Bioceramic s Volume 10
Edited by Lauren t Sedel Laboratoire de Recherches Orthopediques Faculte de Medecine Lariboisiere-Saint Louis Universite Paris 7 - Denis Diderot Paris, France
Christia n Rey Ecole Nationale Superieure de Chimie Institut National Polytechnique de Toulouse Toulouse, France
PERGAMON
Pergamon/Elsevie r Titles of Relate d Interes t
BIOCERAMICS BOOK SERffiS Bioceramic s 4: W. BONFELD, G. W. HASTINGS and K. E. TANNER (ISBN 0 7506 0269 4) Bioceramic s 6: P. DUCHEYNE and D. CHRISTIANSEN (ISBN 0 08 042143) Bioceramic s 7: O.H. ANDERSSON, R.-P. HAPPONEN and A. YLI-URPO (ISBN 0 08 042144 X) Bioceramic s 8: J. WILSON, L. L. HENCH and D. GREENSPAN (ISBN 0 08 0426778) Bioceramic s 9: T. KOKUBO, T. NAKAMURA and F. MIYAJI (ISBN 0 08 0426840) RELATED JOURNAL Biomaterials(ISSN 0142-9612) For details on the above Pergamon/Elsevier Science book series or a free specimen copy of any Elsevier Science journal please contact your nearest Elsevier Science office (see copyright page for addresses).
Bioceramic s Volume 10 Proceedings of the 10th International Symposium on Ceramics in Medicine, Paris, France, 5-9 October, 1997 Editedby Lauren t Sedel Christia n Rey
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Elsevier Science Ltd, The Boulevard, Langford Lane, Kidlington Oxford 0X5 1GB, U.K.
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Elsevier Science Inc., 655 Avenue of the Americas, New York 10100, U.S.A.
JAPAN
Elsevier Science Japan, Tsunashima Building Annex, 3-20-12 Yushima, Bunkyo-ku, Tokyo 113, Japan
Copyright ' 1997 Elsevier Science Ltd All Rights Reserved No part of thispublication may be reproduced, stored in a retrievalsystem or transmittedin anyform or by any means: electronic, electrostatic, magnetic tape, mechanical, photocopying, recording or otherwise, without permission in writing from the publishers. First Edition 1997 Librar y of Congres s Catalogin g in Publicatio n Dat a A catalog record for this book is available from the Library of Congress Britis h Librar y Cataloguin g in Publicatio n Dat a A catalog record for this book is available from the British Library ISBN 0 08 0426921
Cover picture BOURDALOU Factory of Comte d’ Artois Limoges - End of XVIII century Adrien Dubouche National Museum, Limoges
Printed in Great Britain at the University Press, Cambridge
Organizin g Committe e Chairme n
Laurent Sedel Christian Rey
Member s
Alain Meunier Didier Bernache-Assollant
Secretar y
Martine Henry-Amar
Scientific Committe e D. Bernache-Assollant J. Delecrin P. Frayssinet M. Hamadouche P. Marie A. Meunier A. Moroni
M. Nardin R. Nizard N. Passutti H. Petite S. Redey L’H. Yahia
Internationa l Advisor y Committe e A. Barbosa (Portugal) P. Boch (France) W. Bonfield (UK) G. Daculsi (France) P. Ducheyne (USA) P. Griss (Germany) K. de Groot (Netherlands) U. Gross (Germany) L.L. Hench (USA)
S.F. Hulbert (USA) T. Kokubo (Japan) H. Oonishi (Japan) J.A. Planell (Spain) A. Ravaglioli (Italy) C. Rey (France) J. Wilson (USA) T. Yamamuro (Japan) A. Yli-Urpo (Finland)
Scientific Endorsemen t European Society for Biomaterials Society for Biomaterials (USA) European Society for Orthopaedic Research Japanese Society for Biomaterials Ceramics Society of Japan Japanese Society of Orthopaedic Ceramic Implants National Institute for Scientific and Medical Research (I.N.S.E.R.M., France) National Centre for Scientific Research (C.N.R.S., France)
This Page Intentionally Left Blank
