (1997), ‘Development of degradable polyesterurethanes for medical applications: in vitro and in vivo evaluations’, J Biomed Mater Res, 36(1), 65–74. Saad B, Moro M, Tun-Kyi A, Welti M, Schmutz P, Uhlschmid GK, Neuenschwander P and Suter UW (1999), ‘Chondrocyte-biocompatibility of degrapol-foam: in vitro evaluations’, J Biomater Sci Polym Edn, 10(11), 1107–19. Sanders JE, Stiles CE and Hayes CL (2000), ‘Tissue response to single-polymer fibers



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of varying diameters: evaluation of fibrous encapsulation and macrophage density’, J Biomed Mater Res, 52, 231–7. Sanders JE, Lamont SE, Mitchell SB and Malcolm SG (2005), ‘Small fiber diameter fibro-porous meshes: tissue response sensitivity to fiber spacing’, J Biomed Mater Res, 72A, 335–42. Santerre JP, Woodhouse K, Laroche G and Labow RS (2005), ‘Understanding the biodegradation of polyurethanes: from classical implants to tissue engineering materials’, Biomaterials, 26(35), 7457–70. Scherer MA, Ascherl R, Geissdörfer K, Mang W, Blümel G, Lichti H and Fraefel W (1986), ‘Experimental biosynthetic reconstruction of the trachea’, Arch Otorhinolarynol, 243, 215–23. Schimdt MB, Mow VC, Chun LE and Eyre DR (1990), ‘Effects of proteoglycan extraction on the tensile behaviour of articular cartilage’, J Orthop Res, 8, 353–63. Stankus JJ, Guan J and Wagner WR (2004), ‘Fabrication of biodegradable elastomeric scaffolds with sub-micron morphologies’, J Biomed Mater Res A, 70(4), 603–14. Thomson RC, Yaszemski MJ and Mikos AG (1997), ‘Polymer scaffold processing’, in Principles of Tissue Engineering, Lanza RP, Langer R and Chick WL (eds), R.G. Landes, Austin, TX. Vacanti CA, Paige KT, Kim WS, Sakata J, Upton J and Vacanti JP (1994), ‘Experimental tracheal replacement using tissue-engineered cartilage’, J Pediatr Surg, 29(2), 201–4, discussion 204–5. Vunjak-Novakovic G, Martin I, Obradovic B, Treppo S, Grodzinsky AJ, Langer R and Freed LE (1999), ‘Bioreactor cultivation conditions modulate the composition and mechanical properties of tissue-engineered cartilage’, J Orthop Res, 17, 130–8. Watt FM and Dudhia J (1988), ‘Prolonged expression of differentiate phenotype by chondrocytes cultured at low density on a composite substrate of collagen and agarose that restricts cell spreading’, Differentiation, 38, 140–7. Yang L, Korom S, Welti M, Hoerstrup SP, Zünd G, Jung FJ, Neuenschwander P and Weder W (2003), ‘Tissue engineered cartilage generated from human trachea using DegraPol scaffold’, Eur J Cardiothorac Surg, 24(2), 201–7.



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13

Dental regeneration

I. U. R e h m a n, The University of Sheffield, UK and A. S. K h a n, COMSATS Institute of Information Technology, Pakistan

Abstract: Current treatments for diseases and trauma of dental and oral structures rely on durable materials such as amalgam and synthetic materials, or autologous tissue grafts. A paradigm shift has taken place in utilising electrospun nanofibres for tooth restoration and dento-oral regeneration. Use of electrospun fibres and their scaffolds in dento-oral applications involves several considerations, including choice of material, fibre orientation, porosity, surface modification and tissue application. Choices of materials include both natural and synthetic materials and a combination of both and the surface of the scaffold can be modified by osteoconductive components, which can provide an optimal combination of mechanical and biomimetic properties. This chapter highlights recent advances in the application of electrospun nanofibres in dentistry, including the tooth restoration and periodontium regeneration with specific emphasis on resin matrix interaction, cells, biomaterial matrix scaffolds and integrated tissue engineering approaches. Key words: dental restoration, electrospun nanofibres, periodontal regeneration.

13.1

Introduction

Various biomaterials of natural and synthetic origin and their combination have been investigated for their biocompatibility, biodegradability, cell interaction and mechanical properties and have been successfully utilised as cell carrier materials (Hutmacher et al., 2001; Agarwal and Ray, 2001; Ohgo et al., 2003). The prepared scaffolds with different shapes and structures are expected to provide initial support similar to that of extracellular matrix (ECM) (Muschler et al., 2004). Numerous techniques such as phase separation, particulate leaching and electrospinning have been used to fabricate interconnected porous scaffolds (Weigel et al., 2006). Among these tec hniques, the electrospinning process results in nonwoven meshes with a high surface area to volume ratio, where randomly oriented nanofibres serve as cellular support (Inanc et al., 2009). A number of polymers, such as poly(meta-phenylene isophthalamide), polyetherimide, polyethylene oxide, polyethylene terephthalate, polyaniline, polycaprolactone (PCL) and poly(l-lactic acid) have been successfully electrospun into nanofibres (Reneker and Chun, 1996; Liu and Reneker, 2002; Theron et al., 2004; Shin et al., 2004; Tan et al., 2005). Electrospun 280 © Woodhead Publishing Limited, 2011


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fibres mats are suitable for use as scaffold because of their 3D structure and proposed applications for these materials are in areas such as scaffolding for tissue engineering, nanofibre-reinforced composites, supports for enzymes and catalysts, drug delivery, wound healing and guided bone regeneration (Doshi and Reneker, 1995; Buchko et al., 1999; Coombes et al., 2004; Fujihara et al., 2005; Wutticharoenmonkol et al., 2006). Recently, several in vivo and in vitro studies have demonstrated that electrospun scaffolds support the viability and functional characteristics of chondrocytes (Shin et al., 2006), osteoblasts (Xin et al., 2007) and mesenchymal stem cells (Shih et al., 2006), indicating the potential for utilisation in tissue engineering.

13.2

Periodontal regeneration

The complexity of dental, oral and craniofacial (DOC) structures presents special challenges for the regeneration of cartilage, bone, muscle, tendons, cranial sutures, temporomandibular joints (TMJ), salivary glands, periodontium and teeth. A wide range of scaffolds has been applied to accommodate the needs of regenerating DOC tissues (Moioli et al., 2007). Since 2000 numerous researchers have attempted to develop a periodontal membrane with the necessary properties and features. Current clinical periodontal therapy cannot promote the complete regeneration of periodontal tissues (Cortellini and Tonetti, 2000; Crea et al., 2008; Dori et al., 2008). Therefore, new approaches and materials are required to regenerate the lost periodontal tissues (Bartold et al., 2006; Edwards and Mason, 2006). Two surgical regenerative approaches, namely guided tissue regeneration (GTR) and guided bone regeneration (GBR), have been increasingly used for the repair and regeneration and function of damaged periodontal tissues. In both cases, an occlusive periodontal membrane acts as a barrier to prevent epithelial and connective tissue down-growth into the defect, enabling periodontal regeneration (Karring, 2000; Polimeni et al., 2005; Villar and Cochran, 2010). When choosing a scaffold matrix material for periodontal regeneration and drug delivery, careful consideration of the natural ECM in the periodontal connective tissues may elucidate some candidates. Sufficient mechanical and physical properties must be present to allow its placement in vivo and to avoid membrane collapse. Membrane degradation rates should match those of new tissue formation. Overall, these membranes must function for at least 4–6 weeks in order to achieve successful regeneration of the periodontal system (Karring, 2000; Liao et al., 2007). GTR/GBR membranes have been prepared via film casting (Park et al., 2000; Milella et al., 2001), dynamic filtration (Teng et al., 2009) or electrospinning (Fujihara et al., 2005; Yang et al., 2009) using synthetic (e.g. PLA) and/or natural (e.g. collagen) polymers. Some membranes have included therapeutic drugs (Park et al., 2000), growth factors (Milella et al., 2001) and/or calcium phosphate



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particles (Kikuchi et al., 2004). It should be noted that the effects of some of these additions are complex and not universally accepted. The design and fabrication of a functionally graded periodontal membrane (FGM) via electrospinning holds promise as an interface/implant between alveolar bone and epithelial tissue. The implant needs to utilise a graded structure with compositional gradients and sub-compartments to meet the local functional requirements (Pompe et al., 2003). The rationale of having a periodontal membrane with a graded structure again relies on the idea that it can tailor the properties of the different layers to design a membrane that will retain its structural, dimensional and mechanical integrity long enough to permit periodontal regeneration (Bottino et al., 2011). The feasibility of generating electrospun scaffolds with a functionally graded matrix structure with a distinct chemical composition and tuned mechanical and degradation characteristics has been demonstrated (Thomas et al., 2009; McClure et al., 2010). Recently a design for a trilayered electrospun graft combining bioartificial polymers (i.e. polydioxanone (PDS), gelatin and elastin) has been developed (Thomas et al., 2009), as shown in Fig. 13.1, which not only mimicked the dimensions of the extracellular matrix (ECM) (a)

(c)

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13.1 Scanning electron microscopy (SEM) micrographs of electrospun scaffolds: (a) pure polydioxanone (PDS) scaffold, (b) elastin/PDS blend scaffold, (c) PDS/elastin/gelatin blend scaffold, (d) elastin/gelatin scaffold. The scaffolds have a randomly oriented non-woven fibrous structure with a well-interconnected network of pores. Solvent, 1,1,1,3,3,3-hexafluoro-2-propanol (HFP, 10% wt/v), electric field, 1 kV cm–1; flow rate 2 ml h–1 (adapted from Thomas et al., 2009).



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but also controlled degradation behaviour (Zhang et al., 2009). The addition of biopolymers (e.g. collagen, gelatin or elastin) endowed the membrane with bioactive properties and improved hydrophilicity, ultimately translating into better cell adhesion (Kwon and Matsuda, 2005; Jeong et al., 2008) and growth and function (Zhang et al., 2010). Chew et al. (2008) have shown that aligned electrospun scaffolds are beneficial for transplantation of ‘primed’ cells for enhancing regeneration. This organised scaffold architecture provides topographic cues to adherent cells (Fig. 13.2), resulting in an alignment of cells along the axes of the fibres (Lee et al., 2005; Chew et al., 2006). The ability to control cellular alignment on scaffolds can potentially be a powerful approach for recreating the micro- as well as nanoscale architecture of tissues and this can be important for the regeneration of the periodontal ligament (Yin et al., 2010). So far, aligned fibre structures can be obtained using electrospinning technology (Yin et al., 2010; Yang et al., 2005). However, the electrospinning setup only allows aligned fibres with a limited scaffold thickness to be collected. With increasing thickness, the fibres become disordered as the residual charge accumulation on the deposited fibres interferes with the alignment of incoming fibres (Teo and Ramakrishna, 2006). However, the thickness of these scaffolds can be controlled by varying the electrospinning time. The thickness and architectural structures influence the mechanical properties of the electrospun scaffolds. Electrospun nanofibre scaffolds satisfy the requirements of regeneration, enabling the fabrication of living cell/scaffold hybrids with adjustable thickness, with the potential to provide benefits such as convenient handling and allowing integration in the recipient site, a prerequisite for successful outcome in cell-based periodontal ligament tissue engineering (Inanc et al., 2009).

10 µm (a)

(b)

13.2 (a) SEM micrographs of aligned electrospun polycaprolactone (PCL) scaffolds for human Schwann cells culture and (b) confocal fluorescent images of human Schwann cells cultured on aligned PCL scaffolds for 3 days. Solvent, dichloromethane and methanol (ratio, 8:2); flow rate, 4 ml h–1; electric field, 7 kV; distance between supply and collector, 5 cm (adapted from Chew et al., 2008).



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Shang et al. (2010) reported periodontal ligaments cells cultured on electrospun scaffolds compared with cast films and analysed the degradation rate, morphology, mechanical properties and cell proliferation. The degradation results indicated that electrospun fibres degrade much slower than films. The high rate of degradation of films was due to the slow diffusion rate of degraded products containing acid-ending groups, whereas, in electrospun scaffolds, the interconnected porous structure makes it easier for degraded oligomers to diffuse out. For periodontal regeneration, fast degradation of scaffolds reduces its presence in periodontal space. Consequently, the bone fill, tissue coverage and periodontal regeneration achieved will be limited. The orientation of fibres also helped in the cell growth. The aligned fibres promoted directed cell movement; in contrast, the random oriented fibres hampered the migrational behaviour of cells. The aligned fibres can guide cell invasion and provide a highway for cell migration. The key factor in achieving periodontal tissue regeneration is fast migration and proliferation of periodontal ligament cells into the periodontal defect space.

13.3

Reinforcement of dental restorations

Current state-of-the-art dental composites contain modified acrylate resin matrices filled with micro-/nano-scaled ceramic particles. It has been shown that the performance of dental composites can be improved through the use of nanotechnology (Moszner and Klapdohr, 2004). Significant improvements have been made in the microstructure and performance of resin composites, resulting in an enhanced resistance to wear and fracture (Leinfelder, 1995; Suzuki et al., 1996; Ferracane et al., 1997). These include improved filler packing, higher filler levels, and hybrid filler phases (Bowen et al., 1992; Ferracane, 1992; Anusavice, 1996). However, the toughness, strength and durability of resin composites are still inadequate for their use to be expanded to large stress-bearing posterior restorations that involve the replacement of cusps (Wilder et al., 1996). The fillers used to reinforce dental resins are mostly silicate glasses and occasionally glass ceramics. Glass particles are considerably harder than the resin matrices in which they are embedded; much of the stress is transmitted through the particles into the resins. Wherever the submerged portions of the particles are angular or irregular in shape, the stress concentration could become excessively high. Such a condition tends to generate small cracks around the particles, thereby weakening the matrices locally. Chopped glass fibres (with diameters of 10–15 mm and aspect ratios larger than 100) have also been used as fillers for dental resins (Krause et al., 1989; Willems et al., 1992). The resultant composites showed higher mechanical properties, but the improvement was limited probably owing to weak interfacial bonding. The electrospun nano-scaled glass fibres are



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expected to substantially outperform micrometre-scaled glass particles/fibres in the reinforcement of dental composites. This is because the nano-scaled glass fibres possess the desired morphological and mechanical properties including small fibre diameter (and the concomitant large specific surface area), large aspect ratio and high strength and modulus (Dodiuk-Kenig et al., 2008). If a micro-crack is initiated in the dental matrix under contact wear and/or other stresses, the nanoscaled glass fibres remain intact across the crack planes, supporting the applied load. Therefore, crack opening is resisted by the fibres and the matrix is reinforced. Compared to the micrometre-scaled glass fibres, the electrospun nanoscaled glass fibres are over ten times thinner (i.e. their specific surface areas are more than ten times larger) and possess abundant surface silanol (Si—OH) groups that can readily react with silane coupling agents such as 3-methacryloxypropyltrimethoxy silane (MPTMS). Therefore, the interfacial bonding between the surface-silanised electrospun nano-scaled glass fibre and the dental resin matrix can be extremely strong. Among electrospun ceramic nanofibres, silica (SiO2) nanofibres have been investigated with particular interest because of their well-studied precursors and wide dental applications (Shao et al., 2002; Choi et al., 2003). Electrospun nanofibres are low cost and continuous nanofibres are relatively easy to process into applications (Dzenis, 2004; Fong, 2007). The electrospinning method for fabricating nanofibres from polymer blend has been widely reported, but only a few are aimed at dental restorative materials (Fong, 2004; Tian et al., 2007; Gao et al., 2008; Lin et al., 2008; Khan et al., 2008; Sun et al., 2010). Gao et al. (2008) prepared electrospun nano-scaled glass fibres using a spin-dope consisting of 13% tetraethyl orthosilicate (TEOS) and 12% polyvinylpyrrolidone (PVP) in a solvent mixture of N,N-dimethylformamide (DMF)/acetic acid (HAc) and using pyrolysis temperature of 800°C. The prepared fibres were dispersed in cyclohexane and then ultrasonic vibration (Fig. 13.3) converted the continuous fibres into nano-scaled short fibres and mixed surface silanised glass fibres with dental resins (2,2’-bis-(4-2-hydroxy-3-methacryloxypropoxy)-phenyl)-propane (BisGMA) and tri (ethylene glycol) dimethacrylate (TEGDMA), camphorquinone (CQ) and ethyl (4-dimethylamino) benzoate (EDMAB)). It was found that the flexural strength, elastic modulus and work of fracture values of the composite reinforced by silanised electrospun nanofibres were higher than those of composites reinforced by silanised glass powders. The fibre orientation, morphology and distribution strongly influenced the composite properties. Unidirectional longitudinal fibre reinforced the composite to the maximum strength when stress is exerted along the direction of the fibre. The strength of the composite reduces when the stress is applied at an angle to the direction of the fibre. Therefore unidirectional fibres have significantly greater strength than bidirectional



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5 µm

13.3 SEM image showing electrospun nano-scaled glass fibres after ultrasonic vibration, solvent was dimethylformamide and acetic acid (DMF/HAc) (volume ratio of 15/1), using a pyrolysis temperature of 800°C (adapted from Gao et al., 2008).

fibres (Chang and Chai, 2003). The relationship between the quantity of fibre and the polymeric matrix has been demonstrated; with an increase in the quantity of fibre, there is increase in flexural strength (Behr et al., 2000). The distribution of fibres exhibits different properties, depending on their application. If the fibres are equally distributed, the fatigue resistance is enhanced but if they are located in one place they can increase the stiffness and strength (Narva et al., 2001). The incorporation of nanofibres into paste substantially increased the viscosity of dental resins. The nanoscaled glass fibre contained dental composite with rough but no clearly identifiable fracture lines. This suggested that the presence of electrospun nanofibres could deflect the microcracks and effectively increase resistance to the applied force. Wen et al. (2010) controlled the surface roughness of these electrospun fibres by incorporating SiO2 nanoparticles. These nanofibres were prepared by selecting tetraethyl orthosilicate (TEOS) as alkoxide precursor, polyvinylpyrrolidone (PVP) as carrying polymer, N,Ndimethylformamide (DMF)/acetic acid (HAc) as mixture solvent, with a pyrolysis temperature between 600°C and 1000°C. Their results indicated that SiO2 nanofibres were structurally amorphous and demonstrated higher morphological consistency (500 nm) (Fig. 13.4) and SiO 2 nanoparticles appeared to be enriched on the fibre surface. These nanofibres are expected to find important applications in composites (particularly dental composites) as well as catalyst support and adsorption. Fong (2004) investigated dental polymers, Bis-GMA and TEGDMA under the reinforcing effect of electrospun nylon-6 nanofibres. The small diameter of nanofibres provided a high surface area to volume ratio, which



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Dental regeneration (a)

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13.4 SEM images of SiO2 electrospun nanofibres made from spin dopes containing varied amounts of SiO2 nanoparticles: (a) 0 wt%, (b) 1.47 wt%, (c) 2.90 wt% and (d) 4.29 wt%. Electric field, 15 kV; flow rate, 1.0 ml h–1; distance to collector, 10 inches (~25 cm) (adapted from Wen et al., 2010).