Prefac e The 10th Bioceramics volume reports on the meeting held in Paris from 5th to 8th October 1997. It has been a great pleasure for me to edit this book with the highly respected scientist, Christian REY. Bioceramics in Medicine has become one of the major fields in biomaterials. Many different bioceramics, with various material compositions and characteristics producing different biological behaviour, result in a wide range of medical applications. Dense ceramics, such as aluminium oxide, present a range of properties which should allow the material to last for many years within the living body. This highly oxidised material should also theoretically resist corrosion related to oxygen capture. The latter exhibits very interesting tribological properties when compared to either metal or polyethylene, which are more widely employed. Bioceramics made of calcium phosphate of different chemical compositions texture and porosity, are of great interest for the replacement of defective bone lost through tumour excision or major trauma or to enhance bone repair after fracture, fusion or ridge augmentation. Calcium phosphate ceramics have been employed to copy bone mineral. The imitation of bone mineral is very difficult as it presents a complex system within which formation, degradation and even texture are not fully understood. However, artificial calcium/phosphate of various compositions and even textures has been proved to be bone friendly. Osteoconduction has been demonstrated by many of these materials and their capacity to provide a scaffold for bone cells is also well recognised. Is this related to the three dimensional aspects of these materials and to the ability of mesenchymal stem cells to grow on this support? Or is osteoconduction only related to the chemical composition of the ceramic? Even if the answers to these questions are not available at the present, results confirm the excellent osteoconductivity of calcium/phosphate ceramics. Further questions need to be addressed: what are the events following bone apposition on the surface of the ceramic? Will this material resorb allowing natural bone to replace the artificial material, providing either an excellent bone union or bone apposition. Or will this material remain unchanged providing a permanent bone apposition or bone augmentation necessary for a shelf procedure ? Bioglasses, containing different types of silica with phosphorus, represent another very interesting field. These materials bond to bone without any visible interface. Do these bioglasses provide enough bone bonding strength? Do they enhance bone union? Experimental results appear to be positive, but confirmation in the human
viii Preface
body is yet to come. Mixed materials such as apatite wollastonite glass ceramics (AWGC) show both remarkable tolerance as well as strength and an ability to remain unchanged for many years, when implanted in the living body. As well as clinical applications, fundamental subjects are also discussed. What is the mechanism of bone cell osteoconduction? Is it related to chemical or physical factors, or both? What is the role of the micro and the macro geometry of the surface? Would it be possible to imitate these characteristics on other materials, such as metals or polymers? All these questions and many others are discussed in this book. The applications of bioceramics are numerous. Orthopaedic and dental surgery are the two major fields of interest but others such as plastic surgery, E.N.T., problems of percutaneous devices, embolisation materials, calcium phosphate cement are also increasingly important. Finally the bioceramics field is one of the most active biomaterials applications and has grown considerably in the last twenty years. Its high level scientific input from a wide range of backgrounds and the participation of industries create a great interest in this subject. We hope that everybody, whatever his own interest, will find something of value in this book.
Laurent SEDEL
CONTENTS CALCIU M PHOSPHAT E IN VIVO FORMATIO N Precipitatio n of Calcium Phosphat e on Titani a Ceramic s K-L. Eckert, S.-W. Ha, S. Ritter and E. Wintermantel Apatit e Formatio n on Polymer s by Biomimeti c Processin g Using Sodium Silicate Solution F. Miyaji, S. Handa, T. Kokubo and T. Nakamura Bonelike Apatit e Formatio n on the Surfac e of ChemicaU y Treate d Tantalu m Substrates : Effect of Heat Treatmen t T. Miyazaki, H.M. Kim, F. Miyaji, T. Kokubo and T. Nakamura
3
7
11
Influenc e of Hydroxyapatit e Particl e Size and Morpholog y on Hapex^ ^ M. Wang, R. Joseph and W. Bonfield
15
Effect of Fluorid e Substitutio n on the Biocompatibilit y of Hydroxyapatit e K.A. Ring, L. Di-Silvio, I.R. Gibson, C. Ohtsuki, L.J. Jha, S.M. Best and W. Bonfield
19
Apatit e Precipitatio n in Biphasi c Calcium Phosphat e Cerami c After Incubatio n in Rabbi t Serum and Ionic Simulate d Body Fluid (SBF) R. Rohanizadeh, J.M. Bouler, D. Couchourel, M. Padrines and G. Daculsi
23
Apatit e Precipitatio n in Biphasi c Calcium Phosphat e Cerami c After Implantation : Influenc e of Implantatio n Site R. Rohanizadeh, M. Trecant-Viana, J. Delecrin, J.M. Bouler and G. Daculsi
27
GLAS S CERAMIC S BIO ACTIVIT Y Bioactivit y and Structur e of Organicall y Modified Silicate Synthesize d by the Sol-Gel Metho d K. Tsuru, S. Hayakawa, C. Ohtsuki and A. Osaka Hydroxyapatit e Formatio n on Bioactive Glass Coate d Titaniu m C.Y. Kim and S. Kwon
33
37
X Contents
Effect of Multivalen t Cation s in Calcium Silicate Glasses on Bioactivit y N. Imayoshi, C. Ohtsuki, S. Hayakawa and A. Osaka
41
Transformatio n of Bioactive Glass Granule s into Ca-P Shells In Vitr o S. Radin, P. Ducheyne, S. Falaize and A. Hammond
45
Multilayere d Coating s of Hydroxyapatite/Glas s Cerami c Composite s Plasm a Spraye d on Ti-6A1-4V Alloy P.L. Silva, J.D. Santos and FJ. Monteiro
49
The Bony Reactio n to Rapidl y Degradabl e Glass-Ceramic s Based on the New Phas e Ca2KNa(P04) 2 C. Miiller-Mai, G. Berger, C. Voigt, B. Bakki and U. Gross
53
Resorbable , Porou s Phosphat e Inver t Glasses - Firs t in Vitr o and in Vivo Result s J. Vogel, K.-J. Schulze, D. Reif, P. Hartmann, U. Platzbecker and B. Leuner
57
Implantatio n of Bioactive and Iner t Glass Fibre s in Rat s - Soft Tissue Respons e and Short-Ter m Reaction s of the Glass M. Brink, P. Laine, K. Narva and A. Yli-Urpo
61
Spina l Fusion Using Titaniu m Spacer s With Bioglass^ and Autogenou s Bone: A Comparativ e Stud y in Sheep J. Wilson, G. Lowery and S. Courtney
65
DENSE AND POROU S BIOACTIV E CERAMIC S Macroporou s Biphasi c Calcium Phosphat e Ceramics : Influenc e of Macropor e Diamete r and Macroporosit y Percentag e on Bone Ingrowt h O. Gauthier, J-M. Bouler, E. Aguado, P. Pilet and G. Daculsi Mechanica l Fatigu e of Hot Presse d Hydroxyapatit e S. Raynaud, E. Champion, D. Bernache-Assollant Mechanica l Propert y Change s in Macroporou s Cerami c After Implantatio n into Bone and Muscle M. Trecant-Viana, J. Delecrin, J.M. Nguyen, J. Royer and G. Daculsi
71
75
79
Contentsxi
Differenc e of Bonding Behavior Between Four Differen t Kinds of Hydroxyapatit e Plat e and Bone S.S. Chung, C.K. Lee, K.S. Hong and H.J. Yoon
83
Treatmen t of Osteomyeliti s by Antibiotic-Soake d Porou s A-W Glass Cerami c Block K. Kawanabe, Y. Okada, H. lida and T. Nakamura
87
Calcium Hydroxyapatit e Cerami c Implant s Impregnate d With Antibioti c for the Treatmen t of Chroni c Osteomyeliti s Y. Yamashita, T. Yamakawa, K. Kato, Y. Shinto, N. Araki and A. Uchida
91