10000x 15kV

1 µm

10000x 15kV

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13.5 Scanning electron micrographs of (a) electrospun nylon-6 nanofibres in the form of nonwoven fabric, (b) representative fracture surface showing the presence of the nanofibres in the composite resin. The solvent, 1,1,1,3,3,3-hexafluoro-2-propanol (HFP); electric field, 20 kV; distance from supply, 25 cm (adapted from Fong, 2004).

could enhance the intermolecular hydrogen bonding between the nylon-6 nanofibres filler and the matrix of resin polymers as shown in Fig. 13.5. Furthermore, electrospun nanofibres are continuous (Fig. 13.6(a)). If a microcrack is initiated in a matrix under contact wear and/or other stresses,



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Resin remnants 500 nm

500 nm

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13.6 SEM images showing: (a) neat nylon-6 nanofibres and (b) nylon-6/fibrillar silicate nanocomposite nanofibres on the fracture surfaces of three-point flexural testing specimens. Solvent, 1,1,1,3,3,3-hexafluoro-2-propanol (HPF); electric field, 25 kV (adapted from Tian et al., 2007).

10 µm

13.7 SEM image of as-electrospun PAN–PMMA continuous nanofibres. Solvent dimethylformamide (DMF, 10%); electric field, 15 kV, flow rate, 0.4 ml h–1 distance from nozzle to collector, 15 cm (adapted from Sun et al., 2010).

the nanofibres remain intact across the crack planes and support the applied load. Therefore, crack opening is resisted by the nanofibres and the matrix is reinforced as shown in Fig. 13.6(b) (Tian et al., 2007). Sun et al. (2010) successfully fabricated polyacrylonitrile (PAN)– poly(methyl methacrylate) (PMMA) core–shell non-woven nanofibrous membranes, as shown in Fig. 13.7, by co-electrospinning a mixed solution of these polymers from a single-nozzle. The non-woven PAN–PMMA nanofibre



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membranes were laminated on Bis-GMA/TEGDMA (50/50 wt%) resin with photoinitiator CQ and co-initiator 2-(dimethylamino)ethyl methacrylate (DMAEMA) layer by layer, which resulted in the flexural strength (Fs), elastic modulus (Ey) and work of fracture (WOF) increasing by 18.7%, 14.1% and 64.8%, respectively. The SEM images of the fracture surface of the Bis-GMA/TEGDMA composite reinforced with 1.2 wt% PAN-PMMA electrospun nanofibres are shown in Fig. 13.8. The research group hypothesised using this material as dental restorative composite. They owed the improvement effect to the formation of a semi-interpenetrated polymer network (IPN) structure between the nanofibre and the dental resins, which resulted in an ‘in situ nano-interface’ that could provide good interfacial adhesion. A novel bioactive dental restorative composite was synthesised by chemically binding the nano-hydroxyapatite (nHA) to the diisocyanate component in the polyurethane (PU) backbone utilising solvent polymerisation. The procedure involved stepwise addition of monomeric units of the PU and optimising the reagent concentrations (Khan et al., 2008). The interface adhesion of nHA particles and the polymer matrix is a major factor affecting the properties of the composites (Liu et al., 1996; Hong et al., 2005). When nanoparticles (nHA) and polymers form a composite, provided that homogeneous dispersion of the nanoparticles is achieved at the microscopic level, the resultant materials properties are expected to be improved and are an effective way to manipulate the surface properties of nHA. In the study, the selected PU/nHA composites were dissolved in N,N-dimethylformamide (DMF) to form a 10 wt% solution. The composite solution was electrospun to produce nanofibres for morphological characterisation. The SEM images

5 µm

2 µm (a)

(b)

13.8 SEM images of the fracture surface of the Bis-GMA/TEGDMA composite reinforced by 1.2 wt% PAN-PMMA electrospun nanofibres (a) perpendicular to the direction of the nanofibres and (b) parallel to the direction of the nanofibres. Solvent, dimethylformamide (DMF, 10%); electric field 15 kV; flow rate, 0.4 ml h–1; distance from nozzle to collector, 15 cm (adapted from Sun et al., 2010).



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of electrospun mats of PU and PU/nHA shown in Figs 13.9, 13.10 and 13.11. The morphological appearance of fibres showed quite uniform size (400–500 nm). The reduction in diameter from the millimetre scale to the

1 µm

13.9 SEM image of electrospun fibres of polyurethane, 10 wt% DMF was used as solvent. Diameter of nozzle, 330 µm; flow rate, 15 ml min–1; applied voltage, 7–9 kV. Fibres were collected 120 mm below the nozzle.

10 µm

13.10 SEM image of PU/nHA composite. Solvent, 10 wt% DMF; diameter of nozzle, 330 µm; flow rate, 15 ml min–1; applied voltage, 7–9 kV. Fibres were collected 120 mm below the nozzle.



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Spectrum 4

O Ca O

Si P 0

1

Ca Ca 4 5 keV

2 3

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13.11 SEM image and EDS spectrum of electrospun fibres of PU/nHA composite. Solvent, 10 wt% DMF; diameter of nozzle, 330 µm; flow rate, 15 ml min–1; applied voltage, 7–9 kV. Fibres were collected 120 mm below the nozzle.

nanoscale was due to stretching of jet by whipping and bending. The size of nanofibres depends on the incorporation of nHA which causes the viscosity of the solution to increase. The increase in viscosity of the solution is due to the increased molecular entanglement or linkage between polyurethane and nanoparticles. The smoothness and splaying of nanofibres was attributed to the increasing DMF volume fraction. Surface tension and viscosity decreased, while conductivity and dielectric constant increased. The solvent composition and various solution properties conducted an important role in determining fibre formation (Lee et al., 2003). The SEM image of PU (Fig. 13.9) shows the presence of beads in nanofibres. It is reported that the electrospun beaded nanofibres of the polymers were related to the instability of the jet of polymer solution (Lee et al., 2003; Berkland et al., 2004). Whereas, the morphological analysis of PU/nHA composite (Fig. 13.10) shows non-beaded uniform nanofibres. The energy dispersive spectroscopic (EDS) analysis shows (Fig. 13.11) the presence of calcium and phosphorous peaks attributed to the presence of homogeneously mixed nHA. The possible application of these resultant composite nanofibres is to reinforce the dental resins by cutting into short fibres and increasing the mechanical properties such as compressive strength, flexural strength and physical properties such as wear and fracture toughness. The other hypothetical application of these nanofibres is to use them as



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scaffolds for dentino-pulpal regeneration. To date electrospun fibres have not been used as scaffolds for dentino-pulpal regeneration. However, it is expected that electrospun fibres have a potential application for use as scaffolds for tooth regeneration. The mechanisms underlying tooth tissue engineering are largely unknown and further studies into the function of scaffolds and precursor cells are required.

13.4


In recent years electrospinning has gained widespread interest for its potential applications in tissue engineering. This renewed interest can be attributed to its relative ease of use, adaptability and its ability to fabricate fibres with diameters on the nano-scale. While much has been learned about the electrospinning process over its long history, researchers are still gaining new insights and developing new ways of utilising this technique for tooth regeneration applications. The desired scaffold properties such as pore size, density, geometry and alignment can be used to control the mechanical properties as well as the biological response to the scaffold. Research has shown that surface modification of electrospun fibres can be used to enhance further the scaffold’s interaction with cells in tissue engineering applications. The ability to electrospin synthetic (biodegradable or non-degradable), natural and hybrid materials allows precise tailoring of the scaffold properties to the desired application and new materials are constantly being electrospun. In this chapter, recent and state-of-the-art work in electrospinning for applications in dento-oral regeneration and restoration have been discussed. In the future it is important that research focuses on gaining a better fundamental understanding of the electrospinning process, but even more importantly on how this technique can be used as a tool in developing new systems for dento-oral regeneration. Electrospun nanofibres could be utilised to develop the next generation of dental composites, which would be particularly useful for large posterior restorations. If the fibres can be collected individually and aligned to specific orientations, it is possible to increase the maximum volume fraction of fibres in a resin matrix thus increasing the strength of the material. The increased surface area/weight will also allow better bonding with the resin matrix to help prevent failure by pullout of the fibres. Much work is still going on and there is yet to be a definite scaffold for tissue engineering. Electrospun nanofibrous scaffolds hold great potential for use in dento-oral regeneration applications in the future. A nanofibrous scaffold closely mimics the extracellular matrix rather than macroscale matrix. Cell adhesion, proliferation and differentiation of several types of cells have been observed and can be used for osteogenesis. Future developments and processing technologies must find a way of advancing in the right direction.



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References

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A. S u b r a m a n i a n, U. M. K r i s h n a n, and S. S e t h u r a m a n, SASTRA University, India

Abstract: Skin regeneration using electrospun nanofibres has received a great deal of attention for treating patients suffering from acute burn injuries and chronic non-healing wounds like diabetic ulcers, pressure ulcers, venous ulcers and so on. Currently available skin substitutes which restores the tissue function have had limited success. Tissue engineering has been exploited to design a scaffold which mimics the extracellular matrix (ECM) in its structure and function. Recently electrospinning has emerged as an important technique for fabricating these scaffolds because of its unique ability to generate nanofibres with appropriate physical, mechanical and biological properties to support cell growth, proliferation, migration and differentiation. This chapter focuses upon the potential of electrospun nanofibres as skin equivalent. Key words: biomaterials, dermis, electrospinning, epidermis, fibroblast, keratinocytes.

14.1

Introduction

Successful development of skin equivalents has considerable implications for patients suffering from acute burn injuries and chronic non-healing wounds. Skin, the largest organ in the body, performs intricate functions for maintaining the structural integrity of the entire body, thermoregulation, a protective barrier against microorganisms, fluid homeostasis, immunoactive defence, sensory function and vitamin-D production through UV exposure when in direct sunlight (Soh et al., 2009). Additionally, the selfhealing property of skin tissue promotes the repair of wounds. However, when extensive skin loss occurs, owing to deep burns or diabetic ulcers, the skin cannot repair the wound, which subsequently causes loss of function (MacNeil, 2008). This necessitates the need for the fabrication of a skin equivalent, which covers the wound and regenerates the native tissue to restore all of its functions. Tissue engineering is a novel approach for rebuilding the native tissue with its functions through the development of an extracellular matrix (ECM) analogue. Therefore, skin tissue engineering has received immense attention for its ability to restore the tissue function in two distinct types of injuries, namely acute wounds such as burn injuries and chronic non-healing wounds such as diabetic ulcers, pressure ulcers and so on. Currently, there are enormous 298 © Woodhead Publishing Limited, 2011


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numbers of skin substitutes existing in the market for clinical use. But each restores the function with limited success since these skin substitutes do not mimic the ECM topography (Metcalfe and Ferguson, 2007). Recently, research has been focused on the development of nanofibres as an ECM analogue for tissue engineering applications. There are many different methods such as self-assembly, template-directed synthesis, phase separation, melt-blown, drawing, electrospinning and so on, used to fabricate nanofibres (Jian et al., 2008). Among these, electrospinning has been extensively used to develop micro/nanofibres from naturally derived and synthetic polymers. This chapter outlines the challenges in engineering skin tissue, the development and future prospects of electrospun nanofibres as a skin equivalent and the potential of electrospun nanofibres for skin regeneration compared to other skin substitutes.

14.2

Biology of skin and wound healing

A complete understanding of skin structures with their functional relationships is mandatory in developing tissue engineered skin. Normal skin is composed of layers namely epidermis, dermis and hypodermis (Lee, 2000). The histology of native skin for a Wistar rat is shown in Fig. 14.1. The epidermis is the thin outer covering of skin (0.1–0.2 mm in depth) and is comprised of different cell types, such as self-renewing keratinocytes which perform a barrier function (MacNeil, 2008; Lee, 2000), melanocytes for pigmentation (Lee, 2000), adnexal cells which include sebaceous glands, sweat glands, and hair and nails, dendritic cells of the immune system (Lee, 2000) and nerve endings for sensory function. The keratinocytes in the basal layer of the epidermis are present mainly to repopulate the lost epithelial cells, thereby forming stratified squamous epithelium (Brouard and Barrandon,

Epidermis Sebaceous gland Dermis Hair follicle

14.1 Histology image of skin layers from rat species (¥40 magnification).



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2003). Additionally, this outer layer, comprising limited extracellular matrix molecules especially carbohydrate polymers and lipids, provides a barrier against aqueous fluids. Keratin filaments within the epidermis help to maintain its structural integrity (MacNeil, 2008). In contrast, the underlying dermis is composed of two regions namely papillary and reticular dermis. The superficial region of dermis called the papillary dermis has small bundles of loose connective tissue with collagen and immature elastic fibres (Keene et al., 1997). Within the reticular dermis, few cells are present and a dense network of extracellular matrix molecules, predominantly collagen and elastin fibres, and glycosaminoglycans confer mechanical strength, flexibility and elasticity (Metcalfe and Ferguson, 2007). Fibroblasts are the major cell type within both regions of dermis and are responsible for the production of ECM molecules (Lee, 2000). Moreover, these fibroblast cells are found to play a major role in the wound healing process by producing remodelling enzymes such as proteases and collagenases (Metcalfe and Ferguson, 2007). Other cells such as endothelial cells, nerve cells and mast cells are also present in both the regions of dermis. The third layer is the hypodermis, which is rich in adipose tissue and blood vessels, and provides energy and nutrition to the cells within the epidermis and dermis (Andreadis and Geer, 2006), mechanical protection and thermoregulation. Since fat is a heat insulator, this layer insulates the body from cold temperatures and offers shock absorption. In general, skin is continually renewed and remodelled by equilibrating the proliferation, differentiation and apoptosis of the cells (Brouard and Barrandon, 2003). Once every three weeks, the stratified squamous epithelium in the epidermis is completely renewed and similarly the dermis of skin is routinely remodelled with ECM molecules (Brouard and Barrandon, 2003). As previously mentioned, skin tissue possesses a self-wound healing property. Wound healing is a sequence of complex biochemical events comprising three phases: inflammation, proliferation and remodelling (Carlson et al., 2009). Immediately after injury, coagulation of the platelets occurs to arrest bleeding by adhesion and aggregation of platelets to interstitial connective tissue. The inflammatory cytokines released at the site of injury to regulate the blood flow cause lymphocytes and macrophages to infiltrate and thus remove bacteria and debris (Metcalfe and Ferguson, 2007). Angiogenesis and collagen deposition in the proliferative phase promotes granular tissue formation (Clark, 1997; Clark and Singer, 2000). In parallel with granulation, wound contraction is facilitated by myofibroblasts (actin rich fibroblast). Growth factors such as platelet-derived growth factor (PDGF) and transforming growth factor b (TGF-b) recruit myofibroblast that slowly replaces the granulation tissue in the wound bed. Collagenase and matrix metalloproteinase (MMP- 1,2,3, 8 and 13) play a major role in collagen remodelling by removing excess collagen formed during continuous collagen synthesis, thus controlling the



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overall amount of collagen (Koga et al., 2007). Finally the proliferation and migration of keratinocytes into the wound bed regenerates the epidermal layer (Clark and Singer, 2000). However, difficulty remains in the self-healing of larger and more chronic wounds, which may be due to the imbalance of these factors (Parenteau et al., 1997). It is clear that the wound healing mechanism depends solely upon communication between the dermal fibroblasts and epidermal keratinocytes, extracellular matrix molecules and vascularisation, which is regulated by the release of inflammatory cytokines and growth factors such as platelet-derived growth factor (PDGF), transforming growth factor (TGF-b), epidermal growth factor (EGF) and vascular endothelial growth factor (VEGF). Therefore for active skin tissue regeneration, the scaffold should ideally be able to interact well with both the fibroblasts and keratinocytes, thereby depositing appropriate ECM molecules on the scaffold surface, which promotes angiogenesis for tissue ingrowth and stimulates the expression of other signals such as cytokines, growth factors and so on.

14.3

Challenging problems in existing therapies

The current gold standard for skin regeneration is the use of autologous grafts for the permanent coverage of both acute and chronic wounds (Kumbar et al., 2008). For partial thickness burn injuries, a split-thickness skin graft is used to cover the damaged area. A split-thickness graft is normally composed of the entire epidermal layer with some of the underlying dermal layer (MacNeil, 2008), whereas a full thickness skin graft contains both the epidermis and dermis layers. Although the full-thickness graft provides a better performance than the split-thickness graft, it requires a well-vascularised wound bed. The improved performance of such grafts is chiefly due to the migration of epidermal keratinocytes from the basal layer, thereby promoting re-epithelialisation of the wound (MacNeil, 2008). However, the use of autografts is limited by their availability, blister formation, donor-site morbidity, secondary surgery and chronic inflammation (Eaglstein and Falanga, 1997). Alternative treatments for acute or chronic wounds are the use of allogenic cadaver skin grafts or xenografts as a temporary covering to protect the wound (Eaglstein and Falanga, 1997). The major limitations of allografts are immunologic rejection, the requirement for immunosuppressive treatment and possible transmission of infection, such as HIV and hepatitis (Eaglstein and Falanga, 1997). The use of a porcine xenograft, for example, also had further limitations, including a poor shelf life and bacterial contamination (Eaglstein and Falanga, 1997). Cultured ‘sheet grafts’ or cultured epithelial autografts have been developed by growing multilayers of epithelial sheets directly from keratinocytes isolated from the patient as a potential substitute for the irreparable skin (Eaglstein and Falanga, 1997). However, the risk of



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bacterial infection, high cost, difficulty in handling fragile sheets of cells and poor stability limits the use of a sheets as a potential graft (Eaglstein and Falanga, 1997). Currently, there are plenty of temporary or permanent skin equivalents that are commercially available. These are made of living cells, ECM molecules, biomaterials or a combination of all these. Epicel® is an epidermal autograft used for permanent wound coverage that poses a minimum risk of infection. However Epicel® requires 2 to 3 weeks to culture and its fragile nature limits its use (Lee, 2000). Composite graft Integra® is a bilayered artificial skin composed of a disposable silicone sheet as an artificial epidermis bonded to glutaraldehyde cross-linked porous collagen (bovine) – chondroitin (shark) as a dermal layer (Andreadis, 2006). Although it resembles many characteristics of human skin, the risk of infection, haematoma formation and two-step processes (replacement of silicone layer by split thickness autograft) restrict the use of this graft (Lee, 2000; Eaglstein and Falanga, 1997). Alloderm® is composed of human cadaveric skin, forming an acellular dermal matrix; however, it has a major limitation owing to the possibility of disease transmission (Lee, 2000). Apligraf® is the best skin equivalent that closely mimics the native tissue and has the ability to stimulate growth factors and cytokines similar to human skin (Eaglstein and Falanga, 1997). It is made up of human keratinocytes cultured on a matrix of bovine collagen type I and human dermal fibroblast. However, the major limitations of these grafts are expense, limited shelf life and time-consuming process to manufacture which reduces their effectiveness (Lee, 2000). It is apparent that although the available skin substitutes seem to be a promising option for the treatment of both acute and chronic wounds, each one has its own limitations. This is mainly because of their lack of topographical homogeneity when compared to the native tissue. The morphology of native ECM is fibrillar in the size range micrometre to submicrometre, with collagen fibrils in the range of 50–500 nm (Rho et al., 2006; Subramanian et al., 2009). Thus it is desirable for a tissue engineered scaffold to mimic the extracellular matrix in terms of its topography, dimensions, composition and mechanical properties. Additionally, it would be advantageous if the scaffold has antimicrobial properties to prevent bacterial infection, wound healing properties to repair and regenerate the skin and thrombogenic properties to establish haemostasis.

14.4

Restoring functional skin tissue

Engineering skin tissue is mainly focused on the restoration of dermis and epidermis layers rather than the regeneration of skin appendages (such as hair, nail and glands). Various strategies have been attempted to re-establish the native skin without compromising its functions. Bioengineers have aimed



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to enhance the wound healing process in the native skin tissue by delivering drugs such as antibiotics to control infection and growth factors/genes to speed up the re-epithelialisation, angiogenesis and/or granulation tissue phases (Andreadis and Geer, 2006). Other researchers may choose to regenerate the epidermal or both epidermal and dermal layers by implanting acellular or cellular matrix. Moreover, by surface modification of the scaffold, researchers have improved the biological properties of tissue engineered scaffolds such as cell adhesion, proliferation, viability and differentiation. Complete restoration of functional skin tissue chiefly depends upon the vital determinants such as the choice of biomaterial, scaffold fabrication methodology, cell lines and surface topography such as pores, ridges, grooves, fibres, nodes, and so on (Subramanian et al., 2009).

14.5

Nanofibres as extracellular matrix analogue

The ECM plays a significant role in organ development and also in tissue repair under pathological conditions (Kolacna et al., 2007). The ECM architecture and composition are responsible for cell integrity, migration, adhesion, nutrition, differentiation and angiogenesis. The most abundant proteins in the ECM are the collagen family, which provides the biological and structural integrity of many different tissues. The fibre diameter of different collagen types varies depending on the type of subunits; for example, type I collagen forms fibres of 50 nm; type II collagen develops fibres of less than 80 nm; type III collagen exists as 30–130 nm fibrils (Kolacna et al., 2007). The main goal of tissue engineers is to design ECM analogues, incorporating three principle characteristics, namely similar physical properties such as mechanical stiffness and elasticity, the same chemical composition as the ECM and nanoscale geometry similar to the collagen fibrils (Mark et al., 2010; Li and Xia, 2004). An electrospun nanofibrous scaffold seems to be a potential candidate to mimic the architecture of native ECM (Han et al., 2008). The high surface area to volume ratio, high porosity with tunable pore size, tailorable mechanical strength and morphological similarities to native collagen fibrils make nanofibres suitable candidates for ECM analogues (An et al., 2009; Taepaiboon et al., 2007). The high surface area to volume ratio of electrospun scaffolds showed higher adsorption of cell adhesion molecules such as vitronectin and fibronectin than the other scaffolds that control the cell fate (Woo et al., 2003). A higher porosity within the electrospun nanofibrous scaffold could easily allow the diffusion of nutrients, by-products and tissue in-growth (Ashammakhi et al., 2007; Hromadka et al., 2008). Moreover, spinning biodegradable polymers directly onto the site of injury can promote wound healing by the formation of skin tissue and remove scar tissue formation (Xiumei et al., 2007; Huang et al., 2003). The electrospun nanofibrous scaffold should be able to perform the functions



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of ECM such as physical support for cell adhesion, proliferation, migration and differentiation, rigidity and mechanical stiffness to the engineered tissues, active cell–scaffold interaction to control cell fate by surface modification, a reservoir for exogenous growth factors and porous structure for vascularisation and neo tissue formation (Pan et al., 2006; Yoo et al., 2009; Maretschek et al., 2008; Li et al., 2002). The nanofibrous topography could permit the filopodia of cells to adhere firmly through increased serum protein adsorption, thereby contributing better cell adhesive strength than solid walled scaffolds (Woo et al., 2003). Li et al. (2006) reported that chondrocytes seeded onto microfibres dedifferentiated to fibroblast-like morphologies whereas chondrocytes on nanofibrous scaffolds retained chondrocyte-like morphology (Li et al., 2006). It was also hypothesised that the size of the nanofibres are several orders less than normal cells, resembling native ECM architecture, and possess spatial advantages in promoting cell–matrix interaction. However, scaffolds made from microfibres does not possess such merit. Nanofibrous cultures were found to promote the formation of 3D matrix adhesion and an in vivo-like microenvironment for both fibroblast and kidney cells (Schindler et al., 2005).

14.6

Ideal properties of scaffold

There are numerous criteria for the design of a scaffold to promote wound healing and skin regeneration without the formation of scar tissue. Ideally, scaffold for skin tissue should have appropriate wettability, antimicrobial properties, mechanical properties, biodegradability, surgical handleability, wound healing properties and cell adhesive properties. The electrospun nanofibre matrix seems to be a promising candidate for skin tissue engineering applications because of its easily tunable scaffold properties, such as wettability, biodegradability, porosity and mechanical properties by fibre dimension (Kumbar et al., 2008). The high surface area to volume ratio of electrospun nanofibres provides an excellent barrier whilst preventing fluid accumulation and facilitating oxygen permeation at the wound site (Kumbar et al., 2008).

14.6.1 Wettability Major problems in acute and chronic wounds are infection and fluid accumulation at the tissue interface. Therefore, it is desirable that the scaffold should readily adhere to the wound site, mainly to relieve pain and fluid accumulation. Scaffolds made from hydrophilic materials are of considerable interest in skin regeneration, since they exhibit good adherence properties to the wound and can also create hypoxic environments to promote wound healing (Kane et al., 1996). Hypoxia contributes to the upregulation of



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VEGF expression in cells such as monocytes, fibroblast, keratinocytes and endothelial cells by hypoxia inducible factor (HIF-1a). This in turn promotes angiogenesis and tissue perfusion (Bao et al., 2009). Moreover, cells prefer a hydrophilic environment to adhere, spread and grow (Kumbar et al., 2008). Recently chitosan nanofibrous scaffolds have received a great deal of attention for skin regeneration since their haemostatic properties enhance their adherence to the wound bed (Baldrick, 2010). The fibre diameter has been found to affect the surface wettability of the scaffold, that is as the fibre diameter decreases, the hydrophilicity also decreases (Kumbar et al., 2008). This is because the smaller fibre diameter demonstrates evidence of high degrees of roughness, since surface roughness is synonymous with the intrinsic non-wetting characteristic of the substrate (Robinette and Palmese, 2005).

14.6.2 Antimicrobial properties The barrier function of a nanofibrous scaffold is essential to protect the wound from bacterial infection, fluid accumulation and to prevent water and electrolyte loss. Recent research has focused on the development of nanofibres with antimicrobial properties. Specifically, chitosan/PEO ultrafine nanofibres were found to have strong antibacterial activity, which was further enhanced by the addition of less than 5 nm silver nanoparticles to the nanofibres (An et al., 2009). It was observed that although the biocidal property of the scaffold depends on the presence of silver nanoparticles, the developed nanofibre provides more surface area where silver nanoparticles can reside, hence better microbicidal activity (Xu and Zhou, 2008). Further information on antimicrobial wound dressings is included in Chapter 15.

14.6.3 Mechanical properties The choice of scaffold mainly depends upon the type of tissue. Some tissues require a soft, spongy scaffold whereas some require a rigid scaffold. Depending on the type of tissue, both rigid (causes damage to surrounding tissues) and fragile (loses mechanical integrity) scaffolds lead to inflammatory responses, resulting in subsequent implant failure (Greenwald and Berry, 2000). Thus the mechano-compatibility of the scaffold plays a major role in tissue regeneration. An ideal skin scaffold should be flexible, elastic and durable and should resemble the mechanical properties of native skin (Quinn et al., 1985). The major advantage of using electrospun nanofibrous scaffolds for tissue engineering applications is that the mechanical properties can be tailored by varying the fibre diameter and composition of the scaffold. The tensile properties of fibrous scaffolds increases with decreasing fibre diameter (Xie and Wang, 2006). Kumbar et al. (2008) developed poly(lactic-



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co-glycolic acid) (PLGA) nanofibrous scaffolds for skin regeneration with various fibre diameters (150–225 nm, 200–300 nm, 250–467 nm, 500–900 nm, 600–1200 nm, 2500–3000 nm); the tensile modulus was found to vary from 39.23(±8.15) to 79.21(±13.71) MPa which was comparable to normal human skin (15–150 MPa) (Kumbar et al., 2008). Our laboratory has successfully improved the mechanical properties of chitosan nanofibrous scaffolds by incorporating gelatin for skin tissue engineering applications (Brahatheeswaran et al., 2010). Scanning electron micrographs of chitosan and chitosan–gelatin scaffold are shown in Fig. 14.2. The ultimate tensile strength of the chitosan nanofibrous scaffold was measured as 13.61(±0.6) MPa, whereas the maximum tensile strength for chitosan–gelatin (50:50 blend) nanofibre scaffold was increased to 37.91(±4.42) MPa (Brahatheeswaran et al., 2010).

14.6.4 Porosity The porosity of the scaffold is essential to promote vascularisation, cell infiltration, exchange of nutrients and tissue ingrowth (Wu and Ding, 2004). The minimum pore size required for the scaffold to facilitate cell migration is 10 mm (Kane et al., 1996). For skin regeneration, electrospun nanofibrous scaffolds have received considerable interest compared to freeze dried scaffolds, since in the freeze drying process, pore structure and its size mainly rely upon the nucleation and growth rate of ice crystals (Powell et al., 2008). Thus the freeze drying process generates a wider distribution of pore diameters with heterogeneous pore structures throughout the scaffold (Powell et al., 2008; O’Brien et al., 2004). Electrospun collagen nanofibres were found to possess a more homogeneous pore structure with a small variation in pore diameter (0.13–4.5mm) compared to freeze dried collagen sponge, (a)

(b)

10 µm

10 µm

14.2 Scanning electron micrographs of (a) chitosan and (b) chitosan– gelatin nanofibrous scaffolds at a magnification of 1000 times using constant electrospinning parameters; voltage gradient 2.5 kV cm–1 and flow rate 0.005 ml min–1.



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which had a larger variation in pore structure and pore diameter (1–10 mm) throughout the scaffold (Powell et al., 2008). Although a nanofibrous mat has a very large surface to contact with cells, some reports demonstrated its limitations in cell infiltration owing to its small pore size. The porosity of the electrospun fibrous mats enhances nutritional transport. However, there is a theory that the cells can ‘push’ the fibres aside and penetrate through. In addition, electrospinning parameters can be controlled in order to engineer a desired pore size, which could aid cell infiltration and vascularisation since cell infiltration depends upon the cell type.

14.6.5 Biodegradability The biodegradable properties of the scaffold plays a major role in the tissue engineering process. The rate of degradation and its degradation products have an effect on the cell viability, growth and the immune response (Wu and Ding, 2004). In order to maintain the structural integrity and mechanical properties of the scaffold, it is desirable to control its degradation rate (Pham et al., 2006). It is also essential to balance the degradation rate of the scaffold with the rate of new tissue formation and regeneration of neo tissue (Pham et al., 2006).

14.6.6 Surgical handleability The mechanical properties and handleability of the scaffold are very important for in vivo applications as they need to resist structural collapse during implantation. Thus it is advantageous for the scaffold to be able to withstand sterilisation and to be suturable (Kane et al., 1996). Nanofibrous scaffolds can be sterilised by ethylene oxide, g-radiation, UV radiation or 70% ethanol (Vepari and Kaplan, 2007).