BONE CELL S ONT O BIOACTIV E CERAMIC S Measuremen t of Intac t Osteocalci n Content s in the Composit e of Porou s Hydroxyapatit e Cerami c and Allogeneic Marro w Cells M. Akahane, H. Ohgushi, T. Yoshikawa, S. Tamai, Y. Dohi, K. Hosoda and T. Ohta Si-Ca-P Xerogels and Bone Morphogeneti c Protei n Act Synergisticall y on Rat Stroma l Marro w Cell Differentiatio n In Vitr o E.M. Santos, P. Ducheyne, S. Radin, B. Shenker and L Shapiro Effect of Surfac e Instabilit y of Calcium Phosphat e Ceramic s on Growt h and Adhesion of Osteoblast-Lik e Cells Derived from Neonata l Rat Calvari a T. Suzuki, M. Hukkanen, L.D.K. Buttery, J.M. Polak, Y. Yokogawa, K. Nishizawa, F. Nagata, Y Kawamoto, T. Kameyama and M. Toriyama Histologica l Evaluatio n of Culture d Bone Graf t Using Cryopreserve d Marro w CeUs H. Nakajima, T. Yoshikawa, H. Ohgushi, M. Akahane, S. Tamai, K. Mishima and K. Ichijima Interaction s of Bioceramic s on Huma n Osteoarthriti s (OA) Type B Synoviocytes. Effects on Interleuki n Levels and Lipoxygenas e Pathway s B. Liagre, J-L. Charissoux, M-J. Leboutet, D. Bernache-Assollant and J-L. Beneytout A Long Term Implantatio n of Culture d Bone in Porou s Hydroxyapatit e T. Yoshikawa, H. Ohgushi, H. Nakajima, M. Akahane, S. Tamai and K. Ichijima
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101
105
109
113
117
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Contents
SURFAC E BEARIN G CERAMIC S Oxide Ceramic s for Articulatin g Component s of Tota l Hip Replacement s G. Willmann
123
Ex Vivo and In Vitr o Analysis of the Alumina/Alumin a Bearin g System for Hip Join t Prosthese s H.J. Refior, W. Plitz and A. Walter
127
Hybri d Alumina-Alumin a Hip Replacement : A Survivorshi p Analysis and Result s at a Minima l Five Year FoUow-Up M. Hamadouche, P. Bizot, R.S. Nizard and L. Sedel
131
Low Temperatur e Ageing Behaviour of Zirconi a Hip Join t Head s J. Chevalier, J.M. Drouin and B. Cales Characterizatio n of Zirconi a Coate d by Bioactive Glass: Preliminar y Observation s M. Bosetti, M. Santin, M. Mazzocchi, A. Krajewski, M. Rastellino, A. Ravaglioli and M. Cannas
135
139
Calcium Phosphat e Particle s are Found at the Polyethylen e Inser t Surfac e Whethe r Implante d With Ha-Coate d Devices or Not. A SEM-EPM A Study P. Frayssinet, L. Gineste, G. Bonel and N. Rouquet
143
Bone Remodellin g Aroun d Implante d Material s Under Load-Bearin g Condition s M. Oka, Y.S. Chang, S. Yura, K. Ushio, J. Toguchida and T. Nakamura
147
CLINICA L USE OF CERAMIC S Clinica l Comparativ e Stud y Between Porou s Coate d and Hydroxyapatit e Porou s Femora l Implant s Y.H. Kim, J.H. Shon and I.Y. Choi
153
Revision Rate s and Radiographi c Change s Associated With Differen t Socket Interfac e Technologie s : Clinica l Result s from 416 Patient s at 6 To 8 Year s FoUow-Up M.T. Manley, A. Edidin, J. A. Epinette, R.G. Geesink, J.A. D’Antonio and W.N: Capello
157
Contents xiii
Comparativ e Stud y of the Result s Between Custom Non-Coate d Cementles s Hip Implant s and Mirrore d Cementles s HA-Coate d Hip Implant s on the Contra-Latera l Side M. Mulier and G. Deloge
161
Improvemen t of THR With Spongiosa Meta l Surfac e Using the Wear Couple Ceramic-On-Cerami c G. Quack, G. Willmann, H.G. Pieper and H. Krahl
165
Acetabula r Reconstructio n in Revision Tota l Hip Arthroplast y Using a Bone Graf t Substitut e R.P. Pitto and D. Hohmann
169
POSTE R 1 Effect of Solution Ageing on Sol-Gel Hydroxyapatit e Coating s B. Ben-Nissan, C. Chai and K.A. Gross Ionic Cement s : Influenc e of the Liquid/Soli d Rati o on Porosit y and Mechanica l Propertie s F. Betchem, P. Michaud, F. Rodriguez and Z. Hatim
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Sinterin g and Therma l Decompositio n of Hydroxyapatit e Bioceramic s J. Cihlar and M. Trunec