14.6.7 Wound healing properties The wound healing properties of the skin substitute have been enhanced mainly by three approaches: antimicrobial actions to protect the wound from infection, activation of cytokines and other tissue factors, and finally imparting thrombogenic and platelet adhesive properties to control bleeding at the wound site. Chitosan nanofibrous scaffolds have been found to regenerate the skin tissue whilst promoting wound healing. This is due to the thrombogenic and antimicrobial properties of chitosan (Baldrick, 2010). Moreover, the free amine groups in chitosan activate macrophages and cytokines in mice, demonstrating its role in the wound healing process (Baldrick, 2010).



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14.6.8 Cell adhesive properties A chief property of the scaffold is that it should allow the cells to adhere to it and this mainly depends on the surface characteristics such as the charge density and wettability (Subramanian et al., 2009). The cell adhesive properties of the nanofibrous scaffold can also be improved by a coating of ECM proteins such as laminin, fibronectin and collagen or by the incorporation of peptide sequences such as arginine–glycine–aspartic acid (RGD) or by functional groups such as hydroxyl, carboxyl and amino groups (Li et al., 2008; Chua et al., 2006). Moreover, a higher surface area of nanofibres facilitates the immobilisation of cell adhesive molecules onto the surface to a greater extent, thus resulting in better adhesive properties (Yoo et al., 2009). The total cell number was significantly higher when the cells were seeded onto nanofibrous scaffolds rather than onto polymer coated glass (In et al., 2008).

14.7

Choice of biomaterial

The success of a tissue engineered scaffold lies in the choice of the biomaterials. Among the different types of biomaterials, natural materials such as collagen, chitosan and gelatin are currently being investigated as potential scaffolds for skin regeneration. Collagen is the most abundant protein in many human tissues such as bone, skin, ligament, cartilage and tendon. Collagen, a natural ECM protein, is a polypeptide composed of repeating amino acids of glycine, proline and hydroxyproline. Because of its fundamental properties such as biodegradability, biocompatibility, mechanical strength and regulation of cell adhesion, spreading, proliferation and differentiation, it has been extensively investigated for biomedical applications. Moreover, the haemostatic property of collagen controls bleeding at the wound site and subsequently aids wound healing (Nimni and Harkness, 1988). Electrospun collagen scaffolds have been found to produce skin substitutes that closely resemble the cellular organisation, proliferation and maturation in vitro (Powell et al., 2008) of collagen. Additionally scaffold engraftment and a full thickness healing process were evaluated in electrospun scaffolds in athymic mice and found to have a higher engraftment rate and to reduce wound contraction better than freeze dried collagen scaffolds (Powell et al., 2008). In spite of its versatile properties, however, there are limitations to the use of collagen as a biomaterial due to immunological response, disease transmission and high cost. Chitin and chitosan are naturally occurring polysaccharides that have been widely investigated for skin regeneration (Khor and Lim, 2003). Unlike other fibres, a chitosan fibre carries a positive charge. This cationic property of chitosan exhibits electrostatic interactions with negatively charged molecules. The cationic nature of chitosan has an affinity towards the anions on bacterial cell walls, thereby preventing entry across the cell wall. This © Woodhead Publishing Limited, 2011


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microbicidal action of chitosan plays a major role in the wound healing process. Furthermore, chitin possesses wound healing, antifungal, biodegradable and thrombogenic properties, which makes it an attractive material for use in skin tissue engineering applications (Senel and McClure, 2004). Furthermore chitin nanofibrous scaffolds structurally resemble glycosaminoglycans, such as, chondroitin sulfates and hyaluronic acids in the ECM (Noh et al., 2006). Also, chitin nanofibres demonstrated greater cell adhesion and spreading for both human keratinocytes and fibroblasts compared to chitin microfibres (Min et al., 2004). Collagen–chitosan complex nanofibres have been found to resemble native ECM (Chen et al., 2007). Ultra-fine electrospun fibres of hexanoyl chitosan have been investigated as a skin tissue scaffold and were found to be non-toxic when evaluated with mouse fibroblasts. Moreover, this scaffold also supported the adhesion and proliferation of both human keratinocytes and human foreskin fibroblasts (Neamnark et al., 2008). Many researchers have investigated the biomedical applications of silk fibroin nanofibres. These have distinct biological properties such as biocompatibility, good oxygen and water vapour permeability, biodegradability and minimal inflammatory response (Min et al., 2004; Venugopal and Ramakrishna, 2005). Silk fibroin nanofibres were investigated with and without a coating of ECM proteins to study the influence of ECM proteins on both human keratinocyte and fibroblast adhesion and spreading (Min et al., 2004). Min et al. (2004) observed that a type I collagen coating promoted both cell adhesion and spreading, whereas a laminin coating only encouraged cell spreading. The fabrication of chitin–silk fibroin nanofibres showed improved keratinocyte adhesion compared to pure chitin matrices (Park et al., 2006b). In order to mimic the ECM topography, Schenke-Layland et al. (2009), generated the human ECM proteins in vitro, and fabricated them as electrospun nanofibres with polycaprolactone (PCL). This scaffold showed better mechanical properties and binding sites for cell adhesion and proliferation as native ECM with a similar topography (Schenke-Layland et al., 2009). Apart from the use of natural materials, synthetic materials such as poly lactide-co-glycolide could be used as a dermal substitute. Electrospun PLGA nanofibre matrices with seven different fibre diameters (150–225 nm; 200–300 nm; 250–467 nm; 500–900 nm; 600–1200 nm; 2500–3000 nm and 3250–6000 nm) were fabricated in order to evaluate the role of fibre dimension on proliferation of human skin fibroblast (Kumbar et al., 2008). Cells cultured on matrices with diameters ranging from 250–467 nm, 500–900 nm and 600–1200 nm showed better adherence, progressive growth and had a well spread morphology compared to the cells on matrices with diameters of 150–225 nm, 200–300 nm, 2500–3000 nm and 3250–6000 nm since the middle range fibre diameter is more like natural ECM (Kumbar et al., 2008). Electrospun photo-cross-linked dextran–PLGA nanofibrous scaffold was developed by blending PLGA and methacrylated dextran (Jiang et al.,



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2004). The photo-cross-linking of dextran in electrospun fibres improved the stability of the scaffold and also eliminated the use of cross-linking agents which could be toxic to the cells. This dextran/PLGA electrospun scaffold has been found to resemble dermal tissue structure and to encourage the adhesion and proliferation of dermal fibroblasts (Pan et al., 2006). Moreover, the dermal fibroblast was seeded with this scaffold and investigated for the presence of ECM proteins such as collagen and elastin after 5 days, which was confirmed by Sirius Red and Verhoeff’s solutions, respectively (Pan et al., 2006). The staining intensity of both proteins was increased from day 1 to day 5, thereby suggesting ECM protein deposition. Polycaprolactone/poly(trimethylene carbonate) (PCL–PTMC) electrospun nanofibres were successfully developed to release a herbal medicine ‘shikonin’, which possesses antitumour, antioxidant and antimicrobial properties, in a sustained manner for wound healing applications (Han et al., 2009). The antimicrobial property of shikonin-loaded PCl–PTMC fibres was investigated against two typical bacteria frequently found on burn wounds (E. coli and S. aureus); S. aureus was found to be more sensitive than E. coli. Similarly, the antioxidant property of this scaffold was determined using DPPH assay both before and after electrospinning and showed no significant difference in radical scavenging ability. A selection of electrospun biomaterial scaffolds for skin regeneration is listed in Table 14.1.

14.8

Cellular interactions on skin substitute

Cell–ECM interaction plays a major role in controlling cellular processes such as cell shape, adhesion, spreading and proliferation. Tissue engineered nanofibrous scaffolds resemble the ECM not only in terms of topography, but also by activating various intracellular signalling pathways which control cell fate (Toh et al., 2006). The diameter of the electrospun fibre is generally smaller than the cell, which allows the cell to form around it and attach to adsorbed proteins at various focal points (Pham et al., 2006). This cell–scaffold interaction enhances the deposition of ECM molecules, which in turn regulates cellular activities. Fibroblasts cultured on synthetic elastin electrospun nanofibres were found to deposit ECM molecules such as fibronectin and collagen after just 14 days in vitro (Li et al., 2005). Pan et al. (2006) confirmed the presence of ECM proteins, elastin and collagen, deposited by dermal fibroblasts that were cultured on dextran–PLGA electrospun scaffolds after 5 days (Pan et al., 2006). Furthermore, they also identified the expression of ECM proteins, cytokines and focal adhesion complex using real-time polymerase chain reactions (PCR) (Pan et al., 2006). Unlike native cell–ECM interactions occurring via integrins, cells adhere to the biomaterial scaffold via chemical bonds (e.g. covalent, hydrophobic and/ or hydrogen bonds). It has also been observed that cell–matrix interactions



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Table 14.1 Biomaterial nanofibrous scaffold used for skin regeneration Biomaterial Cells cultured nanofibres

Success of cell response

References

Plasma-treated Human epidermal silk fibroin keratinocyte (NHEK), fibroblast (NHEF)

Showed better cell adhesion and spreading than untreated

Jeong et al. (2009)

PCL–collagen

Adhered and proliferated

Venugopal et al. (2006)

Gelatin–poly Human primary dermal fibroblast (l–lactide–co– e–caprolactone) and keratinocyte

Higher cell adhesion and proliferation rate based on gelatin ratio in blend

In et al. (2008)

Gelatin–PCL BMSC, fibroblast

BMSCs spread well and migrated deeper inside gelatin–PCL scaffolds than PCL scaffolds; better fibroblast adhesion, growth and proliferation

Zhang et al. (2004); Chong et al. (2007)

Human dermal fibroblast

Poly (ethylene– co–vinyl alcohol)

Human aortic Support cell growth smooth muscle cell, human dermal fibroblast

PGA–chitin

Normal human Support cell adhesion Park et al. (2006a) epidermal fibroblast and spreading

Chitin

Normal human oral keratinocytes, normal human epidermal keratinocytes and normal human gingival fibroblasts

Chitin nanofibres showed better cell adhesion and spreading than chitin microfibres

Kenawy et al. (2003)

Noh et al. (2006)

promote ECM production promoting interaction with the focal adhesion complex on the cell membrane, subsequently influencing the expression of functional genes (Pan et al., 2006).

14.9


The main objective is to develop an ideal scaffold that offers complete restoration of functional skin, including regeneration of skin appendages, angiogenesis and scar-free interaction with the host’s tissue. Electrospinning has received a great deal of attention since the process is versatile and it is easy to develop scaffolds with tailorable properties. These properties can



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be tuned by either varying the fibre dimensions, their orientation and/or their composition. The biological properties of the synthetic scaffolds can be improved by changing the surface properties such as wettability, ECM protein coating and peptide sequence. In addition, electrospun fibres can be used as a carrier for the delivery of antibiotics, antimicrobial agents and gene/growth factors. Development of an ideal combination of biodegradable nanofibrous scaffold with multiple growth factors may be the future direction of research on electrospun skin substitutes. Tissue growth factors such as PDGF, TGF-b, EGF cover the wound and influence regeneration. Thus incorporation of these facts can improve the regenerative potency of the tissue. Currently, there has been a tremendous increase in research into the development of nanofibrous scaffolds for both cell delivery and growth factor delivery. However, development of a single biomaterial scaffold for skin tissue regeneration, which includes both of these approaches, still remains challenging. Furthermore, the growing interest in the search for a novel biomaterial that is able to promote rapid wound healing, encourage better cell interactions and angiogenesis in the complete absence of wound contraction and fibrosis (in either the short term or long term), should offer new insights into skin regeneration.

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O’Brien, F.J., Harley, B.A., Yannas, I.V. and Gibson, L. (2004). ‘Influence of freezing rate on pore structure in freeze-dried collagen-GAG scaffolds’. Biomaterials, 25(6), 1077–86. Pan, H., Jiang, H. and Chen, W. (2006). ‘Interaction of dermal fibroblasts with electrospun composite polymer scaffolds prepared from dextran and poly lactide-co-glycolide’. Biomaterials, 27(17), 3209–20. Parenteau, N.L., Sabolinski, M.L., Mulder, G. and Rovee, D.T. (1997). ‘Wound research’. In: Chronic Wound Care: A Clinical Source for Healthcare Professionals, 2nd edn, D. Krasner and D. Kane, (eds). Health Management Publications, Wayne, PA, 389–95. Park, K.E., Kang, H.K., Lee, S.J., Min, B.M. and Park, W.H. (2006a). ‘Biomimetic nanofibrous scaffolds: preparation and characterization of PGA/chitin blend nanofibers’. Biomacromolecules, 7(2), 635–43. Park, K.E., Jung, S.Y., Lee, S.J., Min, B.M. and Park, W.H. (2006b). ‘Biomimetic nanofibrous scaffolds: preparation and characterization of chitin/silk fibroin blend nanofibers’, Int J Biol Macromol, 38(3–5), 165–73. Pham, Q.P., Sharma, U. and Mikos, A.G. (2006). ‘Electrospinning of polymeric nanofibres for tissue engineering applications: a review’. tissue Eng, 12(5), 197–211. Powell, H.M., Supp, D.M. and Boyce, S.T. (2008). ‘Influence of electrospun collagen on wound contraction of engineered skin substitutes’. Biomaterials, 29(7), 834–43. Quinn, K.J., Courtney, J.M., Evans, J.H., Gaylor, J.D.S. and Reid, W.H. (1985). ‘Principles of burn dressings’. Biomaterials, 6(6), 369–77. Rho, K.S., Jeong,L., Lee, G., Seo, B.M., Park, Y.J., Hong, S.D., Roh, S., Cho, J.J., Park, W.H. and Min, B.M. (2006). ‘Electrospinning of collagen nanofibers: Effects on the behavior of normal human keratinocytes and early-stage wound healing’. Biomaterials, 27(8), 1452–61. Robinette, E.J. and Palmese, G.R. (2005). ‘Synthesis of polymer–polymer nanocomposites using radiation grafting techniques’. Nuclear Instr Meth Phys Res, 236, 216–22. Schenke-Layland, K., Rofail, F., Heydarkhan, S., Gluck, J.M., Ingle, N.P., Angelis, E., Choi, C.H., MacLellan, W.R., Beygui, R.E., Shemin, R.J. and Heydarkhan-Hagvall, S. (2009). ‘The use of three-dimensional nanostructures to instruct cells to produce extracellular matrix for regenerative medicine strategies’. Biomaterials, 30(27), 4665–75. Schindler, M., Ahmed, I., Kamal, J., Nur, E.K.A., Grafe, T.H., Young Chung, H. and Meiners, S. (2005). ‘A synthetic nanofibrillar matrix promotes in vivo-like organization and morphogenesis for cells in culture’. Biomaterials, 26(28), 5624–31. Senel, S. and McClure, S.J. (2004). ‘Potential applications of chitosan in veterinary medicine’. Adv Drug Delivery Rev, 56(10), 1467–80. Soh, C.L., Lim, J.M.C., Boyd, R.L. and Chidgey, A.P. (2009). ‘Epithelial stem cells and the development of the thymus, parathyroid and skin’. In: Regulatory Networks in Stem Cells, V.K. Rajasekhar and M.C. Vemuri (eds). Humana Press, New York, 405–37. Subramanian, A., Krishnan, U.M. and Sethuraman, S. (2009). ‘Development of biomaterial scaffold for nerve tissue engineering: Biomaterial mediated neural regeneration’. J Biomed Sci, 16, 108. Taepaiboon, P., Rungsardthong, U. and Supaphol, P. (2007). ‘Vitamin-loaded electrospun cellulose acetate nanofiber mats as transdermal and dermal therapeutic agents of vitamin A acid and vitamin E’. Europ J Pharmaceut Biopharmaceut, 67(2), 387–97. Toh, Y.C., Ng, S., Khong. Y., Zhang, X., Zhu, Y., Lin, P.C., Te, C.M., Sun, W. and Yu, H. (2006). ‘Cellular responses to a nanofibrous environment’. Nanotoday, 1(3), 34–43. © Woodhead Publishing Limited, 2011


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Venugopal, J. and Ramakrishna, S. (2005). ‘Biocompatible nanofibre matrices for the engineering of a dermal substitute for skin regeneration’. Tissue Eng, 11(5–6), 847–54. Venugopal, J.R., Zhang, Y. and Ramakrishna, S. (2006). ‘Invitro culture of human dermal fibroblasts on electrospun polycaprolactone collagen nanofibrous membrane’. Artificial Organs, 30(6), 440–6. Vepari, C. and Kaplan, D.L. (2007). ‘Silk as a biomaterial’. Prog Polym Sci, 32(8–9), 991–1007. Woo, K.M., Chen, V.J. and Ma, P.X. (2003). ‘Nanofibrous scaffolding architecture selectively enhances protein adsorption contributing to cell attachment’. J Biomed Mater Res, 67A(2), 531–7. Wu, L. and Ding, J. (2004). ‘In vitro degradation of three-dimensional porous poly(d,llactide-co-glycolide) scaffolds for tissue engineering’. Biomaterials, 25(27), 5821–30. Xie, J. and Wang, C.H. (2006). ‘Electrospun micro- and nanofibers for sustained delivery of paclitaxel to treat C6 glioma in vitro’. Pharm Res, 23(8), 1817–26. Xiumei, M., Zonggang, C. and Weber, H.J. (2007). ‘Electrospun nanofibres of collagenchitosan and P(LLA-CL) for tissue engineering’. Frontiers Mater Sci China, 1(1), 20–3. Xu, X. and Zhou, M. (2008). ‘Antimicrobial gelatin nanofibres containing silver nanoparticles’. Fibres and Polymers, 9(6), 685–90. Yoo, H.S., Kim, T.G. and Park, T.G. (2009). ‘Surface-functionalized electrospun nanofibers for tissue engineering and drug delivery’. Adv Drug Delivery Rev, 61(12), 1033–42. Zhang, Y., Ouyang, H., Lim, C.T., Ramakrishna, S. and Huang, Z.M. (2004). ‘Electrospinning of gelatin fibers and gelatin/PCL composite fibrous scaffolds’. J Biomed Mater Res, Part B: Appl Biomater, 72B(1), 156–65.



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15

Wound dressings

T. R. H a y e s and B. S u, University of Bristol, UK

Abstract: In this chapter, the use of electrospinning for the fabrication of materials for wound healing applications will be discussed. There is a definite overlap between wound healing and skin tissue regeneration, as discussed in Chapter 14. However, this chapter will concentrate on materials that can provide the optimum environment for natural healing to take place, with more emphasis on antimicrobial properties than the interaction between the cells and the materials. Key words: antimicrobial, electrospinning, nanofibres, wound dressings.

15.1

Introduction: wound healing

A wound is a disruption of tissue integrity, which causes severance of capillaries and destroys and damages cells. All wounds initiate bodily reactions with the single purpose of restoring the tissue integrity and therefore the original tissue function. Wound healing is often classified according to severity. Primary wound healing occurs in situations when there is a smooth, closely abutting incision wound without significant loss of tissue. In these instances there is a good prognosis for successful wound healing. Secondary wound healing occurs in deeper wounds where tissue loss occurs or where infection prevents the re-association of the wound edges. In these cases, the wound is closed by the growth of new granulation tissue, which fills the defect and which is eventually transformed into scar tissue.

15.1.1 Healthy wounds After initial haemostasis (arresting of blood flow) and blood clotting, all wound healing, irrespective of the wound type or severity, proceeds in three interrelated phases with overlapping time spans (Harding et al., 2002). The first stage in wound healing is the inflammatory, exudative or cleansing phase. This involves the exudation of blood plasma to the wound area, which delivers a number of inflammatory cells that are involved in wound cleansing and the defence against infections. These cells also secrete substances that create the necessary conditions for the subsequent phases of wound healing. The second stage of wound healing, known as the proliferative phase, is dominated by cellular proliferation with the aim of fast vascular regrowth 317 © Woodhead Publishing Limited, 2011


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and closure of the defect by generation of new tissue. The first, temporary tissue to be formed is called granulation tissue and comprises a network of type III collagen that acts as a ‘quick-fix’ to close the defect and block infection. This primitive tissue also acts as a moist bed for the subsequent regrowth of the epithelial cells. After having fulfilled its function, this granulation tissue is subjected to regression and is gradually replaced by stronger, long-strand type I collagen in the form of scar tissue (Stadelmann et al., 1998). The final stage of wound healing, the regeneration or epithelialisation phase, involves the regrowth of epidermal cells by cell division and migration from the wound edges across the smooth, moist surface of the granulation tissue (Stadelmann et al., 1998).

15.1.2 Non-healing wounds If a wound shows no healing tendencies under appropriate therapy within 12 weeks, it is classed as a non-healing or chronic wound (Harding et al., 2002). Chronic wounds can develop at any time from an acute wound if, for example, persistent infections are not recognised and treated effectively. However, they most commonly arise as the last stage of a progressive condition of tissue breakdown from vascular disease, pressure damage, radiation injury or tumours (Harding et al., 2002; Grey et al., 2006; Boateng et al., 2008). If the normal blood flow to an area of tissue is impaired, irrespective of the cause, the balance of nutrients and supply of oxygen are seriously affected and the tissue dies. This is the worst situation for the initiation of wound healing. The lack of oxygen and nutrients in the wound area prevents the successful cleansing of the wound and, therefore, the subsequent granulation and epithelialisation phases from occurring. Also toxic decay of the necrotic tissue and bacteria in the wound infiltrate the surrounding area causing further tissue degradation. In this situation, therapeutic intervention is required before the wound can heal (Boateng et al., 2008). Many of the common causes of non-healing wounds are most frequently found in elderly patients with, for example, vascular diseases, which account for 90% of chronic wounds. Decubitus ulcers (pressure sores), which result when areas of tissue are subjected to pressures exceeding that of the blood pressure across those regions, for a prolonged period of time, are also most common in elderly patients who are immobile and cannot change position whilst sitting or lying. In an ageing society, the effective treatment of these wounds becomes ever more important. Wound healing can only be initiated and completed when both the blood supply to the affected area is restored and the non-viable tissue in the wound is removed. This often means that the cause of the restriction in blood flow must be resolved before the wound can heal. After the cause has been treated, the next treatment is to clean the wound of necrotic tissue. Often,



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the best method for this is surgical debridement (removal of the non-viable tissue by mechanical means). This leaves a clean wound surface with an improved vascular situation, encouraging the restart of wound healing. At this point it is important to encourage the synthesis of the granulation tissue by maintaining a permanently moist condition in the wound. This can be achieved by using an occlusive or absorbent wound dressing (Jones et al., 2006).

15.1.3 Wound dressings The role of a wound dressing is to provide the optimum conditions for wound healing, whilst protecting the wound from further trauma and invasion by pathogenic microorganisms. It is also important that the dressings can be removed atraumatically, so as to prevent further damage to the wound surface during dressing changes. For the most serious types of wounds, such as burns and chronic ulcerations, which produce large quantities of exudates, it is generally accepted that moist wound therapy plays an important role in effective treatment (Queen et al., 2004; Jones et al., 2006). It is necessary to keep the wound bed moist to prevent cell death by desiccation, to provide a favourable environment for fibroblast proliferation and to allow re-epithelialisation by the migration of keratinocytes from the wound edges, across the surface of the newly formed granulation tissue. It is important however, to maintain a balance in moisture levels, as excess wound fluid can contribute to an increase in bacterial colonisation of the wound and can adversely affect the surrounding area by maceration of the healthy tissue (Jones et al., 2006). It is extremely important, therefore, to select an appropriate dressing material for the particular wound, in order to maintain the optimum moisture levels for wound healing to take place. Currently, there are a number of types of wound dressing used in moist wound therapies. These include hydrocolloids, which consist of a dispersion of absorbent particles within a self-adhesive elastomer, foams, which are usually made from coated polyurethane, hydrogels, which are prehydrated sheets of hydrophilic polymer, and hydrofibres, such as calcium alginates and carboxymethylated cellulose (Boateng et al., 2008; Jones et al., 2006; Queen et al., 2004; Harding et al., 2000). Each dressing variety has unique properties, which makes it suitable for different wound conditions; a disadvantageous property for the treatment of one wound may be beneficial for the treatment of another. Table 15.1 summarises the key properties of these categories of moist wound therapy dressings. Despite the number of alternative dressing types, there are currently few commercially available nanofibrous wound dressings. However, over the last decade there has been a rapidly increasing interest in nanofibre research,



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Table 15.1 Summary of wound dressings currently in widespread use Dressing type

Fluid Prevents Beneficial functions management maceration

Disadvantages/ limitations

Hydrocolloids

Moderate absorption

No

• Creates a barrier against infection • Convenient to apply

Foams

Moderatehigh absorption Moderatehigh absorption

Yes

• Provides thermal insulation and padding • Transforms to gel as exudate is absorbed • Easily removed • Rehydrates dry wounds • Relieves pain • Easily removed

• Wound edges can be damaged on removal • Prevents gaseous exchange • Adheres to wound if exudation is too low • Only effective on highly exuding wounds

Hydrofibres

Hydrogels Donates moisture

Yes

No

• Requires secondary dressing • Can become saturated

Source: Boateng et al., 2008; Jones et al., 2006; Queen et al., 2004; Harding et al., 2000

which has lead to several potential candidates for wound dressing applications (Min et al., 2004; Rho et al., 2006; Chen et al., 2008; Zhou et al., 2008).

15.1.4 Antimicrobials There are several methods for preventing or treating wound infections, the most common being the use of antibiotics and antiseptics. In recent years there has been an increased prevalence of antibiotic-resistant bacterial strains, such as MRSA (methicillin-resistant Staphylococcus aureus) and VRSA (vancomycin-resistant Staphylococcus aureus) amongst others, which have lead to an increased interest in alternative antimicrobial agents. One such antimicrobial is silver, which has been used for several centuries as a bactericide, fungicide and algaecide for applications ranging from food and drink storage to wound treatment. We will focus on silver in this chapter because of its wide application in wound management. Ionic silver The broad spectrum bactericidal efficacy of silver ions on microorganisms has long been known and utilised in the treatment of infected wounds (Klasen, 2000a; Silver et al., 2006). In modern medicine silver is most commonly used in the form of a compound, such as silver nitrate (AgNO3) or silver



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sulfadiazine (C10H9AgN4O2S), which when dissolved releases silver ions (Ag+). Despite the fact that the bactericidal effect of Ag+ is well known, the mechanism of its bactericidal action remains a subject for investigation. It has been shown that the silver interacts with thiol groups (—SH) of important proteins, inactivating them and leading to structural and metabolic disruption and subsequent cell death (Liau et al., 1997; Matsumura et al., 2003), but also that it reacts with bacterial DNA, preventing its replication (Fox and Modak, 1974; Modak and Fox, 1973). Silver compounds can be applied to wounds as a cream, as in the case of 1% silver sulfadiazine cream, which is used on third degree burns, or incorporated into wound dressings, such as Aquacel Ag (Convatec), Actisorb Silver (Johnson & Johnson), Silvasorb (Acrymed) and Tegaderm Ag (3M). Silver nitrate was the first silver compound to be used in modern medicine to prevent wound infection. Cotton gauze pads were soaked in silver nitrate solution and placed over the wound area (Klasen, 2000b). The problem with using some compounds, especially silver nitrate, is that the silver ions dissociate from the nitrate ions very quickly in solution, resulting in a very fast release of silver. The chloride ions in the wound fluid react with the silver ions from the silver nitrate solution to form silver chloride, which is a more stable compound and is not antimicrobial (Dunn and Edwards-Jones, 2004). This means that in order to maintain antimicrobial properties, silver nitrate must be renewed approximately every two hours, requiring dressing changes and potential damage to the wound bed. Another problem with the use of silver nitrate on wounds is that nitrates are toxic to cells and can impair re-epithelialisation of the wound surface. This side effect can out weigh the benefits of the silver ions. An improvement came with the use of silver sulfadiazine, which is in common use today. Silver sulfadiazine does not dissociate as quickly as silver nitrate, meaning that there is a more gradual release of silver ions into the wound environment (Fox and Modak, 1974; Modak and Fox, 1973). Although chloride ions still react with the released silver ions, there is a continual release of silver ions over a longer time period, allowing less frequent dressing changes, compared with the use of silver nitrate. Having said this, silver sulfadiazine treatments do still commonly need to be repeated once per day, which can be disruptive to the healing process, cause pain for the patient and require a considerable amount of carers’ time. An additional problem associated with the use of ionic silver in wound dressings is that many silver compounds are light sensitive. When exposed to ultraviolet (UV) radiation and in the presence of organic matter, such as skin or wound fluid, Ag+ ions are reduced to metallic silver (Ag0), which appears black. If a silver compound, such as silver nitrate, is absorbed into the skin from an ionic silver wound dressing, this reaction is likely to cause



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black staining of the skin. Although this is not generally harmful, it can be disconcerting for the patient and is certainly an undesirable side effect. For these reasons there has been a drive within research to develop alternative antimicrobial delivery mechanisms that can be incorporated into wound dressing materials, to prolong the release of antimicrobials whilst maintaining optimum conditions in which the wound can heal. Nanocrystalline silver In order to overcome the rapid deactivation of silver nitrate and silver sulfadiazine, an alternative type of silver-containing dressing was developed. Currently marketed as Acticoat 7 (Smith & Nephew), this consists of physical vapour deposition (PVD) sputtered coatings of nanocrystalline silver compounds on polyethylene meshes, which can be incorporated into the dressing. This coating comprises Ag2O and Ag2CO3 with crystal grain sizes of less than 20 nm. This small grain structure means that nanocrystalline silver behaves differently from ionic or bulk metallic silver, by slowly dissolving to release Ag+ and Ag0 (Dunn and Edwards-Jones, 2004; Atiyeh et al., 2007). There have been problems associated with the use of such methods, owing to the fact that all of the silver is at the fibre surface. This coating of silver can be susceptible to detachment from the fibre and once the silver falls off, the dressing loses its antimicrobial activity. In addition to this, as all of the silver is contained within the coating, there may be an excessively high concentration of silver in direct contact with the wound, which could potentially result in side effects, such as skin discolouration or cytotoxicity (Burd et al., 2007). Silver nanoparticles In addition to ionic and nanocrystalline silver, metallic silver nanoparticles have also been shown to possess antimicrobial properties. The mechanism of this action has been investigated using transmission electron microscopy (TEM) based techniques, which found that silver nanoparticles smaller than 10 nm can attach to and penetrate the cell membranes of Escherichia coli, a common gram-negative bacterium (Morones et al., 2005; Sondi and Salopek-Sondi, 2004). This observation was found to correlate with bacterial cell death, which was suggested to be as a result of nanoparticles interacting with the sulfur-containing proteins within the bacteria, causing membrane permeability and damage to their DNA (Morones et al., 2005; Sondi and Salopek-Sondi, 2004). It has also been reported that the bactericidal properties of silver nanoparticles are actually attributed to Ag+ ions that are adsorbed on to the



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surface of the particles owing to their high oxidation reactivity (Morones et al., 2005; Lok et al., 2007). One report has determined that free radicals are released from the surface of silver nanoparticles and, after comparing the antimicrobial effect of silver nanoparticles with samples also containing an antioxidant to eliminate free radicals, concluded that they may be responsible for the antimicrobial activity (Kim et al., 2007). If the antimicrobial properties are as a result of Ag+ ions or free radicals on the surface of silver nanoparticles, it would suggest that each particle acts as a reservoir of antimicrobial elements that can be released gradually. It has been reported that although the antimicrobial action of silver nanoparticles may share many of the same mechanisms as ionic silver, silver nanoparticles have a much higher antimicrobial efficacy. The concentration of silver nanoparticles required for antimicrobial efficacy was found to be at nanomolar levels, in contrast to the micromolar levels required for silver ions (Lok et al., 2006). The conclusion of this study is slightly misleading because concentrations of individual silver ions were compared with those of whole silver nanoparticles, which contain many silver atoms. Despite this, the study does show that nanoparticles are effective at extremely low levels and supports the theory that each particle can release many antimicrobial ions or radicals. An additional advantage of metallic silver nanoparticles over ionic silver is that they are not deactivated by compounds in wound fluid, such as halides, as is the case for Ag+ ions. Even if the antimicrobial action of nanoparticles is by ionisation of the surface and release of Ag+, this means that the activity of the nanoparticles will remain for significantly longer than that of a substance that quickly dissociates and releases all of its silver as Ag+.

15.2

Nanofibres

Nanoscale materials are technically defined as materials with dimensions less than one hundred nanometres (Roco, 2003, 2007), therefore nanofibres are fibres with diameters in this range. The term is, however, commonly given to any fibres with diameters less than one micrometre. Some authors prefer to use the term ultra-fine fibres, a more vague, non-specific phrase. In this chapter, the term nanofibre will be used to refer to any fibre with an average diameter in the sub-micrometre range. Fibres with diameters greater than one micrometre will be referred to as microfibres. Nanofibrous materials possess unique qualities and have shown tremendous potential for use in both wound dressing and tissue engineering applications and are therefore attracting significant interest from the research community (Zhang et al., 2005).