183
An Elaboratio n of the New Dissolution Mechanis m for Apatit e S.V. Dorozhkin
187
Ultrastructura l Study of Long Term Implante d Ca-P Particulat e into Rabbi t Bones A. Dupraz, R. Rohanizadeh, J. Delecrin, P. Pilet, N. Passuti and G. Daculsi
191
Bioactive Glass-an d Glass-Cerami c Composite s and Coating s M. Ferraris, E. Yerne, A. Ravaglioli, A., Krajewski, L. Paracchini, J. Vogel, G. Carl, C. Jana
195
Mechanica l Characterisatio n of Bioactive Coating s on Zirconi a E. Verne, M. Ferraris, C. Moisescu, A. Ravaglioli and A. Krajewski
199
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Contents
Vacuu m Plasm a Spraye d Titaniu m and Hydroxyapatit e Coating s on Carbo n Fiber Reinforce d Polyetheretherketon e (Peek )
203
S.-W. Ha, A. Gisep, H. Gruner, J. Mayer and E. Wintermantel Low Temperatur e Crystallizatio n of Hydroxyapatit e Sputtere d Films in an Autoclav e
207
J. Hamagami, K. Nakamura, Y. Sekine, K. Yamashita and T. Umegaki Fabricatio n of In-Cera m Cor e by Sheet Formin g Proces s
211
D-J . Kim , M-H . Lee and C.E . Kim Surfac e Structur e of Bioactive Titaniu m Prepare d by Chemica l Treatmen t
215
H.M. Kim, F. Miyaji, T. Kokubo, T. Suzuki, , F. Itoh, S. Nishiguchi and T. Nakamura Prefabricate d Biological Apatit e Formatio n on a Bioactive Glass-Cerami c Promote s In Vitr o Differentiatio n of Feta l Rat Chondrocyte s
219
C. Loty, S. Loty, T. Kokubo, N. Forest and J.M. Sautier The Effect on Mechanica l Propertie s by Osteoblasti c Cell Ingrowt h in Macroporou s Syntheti c Hydroxyapatit e and Interpor e 200 ^ ^
223
E. Nordstrom, H. Ohgushi, H. Yoshinari, S. Tamai and T. Yokobori Characterizatio n and Cell Reactio n of a-TCP - and HAp-Coating s on Titaniu m Plat e
229
M. Ohgaki, S. Nakamura and M. Akao The Ectopi c Osteoconductio n Model
233
H. Ohgushi, M. Okumura, T. Yoshikawa, H. Ishida, H. Yajima and S. Tamai Condition s of the Coprecipitatio n of Calciu m Hydroxyapatit e With Zr02 , ZrO i " Y2O3, AI2O3 from Aquoeu s Solution s Using Ammoni a
237
V.P. Orlovskii, Zh A. Ezhova and E.M. Koval Transformatio n of a-TC P to Hydroxyapatit e in Organi c Media
241
K. Sakamoto, S. Yamaguchi, A. Nakahira, K. Kijima and M. Okazaki Structur e and Solvation Effects of P04^- , HP04^- , H2P04 ’ and H3PO 4 from AMI and PM 3 A.J. Sahnas, A. Serret, M. Vallet-Regi and L.L. Hench
245
Contents xv
The Detaile d Configuratio n of Carbonat e Ions in Apatit e Structur e Determine d by Polarize d Ir Spectroscop y
249
Y. Suetsugu, I. Shimoya and J. Tanaka Tissue Cultur e on Amorphou s Calciu m Phosphat e Coatin g
253
K. Suzuki, Y. Kageyama, Y. Kita, A. Yoshino, K. Matsushita and T. Kokubo Bonelike Apatit e Layer Forme d on Organi c Polymer s by Biomimeti c Proces s : TEM-ED X Observatio n of Initia l Stag e of Apatit e Formatio n H. Takadama, F. Miyaji, T. Kokubo and T. Nakamura Sinterabilit y and Second Phas e Formatio n of Syntheti c Hydrox y Apatit e : Controllin g Parameter s and Effect on Bond Strengt h
257
261
H-J. Youn, H.S. Ryu, K.S. Hong, S.S. Chung and C.K. Lee Porou s Sol-Gel Bioglassfi from Near-Equilibriu m Dryin g
265
J. Zhong and D.C. Greenspan
POSTE R 2 Ceramic-Cerami c Bearin g System s Compare d on Differen t Testin g Configuration s
271
J. ChevaHer, B. Cales, J.M. Drouin and Y. Stefani Dissolution and Mechanica l Behaviou r of Plasma-Spraye d Cerami c Coating s for Orthopaedi c Application s
275
N. Demonet, P. Benaben, J.L. Aurelle, B. Forest and J. Rieu Design of a Calciu m Phosphat e Bone Cemen t Suitabl e for the Fixatio n of Meta l Endoprosthese s
279