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15.2.1 Advantages of nanofibres in wound dressing applications The unique properties of nanofibres, which can be considered to be advantageous over conventionally produced textile fibres, are their significantly enlarged specific surface areas, superior mechanical properties, such as stiffness and tensile strength, and their flexibility in surface chemistry as a result of the increased proportion of the polymer molecules that are exposed to the fibre surface (Ko, 2004; Deitzel et al., 2001). The proportion of molecules, and hence functional groups, that are exposed to the surface of a fibre (Ms), is estimated by using equation [15.1]. M s = 100

pd D

[15.1]

where d is the diameter of a polymer chain and D is the fibre diameter. This suggests that a conventionally spun textile fibre, with a diameter of 10 mm, would have approximately one hundred times fewer exposed functional groups than a 100 nm diameter fibre of the same material. Figure 15.1 shows the dependence of specific surface area on fibre diameter. It is also of interest to note that nanofibres are in the same size range as, or even smaller than, many biological structures. Because of the reduced fibre diameter and increased specific surface area, nanofibres possess several important advantages over conventional wound dressing materials. 10000 Natural extracellular matrix (ECM)

Specific surface area (m2 g–1)

1000 100

Conventional textile fibres

Diameter of DNA helix

10

Most human cells

Typical electrospun nanofibres

1 0.1 0.01 0.001

0.01

0.1 1 Fibre diameter (µm)

10

100

15.1 Graph showing the effect of fibre diameter on the specific surface area of fibrous materials.



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They include (Zhang et al., 2005): ∑

Enhanced haemostasis: Nanofibrous wound dressings with their small pore size and large surface area can promote the haemostatic process. ∑ Good absorbability: Owing to their high specific surface area, nanofibrous wound dressings have a higher water absorption capacity. This would help provide and maintain the suitable moisture environment that is necessary for wound healing. The high absorbability is also beneficial to the absorption of exudates. ∑ Semi-permeability: The nanofibre structured dressing is porous with pore size 5 ¥ 106 cells ml–1, in order to support colony formation and maintain cell viability and growth. Gao et al. (2010) assessed the attachment of human ES cells to the synthetic, degradable polymer poly(lactic-co-glycolic acid) (PLGA), both in 2D and 3D scaffolds formed by phase separation. In the study PLGA films were coated with collagen I, collagen IV, laminin and fibronectin, with laminin providing the best adhesive substrate for human ES cells, a result that also translated to 3D scaffolds. 3D PLGA sponges with a 250–500 mm pore size have additionally supported the differentiation of human ES cells into neurons, cartilage and liver cell phenotypes (Levenberg et al., 2003). Cells induced towards a neural lineage formed ductular structures resembling the embryonic neural tube. Immature capillary networks formed through the structures (with the exception of RA treated samples) and 2-week-old constructs were permeated by host blood vessels when transplanted into immunodeficient mice. These studies highlight the potential of biomaterial scaffolds to differentiate ES cells into organised, multilayered structures. However, Smith et al. (2009) demonstrated the advantage of a nanofibrous substrate in the differentiation of human ES cells. Phase separation of poly(lactic acid) (PLA) scaffolds under different parameters produces both



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nanofibrous and solid walled scaffolds. Enhanced osteogenic differentiation of mouse ES cells was observed on the nanofibrous architecture compared to smooth walled scaffolds, with differential requirements for supplements and growth factors observed between the two surfaces.

18.5

Combining ES cells with electrospun scaffolds

18.5.1 Maintenance of pluripotency As discussed above, the architecture of electrospun meshes can itself alter ES cell signalling and enhance expression of pluripotent markers such as Nanog. Nur et al. (2006) cultured mouse ES cells on gelatinised polyamide meshes. Cells formed colonies on the cell surface of electrospun scaffolds but failed to attach to flat films, highlighting the effects of surface topography on cell attachment and proliferation. The effect of the topography on ES cell behaviour was also demonstrated by enhanced proliferation compared to flat gelatinised coverslips. In support of previous research by the same group (Nur et al., 2005), the study linked the increased proliferation to significantly increased Rac GTPase activity. The study also observed enhanced PI3K activity on the electrospun scaffolds which was associated with an increase in Nanog expression compared with flat 2D gelatinised glass coverslips. Although cells were cultured with MEFs and long-term culture was not assessed, the study highlights the potential of electrospun meshes to govern or enhance the culture of mouse ES cells. Our laboratory has focused on developing PLGA scaffolds for mouse ES cell culture and has cultured cells on the surface of sub-micrometre fibre meshes with an average fibre diameter of 570 nm ±380 nm (Fig. 18.5). After 24 hours, mouse ES cells attach to the cell surface and go on to proliferate over a 7 day culture period, forming heaped masses covered with a dense layer of deposited ECM (Fig. 18.5). Despite the promise electrospun meshes have demonstrated in the culture of other stem cells, such as mesenchymal stem cells (MSCs) (Kolambkar et al., 2010; Xin et al., 2007), few studies have combined ES cells with electrospun scaffolds. However, several meshes developed and tested using different cell types are promising constructs for mouse and human ES cell culture. As described above, laminin and fibronectin have both proved effective substrates for human ES cell attachment. Prospective meshes that may support human ES cell culture include electrospun laminin scaffolds composed of fibres ranging between 113 nm and 223 nm, depending on spinning parameters (Neal et al., 2009). Despite both the electrospinning process and the solvent potentially denaturing the protein, electrospun laminin retained its biological activity, enhancing the attachment of adipose stem cells and supporting their neural differentiation (Neal et al., 2009). Although meshes were not © Woodhead Publishing Limited, 2011


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Day 1

386

2 µm

6.67 µm

2 µm

Day 7

Day 3

3.33 µm

50 µm

14.3 µm

18.5 Culture of mouse ES cells on electrospun PLGA meshes. Mouse ES cells were seeded at 5 ¥ 104 cells cm–2 onto the surface of PLGA sub-micrometre fibre scaffolds and cultured for 7 days. Samples were analysed using scanning electron microscopy (SEM). Cells attached and proliferated on the scaffold surface, progressively depositing ECM over the 7 day period and forming rounded aggregates.

chemically cross-linked, assessment after 24 hours in culture media revealed that the fibres remained intact (Neal et al., 2009). The fibrous morphology of the laminin network is reminiscent of its structure within the assembled basement membrane; therefore the presentation of the ECM molecule in this orientation may have advantages over laminin adsorbed on TCP for the culture of human ES cells. Presentation of bioactive ligands from ECM molecules such as laminin and fibronectin on fibrous scaffolds could also have similar advantages. Kim and Park (2006) successfully functionalised electrospun PLGA with arginine-glycine-aspartic acid (RGD) peptides (an a5b1 engaging peptide motif derived from fibronectin). The ligand was orientated to the surface of the fibres and enhanced fibroblast attachment and proliferation. Incorporation of poly(ethyleneimine) in polyurethane electrospun scaffolds enabled functionalisation with GRGDSPK peptide and recombinant Fibrillin-1 fragment PF9, leading to enhanced spreading of murine fibroblasts (Jozwiak et al., 2008).



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Electrospinning has also been utilised to create growth factor loaded meshes for tissue engineering. Sahoo et al. (2010) compared polymer blending (electrospinning of a blended polymer/growth factor solution) and coaxial spinning to incorporate FGF-2 into PLGA fibres. Two separate solutions of FGF-2 and PLGA were prepared for the coaxial method. The aqueous FGF-2 solution was loaded into a separate inner syringe which was placed within a larger syringe containing the PLGA spinning solution. Simultaneous electrospinning of the outer and inner solutions combined them in a core–shell fibre structure, with the inner FGF-2 encapsulated within an outer PLGA shell. By keeping the two solutions separate, the growth factor was protected from the harsh solvents used to dissolve the polymer. This is in contrast to the blended solution which produced a random distribution of FGF-2. The incorporation of FGF-2 led to enhanced attachment and proliferation of rabbit MSCs. However, despite different distributions of FGF-2 within the fibres, little difference was observed between the two conditions. Covalent immobilisation of growth factors to scaffolds, including electrospun meshes, has demonstrated advantages over release in a soluble form. Nur et al. (2008) covalently bound FGF-2 to amine-modified polyamide electrospun meshes. The covalent attachment of FGF-2 dramatically increased its stability in culture, with approximately 20% remaining after 48 hours at 37°C. The tethered FGF-2 signalled through FGF receptors (FGFRs) on the cell surface, enhancing spreading and proliferation of fibroblasts on the scaffold surface. The modified scaffolds were also applied to human ES cells, with the tethered FGF-2 supporting enhanced attachment and colony formation. Unfortunately, the study did not assess the level and persistency of FGF-2 signalling compared to soluble FGF-2. However, it could be hypothesised that the increase in stability, inhibition of endocytosis and increased local concentration of FGF-2 led to increased signalling and enhanced cellular responses, as reported in previous studies (Fan et al., 2007).

18.5.2 Differentiation ES cells have been differentiated into a number of cell phenotypes on electrospun meshes, including adipocytes (Kang et al., 2007), cardiomyocytes (Fromstein et al., 2008) and neurons (Carlberg et al., 2009; Xie et al., 2009). Kang et al. (2007) reported enhanced adipogenic differentiation of mouse ES cells on polycaprolactone (PCL) meshes with the formation of spherical lipid laden cells surrounded by flattened non-adipogenic cells, reminiscent of adipocyte cell architecture in vivo. The ability of the fibre architecture to enhance differentiation was further supported by adipocytic differentiation without hormonal induction. Fromstein et al. (2008) compared polyurethane thermally induced phase separated (TIPS) scaffolds with electrospun meshes for the differentiation of cardiomyocytes from mouse ES cells. Polyurethane



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meshes comprised fibres ranging between 2 mm and 10 mm and created a scaffold 70 mm thick. In comparison, TIPS scaffolds were 1000 mm thick and were composed of interconnected pores with a rough textured surface. Cells adopted different morphologies on the two scaffolds. Cells on the TIPS scaffold adopted a rounded morphology, whilst cells on the electrospun meshes elongated along the fibres and adopted a spread morphology more characteristic of cardiomyocytes. Electrospun meshes demonstrated enhanced cell viability, as the thicker TIPS scaffolds restricted nutrient diffusion to the scaffold centre. However, despite the differences in morphology, both scaffolds supported cardiomyocyte contraction. Owing to the potential application of electrospun scaffolds in nerve regeneration, ES cells have been differentiated into a neural lineage on random polystyrene (Carlberg et al., 2009) and aligned PCL fibres (Xie et al., 2009). Human ES cells differentiated on random polystyrene meshes in N2B27 media supplemented with EGF and FGF-2 formed neurons that were positive for Map2 and bIII-tubulin, and the dopaminergic marker tyrosine hydroxylase (TH) (Carlberg et al., 2009). In comparison, 2D cultures consisted of a large population of glial fibrillary acidic protein (GFAP) positive astrocytes with a lower proportion of Map2, bII tubulin and TH positive neurons. Xie et al. (2009) differentiated mouse EBs on both aligned and random fibres. On both orientations, EBs remained in a rounded morphology on the scaffold surface as cells differentiated into neurons, oligodendrites and astrocytes. However, during the differentiation process, the differentiating cells migrated away from the main EB body, a behaviour which was enhanced on aligned meshes. Both migration and neurite extension were guided by the direction of fibre alignment, as the aligned fibres served as an instructive surface that directed and enhanced neurite elongation. The mechanical properties of electrospun meshes could also be tailored to facilitate differentiation into particular cell lineages. For instance, the matrix stiffness of polyacrylamide gels induced the differentiation of human MSCs, in the absence of supplementary growth factors, into cell phenotypes which reside in tissues with similar mechanical properties (Engler et al., 2006). In addition, mesodermal and osteogenic differentiation of mouse ES cells was promoted by substrates with increasing stiffness (Evans et al., 2009).

18.5.3 Creating a three-dimensional culture system One of the drawbacks of electrospun matrices is the lack of cell infiltration, creating a topographically distinct but essentially 2D culture environment. This is evident in Fig. 18.5, where the ES cells reside solely on the surface of the scaffolds rather than infiltrating into the mesh. A scaffold with accessible cell volume could potentially allow large scale, defined culture of ES cells, a necessary requirement if ES cells are to be used therapeutically. In addition, a



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truly 3D environment will facilitate the formation of multilayered, organised tissue constructs as reported for PLGA sponges (Levenberg et al., 2003). During standard electrospinning, fibres are deposited on top of one another creating a dense meshwork that does not encourage cell infiltration. A number of studies have attempted to overcome this using a variety of techniques. Nam et al. (2007) incorporated salt crystals into PCL fibre meshes using a sheath surrounding the electrospinning capillary. The scaffolds had a delaminated layer structure with 100–200 mm gaps between layers created by the salt deposits, with cell infiltration observed from the edge inwards. However, the scaffold was not uniform in structure and there was little evidence of cell infiltration between the delaminated areas. More promising scaffolds were produced by a novel ‘hydrospinning’ method (Tzezana et al., 2008). Sheets of PCL fibres were electrospun into a water bath and layered on top of one other before being placed in a vacuum oven. The vacuum oven treatment led to the evacuation of trapped water, which stretched and expanded the mesh creating a looser 3D scaffold, enabling the infiltration of myoblasts and EBs formed from human ES cells. Another approach is to co-electrospin water-soluble fibres such as poly(ethylene oxide) (PEO) or gelatin, the leaching of which would leave behind a looser network of synthetic fibres (Ekaputra et al., 2008). The infiltration of human foetal osteoblast derived cells was assessed in comparison to meshes composed of fibres in the micrometre range (1.61 mm). Larger fibres were also electrospun in conjunction with electrosprayed heprasil (a commercially available hydrogel of heparin and HA), under the hypothesis that the incorporation of heprasil would reduce the volume density of the fibres and provide a biodegradable element available for cellular remodelling. Cells demonstrated a more spread phenotype on all meshes composed of sub-micrometre fibres with limited infiltration observed, the incorporation of PEO and gelatin offering little or no advantage. On meshes composed of larger fibre diameters without electrosprayed heprasil, cells adopted a more rounded morphology, migrating 50 mm into scaffolds. However, the incorporation of heprasil led to full penetration of cells into the scaffolds. Enhanced migration into meshes composed of larger fibres was also reported by Pham et al. (2006). By increasing the flow rate and concentration of the PCL spinning solution, the study produced fibres 5 mm in diameter which in turn created an average pore size greater than 20 mm. This enabled infiltration of rat MSCs with full perfusion achieved in conjunction with a bioreactor. The increase in pore size with increasing fibre diameter is in agreement with the mathematical model proposed by Eichhorn and Sampson (2005) which concluded that the number of fibre contacts and distance between crossings is a function of fibre diameter; larger fibre diameters decrease the number and increase the distance between the fibre crossings, creating a larger pore size.



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18.6


Future trends

Considering the versatility and well established nature of electrospinning, it is surprising that so few studies have combined electrospun scaffolds with ES cells. A suitable, universal substrate for human ES cell culture has yet to be fully realised and fibrous electrospun meshes present an excellent opportunity to create a surface that replicates both the architecture and biological cues of the natural ECM. The requirement for both these attributes is partially supported by the superiority of matrigel, a preassembled ECM, in comparison to ECM molecules adsorbed onto a flat TCP surface. The ability to tailor fibre diameter and decorate the meshes with ECM molecules and peptide motifs is an attractive method of producing a fully defined, versatile substrate for ES cell culture. The modification of synthetic polymer-based electrospun scaffolds with bioactive ligands or the incorporation of biologically relevant molecules is a more feasible approach, as synthetic polymers scaffolds have high reproducibility and increased stability compared to meshes composed of purely natural polymers. The enhanced stabilisation of covalently immobilised growth factors also adds an extra dimension to mesh functionalisation. An avenue that has yet to be explored is the application of HS or heparin functionalised meshes to ES cell culture to facilitate and enhance growth factor binding. ES cells appear to manipulate their immediate microenvironment by expressing an unusually low sulphated variant of HS, which becomes decorated with specific sulphation motifs upon differentiation, differentially regulating the binding of growth factors. Such a response could be mimicked by HS on the electrospun mesh, facilitating growth factor signalling which functions either to enhance differentiation or maintain pluripotency. As discussed, electrospun meshes have also been applied to the differentiation of ES cells into several cell types. In each case, the fibrous architecture enhanced differentiation compared to 2D surfaces and in some cases fibre alignment had a direct impact on cell behaviour (Xie et al., 2009). However, in order to create multilayered organised tissue, cell ingrowth into the mesh is necessary. Progress has been made in creating 3D meshes for cell culture, by increasing fibre diameter and therefore pore size or by creating looser meshes. As well as enhancing differentiation, a 3D electrospun culture system may also provide increased volume for cell culture, a characteristic that is required if ES cells are to be propagated on a large scale.

18.7

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Index

abiotic degradation, 35, 45–8 hydrolytic degradation of ester, anhydride and carbonate linkages, 46 Achilles tendon, 151, 152, 158 ruptures treatment, 153 Acticoat 7, 322 Actisorb Silver, 321 advanced therapy medicinal product, 76–7, 85 Advanced Therapy Medicinal Products Regulation (EC) 1394/2007, 68, 75 aggrecan, 111 Alamar Blue, 364 albumin, 231 alkaline phosphatase, 95 Alloderm, 302 allografts, 94, 152–3 antimicrobials, 320–3 ionic silver, 320–2 nanocrystalline silver, 322 silver nanoparticles, 322–3 Apigraf, 302 Aquacel Ag, 321 ATMP see advanced therapy medicinal product augmentation cystoplasty, 227 autofluorescence, 351 autografts, 93–4 axons, 171, 172 Bakelite, 34 bands of Büngner, 183, 184 formation, 185 BDNF see brain-derived neurotrophic factor

bioactive scaffolds, 217 biomimetic effect, 334 biopolymers, 153–4, 282, 283 biotic degradation, 35, 48 bisphosphonates, 102 bladder tissue regeneration, 225–38 bladder disease and the need for bladder substitution, 227–8 electrospinning fit for purpose, 234–7 control over architecture, 234 porcine bladder tissue optical micrograph, 235 promoting cellular alignment, 234–6 surface modification and functionalisation, 236–7 electrospun and other scaffolds, 228–33 micrograph from porcine bladder matrix extract, 233 natural/biologically derived polymer scaffold, 232–3 synthetic polymer scaffolds, 230–1 urothelial cells micrograph, 229 in vitro bladder tissue development, 230 future trends, 237 structural/functional properties, 225–7 urinary bladder haematoxylin- and eosin-stained section, 226 urothelium, 226–7 BM-MSCs see bone marrow derivedMSCs Bombyx mori, 37 bone formation, 95–6 bone marrow derived-MSCs, 187 bone morphogenetic proteins 4 (BM 4), 376


Electrospinning-Bosworth-Index.indd 397

6/6/11 6:53:00 PM

398

Index

bone morphogenetic proteins (BMPs), 95, 101 bone remodelling, 93 bone resorption, 96–7 bone tissue regeneration, 93–106 bone biology principles, 94–7 differentiated and TRAP positive stained osteoclasts, 96 osteoblasts and bone formation, 95–6 osteoclasts and bone resorption, 96–7 future trends, 105–6 materials for scaffolds, 99–101 composite materials, 100–1 functionalisation of fibres, 101 polymeric materials, 99–100 osteoporosis, 101–2 scaffolds fabrication, 97–9 extracellular matrix structure, 99 strategies, 97, 98 properties required by bone tissue engineered scaffold, 98 terminology, 94 treatment strategies, 102–5 osteoblast cells cultured on PCL/PVPA scaffolds, 103 parameters used to produce PCL fibres, 102 PCL fibres functionalised with PVPA, 103 PVPA on bone formation, 103–5 PVPA on bone resorption, 105 PVPA structural formula, 102 water droplet on PCL and PCL surface functionalised with PVPA, 104 brain-derived neurotrophic factor, 179, 188 carbon nanotubes, 136–7, 139–40 PCL-0.05% MWCNTs–PAA/PVA, 140 PCL fibre containing a functionalised MWCNT, 139 rat muscle cell activity on PCL scaffolds, 141 cartilage tissue regeneration, 111–24 articular cartilage, 112 culture of chondrogenic cells for implantation, 113–19

human articular chondrocytes, 114–16 human stem cells, 116–19 electrospun nanofibre scaffolds, 119–24 choice of fibre diameter and control of construct porosity, 122–3 choice of materials for cartilage repair, 120–2 polymers tested, 120 scaffolds with graded properties, 123–4 future, 124 CE mark, 73–4, 77 cell culture systems kidney research, 343–57 cells immunostaining on electrospun scaffolds, 352–3 comparison of culture methods, 353–4 current work, 345–8 extracellular matrix proteins immunostaining on electrospun scaffolds, 351–2 future trends, 355–7 results, 348–50 scanning electron microscopy of cells on electrospun scaffolds, 350–1 pancreatic research, 359–70 future trends, 369–70 methods, 363–4 Min6 cell line, 361–2 Nes2y cells, 362 novel scaffolds and production methods, 362–3 results, 364–8 stem cell research, 372–90 combining ES cells with electrospun scaffolds, 385–9 current culture techniques, 375–84 3D scaffolds, 384–5 embryonic stem cells, 373–5 future trends, 390 CellCrown, 364 cellulose, 38 Centella asiatica, 334 central nervous system, 169–70 stem cell treatment for diseases, 170 chemotaxis processes, 179 chitin, 308–9



6/6/11 6:53:00 PM

Index chitosan, 308–9 ciliary neurotrophic factor (CNTF), 179 CNS see central nervous system collagen, 40–1, 149, 150, 154, 174, 208, 211, 308 type I, 94–5, 104 types, 150 collagen scaffolds, 130–1 collectors, 57 PCL fibres spun with HFIP onto a mandrel, 58 PCL fibres spun with HFIP onto a stationery plate, 58 Common Technical Document (CTD), 82 composite cystoplasty, 228, 237 composites, 20 conductive polymers, 133 cellular attachment for scaffolds, 134 rat muscle cell proliferation on 7% Au–PLLA scaffolds, 136 conductivity, 53–4 contractile force, 131–3 electrical stimulation, 132 mechanical stimulation, 132 corticosteroids, 151 cotton, 36 curcumin, 334 cylindromas, 244 decentralised procedure (DP), 71 degradable polymers, 35–6 degradation, 35–6 DegraPol, 255–8, 259 characteristics, 256 degradable block polyesterurethane for tissue engineering, 255 degradation in buffered aqueous solution, 256 electrospun tubular scaffold, 256–8 chloroform DegraPol solutions, 257 collector spindle, 257 electrospun scaffold section image with fibre spacings, 258 operational conditions and solution properties in electrospun reticulate fibres production, 259 scaffold surfaces obtained under different thermic conditions, 257 demineralised bone matrix (DBM), 94 dental regeneration, 280–92

399

dental restorations reinforcement, 284–92 future trends, 292 periodontal regeneration, 281–4 aligned electrospun polycaprolactone scaffolds micrographs, 283 electrospun scaffolds micrographs, 282 SEM images as-electrospun PAN–PMMA continuous nanofibres, 288 Bis-GMA/TEGDMA composite fracture surface, 289 and EDS spectrum of PU/nHA composite electrospun fibres, 291 electrospun nano-scaled glass fibres, 286 electrospun nylon 6 nanofibres, 287 neat nylon 6 nanofibres and nylon 6/fibrillar silicate nanocomposite nanofibre, 288 polyurethane electrospun fibres, 290 PU/nHA composite, 290 SiO2 electrospun nanofibres, 287 dermis, 300 dextran, 309–10 Diabetes mellitus, 359–60 dielectric constant, 54 Directive 2000/70/EC, 69 Directive 2001/83/EC, 79 Directive 2007/47/EC, 84 Directive 90/385/EEC, 85 drug delivery, 17–20 ECM see extracellular matrix elastin, 208, 211 electrospinning, 3–21, 326–7, 345–6 ambient parameters, 61–4 gas composition and movement, 64 humidity, 62 temperature, 62–3 annual number of publications, 4 apparatus, 10–11 applications, 15–21 composites and templates, 20 drug delivery and tissue engineering, 17–20 filters and textiles, 16–17



6/6/11 6:53:00 PM

400

Index

large-scale electrospinning setup, 17 others, 20–1 PLA/collagen electrospun fibres, 19 scaffold architecture, 19 basic concepts, 4–5 setup schematic, 4 conditions and control, 51–65 process variables divided into classifications, 51 future trends, 21 materials, 11–15 endothelial lined PCL/collagen scaffolds, 14 polymer composite fibre morphologies, 15 schematic of different setups, 12 morphology and structural formation, 5–7, 8 electrospun fibre morphologies, 6 tetrahydrofuran annealing process, 8 parameters, 7–10 PMMA-b-PDEA-b-PMMA fibres, 9–10 process parameters, 9–10 substance parameters, 7 processing parameters, 56–61 additional electromagnetic fields, 59–60 collectors, 57 electric field strength, 56–7 flow rate, 60 PLGA fibres spun with HFIP, 56 spinneret, 61 spinning distance, 57 regulation in regenerative medicine materials, 69–84 European regulatory frameworks, 69–75 regulatory input in development projects, 82–4 tissue regeneration products regulation, 75–82 regulatory issues, 67–89 (see also regulatory frameworks) future trends, 84–7 materials in regenerative medicine, 69–84 medical products, 67–8 tissue regeneration products, 68–9 solution parameters, 52–6

conductivity, 53–4 dielectric constant, 54 polymer choice, 56 solvent choice, 55–6 surface tension, 54 viscosity, 52–3 electrospun nanofibres, 285, 327–8 embryonic stem cells, 177, 187, 372, 373–5 derivation and definition, 373–5 mouse embryonic cells derivation, 374 encapsulation, 362, 368 energy dispersive spectroscopic (EDS) analysis, 291 enterocystoplasty, 227 environmental scanning electron microscopy, 265–6 scaffold culture in a bioreactor, 267 scaffold microstructure at the grooves bottom, 268 Epicel, 302 epidermal growth factor, 188 epidermis, 299–300 ESC see embryonic stem cells ESEM see environmental scanning electron microscopy essential requirements (ERs), 73 European Committee for Standardisation, 80–1 European Medicines Agency (EMA), 71 extracellular matrix, 148–9, 150–1, 173, 208 extracellular matrix ligands, 347 fibrinogen, 211, 233 fibronectin, 150, 154 filters, 16–17 functionally graded periodontal membrane (FGM), 282 gel permeation chromatography (GPC), 256 gelatin, 211 glass ceramics, 284 glass fibres, 284, 285 Glial cell line-derived neurotrophic factor (GDNF), 179 Glial cells, 175 glomerular podocyte culture, 346–7 glycolic acids, 211



6/6/11 6:53:00 PM

Index glycoproteins, 150 gold nanoparticles, 133, 135–6 elastic modulus and yield stress for PLLA scaffolds, 138 rat muscle cell proliferation on 7% Au, 137 rat primary skeletal growth, 134 granulation tissue, 318 guided bone regeneration (GBR), 281 guided tissue regeneration (GTR), 281 heart valve tissue regeneration, 202–18 future trends, 217–18 scaffold material selection, 211–12 scaffold properties to meet tissue requirements, 212–17 balancing scaffold degradation and tissue formation, 212–13 bioactive scaffolds, 217 3D electrospun heart valve scaffolds, 212 fibre alignment, diameter and morphology, 213–17 geometry, 212 PLA fibres electrospun pictures, 216 semilunar heart valves, 206–11 cell–matrix interactions, 208–11 fibrosa, spongiosa and ventricularis configuration within the aortic valve leaflet, 209 geometry and function, 206–7 haemodynamic parameters that apply on the aortic valve, 207 human aortic heart valve, 206 structure and composition, 207–8, 209 specific tissue requirements for scaffold properties, 205–11 tissue replacement, 203–5 heparan sulphate proteoglycans (HSPG), 378 hexafluoro-1,1,1,3,3,3-isopropanol (HFIP), 232 hexafluoroisopropanol, 56 HFIP see hexafluoroisopropanol high-density polyethylene (HDP), 249 homografts, 204 human articular chondrocytes, 114–16 cartilage from bone marrow stem cell differentiation, 116

401

loss of phenotype of chondrocytes in monolayer culture, 114 human bladder smooth muscle cells (HBSMC), 133 human stem cells, 116–19 chondrocytes attached to macrofibres and nanofibres, 118 hydrocolloids, 319 hydrofibres, 319 hydrogels, 319 hydrospinning method, 389 hypodermis, 300 Image Pro Plus, 235 immunofluorescence, 348 Integra, 302 International Conference on Harmonisation (ICH), 72, 82 islet transplantation, 360–1, 368 ISO 9001, 80, 81 ISO 13485, 82 ISO 14971, 82 ISO EN 13485, 81 kidney research, 343–57 cells immunostaining on electrospun scaffolds, 352–3 use of actin cytoskeleton and antinephrin antibody to immunostain cells on electrospun scaffolds, 353 comparison of culture methods, 353–4 podocytes cultured on glass coverslips vs cells cultured on electrospun PCL nanofibres, 354 current work, 345–8 commercially immortalised human podocytes, 346 electrospinning, 345–6 electrospinning parameters and solution conditions for PCL, PLGA, and polystyrene, 346 glomerular podocyte culture, 346–7 immunofluorescence, 348 scanning electron microscopy, 347–8 use of extracellular matrix ligands, 347 electrospun materials, 348–50



6/6/11 6:53:01 PM

402

Index

PS, PLGA and PCL scanning electron micrographs, 349 extracellular matrix proteins immunostaining on electrospun scaffolds, 351–2 extracellular matrix proteins immunofluorescence image, 352 future trends, 355–7 glomerular capillary lumen demonstrating three components of glomerular filtration barrier, 344 scanning electron microscopy of cells on electrospun scaffolds, 350–1 human podocytes cultured on materials coated with extracellular matrix proteins, 350 lactic acids, 211 leaflets, 208 leukaemia inhibitory factor (LIF), 376 macrophage colony stimulating factor (M-CSF), 96 marketing authorisation (MA), 71 Mark–Houwink–Sakurada equation, 53 matrix metalloproteinases (MMPs), 209 matrix remodelling, 209 Media Cybernetics, 235 medical devices, 78 regulatory framework, 72–5 Medical Devices Directive 93/42/EEC, 69, 78, 85, 87 medical products, 67–8 medicinal products, 71–2 Mefoxin, 17 mesenchymal stem cells, 177 3-methacryloxypropyltrimethoxy silane (MPTMS), 285 Min6 cell line, 359, 361–2 molecular self-assembly, 325–6 MSC see mesenchymal stem cells mulberry silkworm see Bombyx mori multiwalled carbon nanotubes, 137, 139 muscle fibres, 127 muscle tissue regeneration, 127–42 conductive elements, 133–40 carbon nanotubes, 136–7, 139–40 conductive polymers, 133, 134

elastic modulus and yield stress for PCL scaffolds, 142 gold nanoparticles, 133, 135–6, 138 contractile force, 131–3 electrical stimulation, 132 mechanical stimulation, 132 future trends, 140–2 skeletal muscle injuries, 128 soft tissue exploration/debridement and amputations, 128 skeletal muscle mechanical properties, 129–30 tissue engineering, 130–1 collagen scaffolds, 130–1 electrospinning and myotube formation, 131, 132 immunofluorescent human skeletal muscle cells, 132 mutual recognition procedure (MRP), 71 MWCNTs see multiwalled carbon nanotubes myotendinous junction (MTJ), 148 myotube formation, 131, 132 nanofibres, 323–8 as extracellular matrix analogue, 303–4 nanotechnology, 284 natural polymers, 36–41, 211 degradation, 41 polysaccharides, 41 proteins, 41 sources and structure, 37–40 polyesters, 38–40 polysaccharides, 38 proteins, 37 uses, 40–1 polysaccharides, 41 proteins, 40–1 nephrin, 344 nerve conduit, 181–3, 184 electrospinning set up for nanofibrous tubes fabrication, 183 nanofibrous nerve construct macrographs, 184 nerve growth factor, 179 nerve tissue engineering, 171–81 cells, 175, 177–9 mammalian stem cells classes, 178 current materials and scaffolds, 175, 176



6/6/11 6:53:01 PM

Index materials used to fabricate nerve conduits, 176 growth factors, 179–81 neurotrophic factors for peripheral nerve regeneration, 180 native nerve tissue, 171–4 extracellular matrix molecules description of peripheral nervous system, 173 peripheral nerve anatomical layers, 173 nerve tissue regeneration, 168–93 biomimetic nanoscaffolds for peripheral nerve regeneration, 181–7 bands formation of Bungner, 185 fillings and scaffolds in nerve conduits lumen, 185 intraluminal guidance channel, 183–7 nerve conduit, 181–3, 184 clinical problems in nerve tissue therapy, 169–71, 172 central nervous system, 169–70 peripheral nerve system, 170–1, 172 stem cell treatment for central nervous system diseases, 170 nerve tissue engineering, 171–81 cells, 175, 177–9 current materials and scaffolds, 175, 176 extracellular matrix molecules description of peripheral nervous system, 173 growth factors, 179–81 mammalian stem cells classes, 178 materials used to fabricate nerve conduits, 176 native nerve tissue, 171–4 neurotrophic factors for peripheral nerve regeneration, 180 peripheral nerve anatomical layers, 173 perspective, 192–3 stem cell therapy with nanofibre for nerve regeneration, 187–92 immunostained neurofilament laser scanning confocal microscopy micrographs, 190