FCM. Driessens, L Khairoun, MG. Boltong and Ja. Planell Quantitativ e Compariso n of In Vivo Bone Generatio n With Particulat e Bioglassfi and Hydroxyapatit e as a Bone Graf t Substitut e
283
Y. Fujishiro, H. Gonishi and L.L. Hench Test of Bioactivit y in Four Differen t Glasse s A.M. Gatti, L.L. Hench, E. Monari, F. Gonella and F. Caccavale
287
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Contents
Compariso n of Bone-Implan t Attachmen t Strengt h Between the Implant s With Hydroxyapatite-Coatin g and Tricalciumphosphate-Coatin g on Titaniu m Arc Spraye d Titaniu m
291
K. Hayashi, T. Hara, T. Imamura and Y. Iwamoto Effect of Hydroxyapatit e Coatin g on Bony Ingrowt h into Groove d Titaniu m Implant s
295
K. Hayashi, T. Mashima, K. Uenoyama, T. Hara and Y. Iwamoto Mechanis m of the Inflammator y Reactio n of Conventiona l Calciu m Phosphat e Cemen t
301
K. Ishikawa, Y. Miyamoto, M. Nagayama and K. Suzuki Fractur e of Alumin a Cerami c Head in Tota l Hip Arthroplasty . - Repor t of a Case With Histologica l Examinatio n and Particl e Characterisatio n
305
Y. Kadoya, A. Kobayashi, P.A. Revell, H. Ohashi, Y. Yamano, G. Scott and M.A.R. Freeman Experimenta l Comparativ e Stud y Between Rough-Blaste d and Hydroxyapatit e Coate d Implant s
309
Y.H. Kim, J.S. Park, I.Y. Choi, M.R . Park and T.S. Park Mechanica l and Biological Propertie s of Alumin a Bead Composit e
313
M. Kobayashi, T. Nakamura, T. Kikutani, Y. Okada, N. Ikeda, S. Shinzato and T. Kokubo Remodelin g of Bone Aroun d Hydroxylapatite-Coate d Femora l Stem s
317
A.A. Edidin and M.T. Manley Processin g and Characterisatio n of Biological Hydroxyapatit e Derived from Cattle , Sheep and Deer Bone
321
M.R . Mucalo , G.S. Johnson and M.A. Lorie r Catastrophi c Wear of Meta l Ball of Bipolar Hip Prosthesi s After Fractur e of Alumin a Cerami c Screws Used for Acetabula r Bone Graf t
325
H. Ohashi, Y. Yutani, A. Kobayashi, Y. Kadoya and Y. Yamano Antibacteria l Propert y of Ag-Doped Calciu m Phosphat e Compound-Cellulos e Composite s K. Okada, Y. Yokogawa, T. Kameyama, K. Kato, Y. Kawamoto, K. Nishizawa, F. Nagata, M. Okuyama
329
Contents xvii
Wear Behaviour of Polyethylen e Cup Against 28 mm Alumina Ball in Tota l Hip Prosthese s H. Oonishi, N. Murata, S. Kushitani, S. Wakitani, K. Imoto, Y. Iwaki and N. Kin In Vitr o Cell Behavior of Osteoblast s on Pyros t Bone Substitut e J-S. Sun, F.H. Lin, Y-H. Tsuang, Y-S. Hang, C.Y. Hong and H.C. Liu The Efficacy of Hydroxyapatite-Tricalciu m Phosphat e Filler for Bone Defects Associated With Humera l Pseudoarthrosis : Compariso n With Autogenou s Iliac Bone Graft s K. Suzuki and M. Yamada
333
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341
Experimenta l Study of Apatit e Cement Includin g Cisplati n Y. Tahara, Y. Ishii, S. Sasaki, I. Takano and K. Ohzeki
345
In Vivo Evaluatio n of Sol-Gel Bioglassfi. - Biomechanica l Finding s D.L. Wheeler, R.G. Hoellrich, S.W. McLoughlin, D.L. Chamberland and K.E. Stokes
349
Fixatio n of Hip Prosthese s by Hydroxyapatit e Coatin g G. Willmann
353
Acetabula r Reconstructio n With an Artificia l Bone Block S. Yoshii, M. Oka, T. Yamamuro, H. lida, Y. Kakutani, K. Ikeda, H. Murakami and T. Nakamura
357
Participatio n of Calcium Phosphat e Cement s for Healin g of Alveolar Bone M. Yoshikawa, H. Oonishi, Y. Mandai, K. Minamigawa and T. Toda
361
BIOCERAMIC S SYNTHESI S AND EVALUATIO N Synthese s of Rapid Resorbabl e Calcium Phosphat e Ceramic s With High Macr o or High Micr o Porosit y G. Berger, R. Gildenhaar, U. Ploska and M. Willfahrt
367
Composit e Bioceramic s Mad e of Macroporou s Calcium Phosphat e Ceramic s FiUed With a Self-Settin g Cement . Histologica l Evaluatio n P. Frayssinet, A. Lerch, L. Gineste and N. Rouquet
371