403

nanofibres effects on neuronal differentiation, 189–92 PLLA electrospun nanofibres micrographs, 191 stem cells neuronal differentiation, 187–9 Nes2y cells, 359, 362 neural stem cells, 177 NGF see nerve growth factor non-steroidal anti-inflammatory drugs (NSAIDs), 151 NSC see neural stem cells osteoarthritis, 112 osteoblasts, 95–6 osteocalcin, 95 osteoclasts, 96–7 osteoporosis, 101–2 osteoprotegrin (OPG), 96 osteotendinous junction (OTJ), 148 PAA see polyacrylic acid pancreatic research, 359–70 future trends, 369–70 methods, 363–4 analysis of cells on scaffolds, 364 coated coverslips preparation, 364 electrospun scaffolds preparation, 363–4 Min6 cell line, 361–2 Nes2y cells, 362 novel scaffolds and production methods, 362–3 results, 364–8 Min6 cell equivalence after 48 hours cell culture, 367 Min6 cells growth on different surfaces, 364–5, 366, 367 Min6 seeded on PCL scaffolds and coverslip, 366 Nes2y cell equivalence after 48 cell culture, 368 Nes2y cells growth on different surfaces, 365–8 Nes2y cells on PCL scaffold and coverslip, 367 scaffold morphology, 364 uncoated and coated PCL scaffold, 365 papillary dermis, 300



6/6/11 6:53:01 PM

404

Index

Parkinson’s disease, 169 PCL see poly (e -caprolactone) PEGylation, 231 percutaneous operative, 151–2 periodontal regeneration, 281–4 aligned electrospun polycaprolactone scaffolds micrographs, 283 electrospun scaffolds micrographs, 282 peripheral nerve system, 170–1, 172 list of commercially available artificial nerve grafts, 172 phase separation, 325 PLA see polylactic acid plasma treatments, 236 plastic, 34 PLGA see poly(lactic-co-glycolic acid) PLLA see poly(L-lactic acid) podocytes, 343 Poisson modulus, 252 poly (a -hydroxy acids), 211–12 poly (e -caprolactone), 19, 100, 120, 121, 155, 205, 363 fibres functionalised with PVPA, 103 micrographs and fibre diameter, 162 nanofibre micrographs, 161 surface functionalised with PVPA, 104 tensile properties, 161 tensile testing data for 3D electrospun bundles, 163–4 poly-lactide-co-glycolide (PLGA), 309–10 poly (vinyl phosphonic acid-co-acrylic acid), 102 effect on bone formation, 103–5 effect on bone resorption, 105 structural formula, 102 polyacrylic acid, 139, 140 polyanhydrides preparation, 44 anhydride group, 44 condensation polymerisation, 44 polyaniline, 133 polycaprolactone (PCL), 309 polycaprolactone/poly(trimethylene carbonate) (PCL–PTMC), 310 polycarbonates, 44 polytrimethylene carbonate synthesis, 44 poly(d,L-lactide-co-glycolide), 253 polyesters preparation, 42–3

caprolactone, 43 lactide monomers, 43 polyethylene terephthalate and polycaprolactone, 43 sources and structure, 38–40 polyethylene glycol (PEG), 362 poly(ethylene oxide) (PEO), 7 polyethylenimines, 135 polyglycolic acid, 152, 175, 246 poly(L-lactic acid), 120, 189, 191 polylactic acid, 19, 253 poly(lactic acid-co-glycolic acid), 211–12 poly(lactic-co-glycolic acid), 152, 182 poly(lactide-co-glycolide), 120, 121 polymer chemistry, 34–48 degradable polymers, 35–6 natural polymers, 36–41 degradation, 41 sources and structure, 37–40 uses, 40–1 polymers, 34–5 synthetic degradable polymers, 41–8 common functional groups, 42 degradation, 45–8 preparation, 42–4 uses, 45 polymer composite fibres, 15 polymerase chain reaction (PCR), 268 polymers, 34–5 choice, 56 polystyrene formation by free radical polymerisation, 35 polysaccharides, 38, 211 degradation, 41 sources and structure, 38 basic hexose sugar rings, 39 cellulose, 40 sucrose, 40 uses, 41 polystyrene scaffold, 230 polytetrafluoroethylene (PTFE), 175 porous implants, 253 primary anastomosis, 245 primary wound healing, 317 product development, 70 proteins degradation, 41 sources and structure, 37 uses, 40–1 proteoglycans, 150



6/6/11 6:53:01 PM

Index pseudoislets, 361 PVPA see poly (vinyl phosphonic acid-coacrylic acid) RANK-L, 96 regulatory approval, 70 regulatory frameworks electrospinning, 67–87 medical products, 67–8 tissue regeneration products, 68–9 European regulatory framework, 69–84 medical devices, 72–5 medical products, 69–71 medicinal products, 71–2 regulatory input into development projects, 82–4 future trends, 84–7 regenerative medicine products, 86–7 materials in regenerative medicine, 69–84 tissue regeneration products, 75–82 devices and medicines, 75–80 gamma-sterilised medical devices, 81 materials regulated, 75 standards and guidelines, 80–2 resorptive pit, 97 reticular dermis, 300 satellite cells, 128 scaffolds, 77, 153, 211–17, 230–3 bioactive scaffolds, 217 bone tissue engineering fabrication, 97–9 materials, 99–101 3D electrospun heart valve scaffolds, 212 fibre alignment, diameter and morphology, 213–17 PLA fibres electrospun pictures, 216 material selection, 211–12 natural/biologically derived polymer scaffold, 232–3 properties, 205–17 balancing scaffold degradation and tissue formation, 212–13 to meet tissue requirements, 212–17 specific tissue requirements as a

405

blueprint for scaffold properties, 205–11 synthetic materials, 204–5 synthetic polymer scaffolds, 230–1 tendon tissue engineering, 154 scanning electron microscopy, 256–7, 347–8 Schwann cells, 172, 173, 177 SCI see spinal cord injury secondary wound healing, 317 self-certification, 73 semilunar heart valves, 206–11 cell–matrix interactions, 208–11 geometry and function, 206–7 haemodynamic parameters that apply on the aortic valve, 207 human aortic heart valve, 206 structure and composition, 207–8, 209 fibrosa, spongiosa and ventricularis configuration within the aortic valve leaflet, 209 Shandom ‘Microtome’ 5030, 260 shikonin, 310, 334 silica nanofibres, 285 silicate glasses, 284 silicone, 175 silk, 36–7, 211 silk fibroin, 309 Silvasorb, 321 silver, 320–1 silver nitrate, 320, 321 silver sulfadiazine, 321 size effect, 334 skeletal muscle, 127 injuries, 128 soft tissue exploration/debridement and amputations, 128 mechanical properties, 129–30 elastic modulus and yield stress, 129 skin tissue regeneration, 298–312 biology of skin and wound healing, 299–301 cellular interactions on skin substitute, 310–11 choice of biomaterial, 308–10, 311 biomaterial nanofibrous scaffold used for skin regeneration, 311 future trends, 311–12 ideal scaffold properties, 304–8 antimicrobial properties, 305



6/6/11 6:53:01 PM

406

Index

biodegradability, 307 cell adhesive properties, 308 chitosan and chitosan-gelatine nanofibrous scaffolds, 306 mechanical properties, 305–6 porosity, 306–7 surgical handleability, 307 wettability, 304–5 wound healing properties, 307 nanofibres as extracellular matrix analogue, 303–4 problems in existing therapies, 301–2 restoring functional skin tissue, 302–3 skin layers from rat species, 299 slit diaphragms, 344 solvent, 55–6 PCL fibres spun with HFIP, 55 PCL showing beading spun, 55 spinal cord injury, 168 spinneret, 61 spinning distance, 57 PLGA fibres spun in 20% w/v solution with HFIP over 25 cm with 25kV potential, 59 stem cell research, 372–90 combining ES cells with electrospun scaffolds, 385–9 differentiation, 387 maintenance of pluripotency, 385–7 mouse ES cells culture on electrospun PLGA meshes, 386 three-dimensional culture system creation, 388–9 current culture techniques, 375–84 culture of mouse ES cells, 375 differentiation, 382–4 EB formation, 382 microenvironment of human ES cells in culture, 377 microenvironment of mouse ES cells in culture, 376 3D scaffolds, 384–5 embryonic stem cells, 373–5 derivation and definition, 373–5 mouse embryonic cells derivation, 374 future trends, 390 human ES cells culture, 379–81 adhesive substrate, 380–1 media composition, 379–80

stem cells, 116–19 stenosis, 203 sural nerve, 171 surface tension, 54 surgical débridement, 319 synthetic degradable polymers, 41–8 common functional groups, 42 degradation, 45–8 abiotic, 46–8 biotic, 48 preparation, 42–4 polyanhydrides, 44 polycarbonates, 44 polyesters, 42–3 uses, 45 synthetic polymers, 230–1 tartrate resistant alkaline phosphatase, 96 Taylor cone, 5 technical standards, 80 Tegaderm Ag, 321 templates, 20 tendinopathy, 151 tendon tissue regeneration, 148–64 cell response to electrospun bundles, 155, 157–8, 159–60, 161 clinical need, 152–3 electrospun bundles mechanical properties, 158–9, 161–4 future trends, 160, 163 pathology, 151–2 poly (e -caprolactone) micrographs and fibre diameter, 162 nanofibre micrographs, 161 tensile properties, 161 tensile testing data for 3D electrospun bundles, 163–4 structure and composition, 148–51 cells, 150 extra cellular matrix, 150–1 hierarchical layers, 149 pathology, 151–2 structure and composition, 148–51 tissue, 148 tenocytes SEM micrographs cultured on 3D electrospun bundles, 159 cultured on 3D electrospun bundles that have been plaited together, 160



6/6/11 6:53:01 PM

Index tissue engineering, 153–5, 156–8 electrospinning process and fibre manipulation, 158 material selection – tendon repair, 154–5 nanofibres effect on cell behaviour, 155 research groups overview to developed tendon tissue engineered, 156–7 tendonitis, 151 tenocyte, 150 micrographs cultured on 3D electrospun bundles, 159 micrographs cultured on 3D electrospun bundles that have been plaited together, 160 TET see tissue engineered trachea thermally induced phase separated scaffolds (TIPS), 388–9 tissue engineered product, 76 tissue engineered trachea, 245–52 tissue engineering, 17–20, 75–6, 298 tissue inhibitors of metalloproteinases (TIMPs), 209 tissue regeneration bladder, 225–38 disease and the need for bladder substitution, 227–8 electrospinning fit for purpose, 234–7 electrospun and other scaffolds for bladder tissue engineering, 228–33 future trends, 237 structural/functional properties, 225–7 bone, 93–106 biology principles, 94–7 future trends, 105–6 materials for scaffolds, 99–101 osteoporosis, 101–2 scaffolds fabrication, 97–9 strategies, 97 treatment strategies, 102–5 cartilage, 111–24 culture of chondrogenic cells for implantation, 113–19 electrospun nanofibre scaffolds, 119–24

407

future trends, 124 heart valve, 202–18 future trends, 217–18 scaffold material selection, 211–12 scaffold properties, 212–17 specific tissue requirements for scaffold properties, 205–11 tissue replacement, 203–5 muscle, 127–42 collagen scaffolds, 130–1 conductive elements, 133–40 contractile force, 131–3 electrospinning and myotube formation, 131, 132 future trends, 140–2 immunofluorescent human skeletal muscle cells, 132 mechanical properties, 129–30 skeletal muscle injuries, 128 tissue engineering, 130–1 nerve, 168–93 biomimetic nanoscaffolds, 181–7 clinical problems in nerve tissue therapy, 169–71, 172 perspective, 192–3 stem cell therapy with nanofibre for nerve regeneration, 187–92 product regulation, 68–9, 75–82 devices and medicines, 75–80 gamma-sterilised medical devices, 81 materials regulated, 75 standards and guidelines, 80–2 skin, 298–312 biology of skin and wound healing, 299–301 cellular interactions on skin substitute, 310–11 choice of biomaterial, 308–10, 311 future trends, 311–12 ideal scaffold properties, 304–8 nanofibres as extracellular matrix analogue, 303–4 problems in existing therapies, 301–2 restoring functional skin tissue, 302–3 tendon, 148–64 cell response to electrospun bundles, 155, 157–8, 159–60, 161



6/6/11 6:53:01 PM

408

Index

clinical need, 152–3 electrospun bundles mechanical properties, 158–9, 161–4 future trends, 160, 163 pathology, 151–2 structure and composition, 148–51 tissue engineering, 153–5, 156–8 trachea, 242–76 anatomy, 242–4 electrospun biodegradable tubular tracheal scaffold, 252–8 scaffold fulfilment, 258–63 tissue engineered trachea, 245–52 Topcon SM300 SEM, 347 tracheal tissue regeneration, 242–76 cell and tissue response in vitro and in vivo evaluation, 263–75 chondrocytes adhesion and infiltration micrographs, 266 complex vascular pedicle scaffold macroscopic evaluation, 271 Degrapol scaffold images after implantation, 272 excised tracheal scaffold semi-thin cross-section, 274 gene expression for chondrocytes at two different time points, 269 histological analysis, 273 scaffold surface immunocytochemical images, 268 static culture micrograph for chondrocytes, 265 vascular axis formed by the carotid artery and internal jugular vein of rabbit, 270 in vitro dynamic culture, 265–9 in vitro experimental procedures, 269–75 in vitro static culture, 265, 266 DegraPol, 255–8, 259 characteristics, 256 chloroform DegraPol solutions, 257 collector spindle, 257 degradable block polyesterurethane for tissue engineering, 255 degradation in buffered aqueous solution, 256 electrospun scaffold section image with fibre spacings, 258 electrospun tubular scaffold, 256–8

operational conditions and solution properties in electrospun reticulate fibres production, 259 scaffold surfaces obtained under different thermic conditions, 257 electrospun biodegradable tubular tracheal scaffold, 252–8 bidimensional model for fibreporous meshes, 254 encapsulation dependence on fibres spacing, 254 scaffolds and tissue engineering, 252–5 ESEM images scaffold culture in a bioreactor, 267 scaffold microstructure at the grooves bottom, 268 made by electrospinning trachea-scaffold type 1, 261 trachea-scaffold type 2, 262 trachea-scaffold type 3, 263 scaffold fulfilment, 258–63, 264 inner degrapol electrospun microporous layer image, 264 inner layer, 262–3 ringed profile development, 259–60, 261–2 ringed profile image showing scaffold level section, 261 scaffold models and relevant measurements investigated for tracheal scaffolds, 260 scaffold profile selection, 258–9, 260 spiral profile development, 259, 261 toothed profile development, 260, 262, 263, 264 surgical concern pathologies, 244 tissue engineered trachea, 245–52 first transplant of a human tissue engineered tracheal prosthesis, 250 general approaches used for tracheal replacement, 246 helical template with silicone moldmaking kit and chondrocyteseeded matrix, 248 human tracheal cartilage mechanical features, 252



6/6/11 6:53:01 PM

Index tissue engineering and tracheal replacement, 245–50 tracheal cartilage, structure and biomechanics, 250–2 tracheal cartilage zones images, 251 tracheal chondrocytes appearances, 247 tracheal anatomy, 242–4 human trachea, 243 tracheal wall ultrastructural section, 244 transmission electron microscopy (TEM), 271 TRAP see tartrate resistant alkaline phosphatase ultimate tensile strength (UTS), 158 urothelium, 226–7 valvular interstitial cells (VICs), 211 vascular endothelial growth factor (VEGF), 237 viscosity, 52–3 wound, 317 wound dressings, 317–34 antimicrobial nanofibrous wound dressings, 328–34 cumulative release of silver from nanofibrous alginate samples, 333

409

electrospun alginate/Ag nanoparticle nanofibres, 330 nanofibres with other therapeutics, 329–3 nanofibres with silver nanoparticles, 328–9 zone-of-inhibition assay for electrospun alginate discs, 331–2 antimicrobials, 320–3 ionic silver, 320–2 nanocrystalline silver, 322 silver nanoparticles, 322–3 currently in widespread use, 320 nanofibres, 323–8 advantages and disadvantages of fabrication techniques for medical applications, 326 advantages in wound dressing applications, 324–5 electrospun nanofibres, 327–8 fibre diameter effect on surface area of fibrous materials, 324 nanofibre fabrication methods, 325–7 wound healing, 317–23 healthy wounds, 317–18 non-healing wounds, 318–19 wound dressings, 319–20 wound healing, 300 Young’s modulus, 158, 252



6/6/11 6:53:01 PM


6/6/11 6:53:01 PM

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