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Contents
Physica l Propertie s of an Apatiti c Cerami c Containin g Tricalciu m Phosphat e Prepare d by the Way of a Cement Z. Hatim, M. Freche and J.L. Lacout Cytocompatibilit y of Calcium Phosphat e Coating s With Variou s Ca/P Ratio s P. Frayssinet, L. Arbore and N. Rouquet Compariso n of Resorptio n and Bone Conductio n of Two CaCO a Bone Substitute s J.C. Fricain, Ch. Baquey, B. Basse-Cathalinat and B. Dupuy Reliabilit y of Dual Energ y X-Ray Absorptiometr y in Evaluatio n of Phospho Calcic Bioceramic s in Rabbi t J.X. Lu, O. Legrand, B. Flautre, A. Gallur, M. Descamps, B. Thierry, P. Hardouin and B. Sutter The Evaluatio n of Degradabilit y of Melt and Sol-Gel Derived Bioglassfi InVitr o D.C. Greenspan, J.P. Zhong, X.F. Chen and G.P. LaTorre
375
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391
COMPOSIT E CERAMIC S Upgradin g of Hydroxyapatit e Cerami c Biocompatibilit y by Incorporatio n of a-Tricalciu m Phosphat e S. Sarig, F. Apfelbaum and F. Kahana
397
Preparatio n of Composit e Material s Calcium Hydroxyapatite/CoUage n by Coprecipitatio n Metho d O.I. SHvka and V.P. Orlovskii
401
Bony Reactio n of Severa l Kinds of Ca-P-Collage n Conjugate d Sponges H. Oonishi, F. Sugihara, K. Minamigawa, Y. Mandai, K. Nagatomi, S. Kushitani, H. Iwaki, N. Kin, E. Tsuji
403
In Vitr o and In Vivo Tests of Newly Developed TCP/CPL A Composite s M. Kikuchi, S-B. Cho, Y. Suetsugu, J. Tanaka, T. Kobayashi, M. Akao, Y. Koyama and K. Takakuda
407
Occlusion of Dentin Tubule s by 45S5 Bioglassfi L.J. Litkowski, G.D. Hack, H.B. Sheaffer and D.C. Greenspan
411
Contents xix
Bioiner t and Biodegradabl e Polymeri c Matri x Composite s Filled With Bioactive Si02-3CaO-P205-Mg O Glasses and Glass-Ceramic s R.L. Reis, A.M. Cunha, M.H. Ferdandez and R.N. Correia
415
The Healin g of Segmenta l Bone Defects, Induce d by Bioresorbabl e Calcium Phosphat e Cement Combine d With rhBMP- 2 ; Using as Past e K. Ohura, C. Hamanishi, S. Tanaka and N. Matsuda
419
DENTA L AND E.N.T . APPLICATION S Implan t Placemen t Enhance d by a New Bioactive Materia l E. Schepers and L. Barbier
425
Behaviour of Bioactive Glass (S53P4) in Huma n Fronta l Sinus Obliteratio n K. Aitasalo, J. Suonpaa, M. Peltola and A. Yli-Urpo
429
All-Cerami c Denta l Bridge s by the Direct Cerami c Machinin g Proces s (DCM) F. Filser, P. Kocher, H. Liithy, P. Scharer and L. Gauckler
433
Grindin g of Zirconi a - TZP in Dentistr y - CAD/CAM-Technolog y for the Manufacturin g of Fixed Dentures R. Luthardt , W. Rieger and R. Musil Zirconi a Implant s With a Plasma-Spraye d SiOi -HA Bioactive Coatin g A. Pedra and P. Sharroc k
437
445
ORTHOPAEDI C APPLICATION S Effect of Time and Temperatur e on the Productio n of Porou s Electrolyti c Hydroxyapatit e Coating s N. Asaoka, S. Best and W. Bonfield
447
Calcium Phosphat e Formatio n on Chemicall y Treate d Vacuum Plasm a Spraye d Titaniu m Coating s S.-W. Ha, K-L. Eckert , H. Gruner and E. Wintermante l
451
Propertie s of Plasm a Spraye d Bioactive Fluorhydroxyapatit e Coating s X. Ranz, C. Rey, N. Antolotti, M.F. Harmand, A. Moroni, L. Orienti, G. Viola, S. Bertini, A. Scrivani
455
XX
Contents
Longer-Ter m Mechanica l and Biological Evaluatio n of Titaniu m Alloy Coate d With Apatit e Layer
459
W.Q. Yan, K. Kawanabe, T. Nakamura and T. Kokubo Electrophoreti c Coating s of Porou s Apatit e Composit e onto Alumin a Ceramic s
463
K. Yamashita, E. Yonehara, J-i. Hamagami and T. Umegaki Osseointegratio n in Experimenta l HA-Coate d Femora l Stems
467
E. De Santis, G. Rinonapoli, C. Doria, A. Manunta and M.C. Sbernardori Effect of Hydroxyapatite-Coatin g on the Bondin g of Bone to Titaniu m Implant s in the Femu r of Ovariectomize d Rat s
471
T. Hara, K. Hayashi, Y. Nakashima and Y. Iwamoto
BI O ACTIV E BON E CEMEN T Optimizatio n of Settin g Time and Mechanica l Strengt h of ^-TCP/MCP M Biocement s
477
P. Van Landuyt, C. Lowe and J. Lemaitre Influenc e of the Particl e Size of the Powder Phas e in the Settin g and Hardenin g Behaviou r of a Calciu m Phosphat e Cemen t
481
M.P. Ginebra, E. Fernandez, F.C.M. Driessens, M.G. Boltong and J.A. Planell Subcutaneou s Tissue Response s and Kinetic s of Cells to Tetracalciu m Phosphat e Cement s
485
M. Yoshikawa, H. Oonishi, Y. Mandai, F. Sugihar and T. Toda Biological Behaviou r of a Bioactive Bone Cemen t Implante d in Rabbi t Tibia e
489
A. Afonso, M. Vasconcelos, R. Branco and J. Cavalheiro Histologica l Stud y of a DCPD-Base d Calciu m Phosphat e Cemen t
493
P. Frayssinet, L. Gineste, P. Conte, J. Fages, N. Rouquet and A. Lerch Bioactive Bone Cemen t Studie d in Canin e Tota l Hip Arthroplasty , 2 Year s FoUow-Up Stud y H. Fujita, T. Nakamura, K. Ido, Y. Matsuda, H. lida, M. Kobayashi, M. Oka and Y. Kitamura
497
Contents xxi
NE W MATERIAL S AND TECHNOLOGIE S
Experimenta l Stud y on Hydroxyapatite/N-Carboxymethy l Chitosa n Filler s
503
R. Martinetti, L. Dolcini, A. Ravaglioli, A. Krajevski and C. Mangano
Injectabl e Chitosamin e Hydroxylapatit e Bone Past e
507
J.J . Railhac, P. Sharrock, D. Galy-Fourcade, C. Zahraoui and N. Sans
Manufactur e of a Hydroxyapatite-Chiti n Composit e
511
A.C.A. Wan, E. Khor and G.W. Hastings
Load-Bearin g and Ductil e Hydroxylapatite/Polyethylen e Composite s for Bone Replacemen t
515
R.L. Reis, A.M. Cunha and M.J. Bevis
In Vitr o Assessment of Hydroxyapatite - and Bioglassfi - Reinforce d Polyethylen e Composite s
519
J. Huang, L. Di Silvio, M. Wang, K.E. Tanner and W. Bonfield
Osteoconductiv e Propertie s of Pur e and Type-A Carbonate d Hydroxyapatite s
523
S.A. Redey, D. Bernache-Assollant, C. Rey, P.J. Marie, M. Nardin and L. Sedel
Coagulatio n Times of Blood in Contac t With Gel-Derive d Silica-Alumin a Composit e Powder s
527
S. Takashima, C. Ohtsuki, S. Hayakawa and A. Osaka
Preparatio n of P"*"-Implante d Y203-Al203-Si02 Glas s for Radiotherap y of Cance r
531
M. Kawashita, F. Miyaji, T. Kokubo, G.H. Takaoka, I. Yamada, I. Suzuki and M. Inoue
New Ferromagneti c Bone Cemen t for Local Hyperthermi a K. Takegami, T. Sano, H. Wakabashi, J. Sonoda, T. Yamazaki, S. Morita, T. Shibuya and A. Uchida
535
xxii
Contents
BIOCERAMIC S PROCESSIN G Adsorptio n of L-Lysin e onto Silica Glass : A Synergisti c Approac h Combinin g Molecula r Modelin g With Experimenta l Analysi s
541
R.A. Latour Jr., J.K. West, L.L. Hench, S.D. Trembley, Y. Tian, G.C. Lickfield and A.P. Wheeler Effects of Divalent Cation s on Calciu m Phosphate s Precipitatio n on a Langmuir-Blodgett e Monolaye r
545
S.B. Cho, Y. Suetsugu, J. Tanaka, R. Azumi and M. Matsumoto Effect of Processin g on the Characteristic s of a 20 Vol.% AI2O3 Platelet Reinforce d Hydroxyapatit e Composit e
549
S. Gautier, E. Champion, D. Bernache-Assollant Mechanica l Evaluatio n of Phosphat e Biodegradabl e Glasse s by Mean s of Indentatio n Method s
553
F J . Gil, R. Terradas, J. Clement, G. Avila, S. Martinez and J.A. Planell Silicon in Connectiv e Tissue: Semi-Empirica l Molecula r Orbita l Model s
557
K.D. Lobel, J.K. West and L.L. Hench The Effect of Hea t Treatmen t on Bone Bondin g Ability of Alkali-Treate d Titaniu m
561
S. Nishiguchi, T. Nakamura, M. Kobayashi, W-Q. Yan, H-M. Kim, F. Miyaji and T. Kokubo Therma l Processin g of Compac t Bovine Bone
565
G. Vargas, M. Mendez, J. Mendez and J. Lopez
EVALUATION S M E T H O D S AND NE W APPLICATION S Characterizatio n of Syntheti c and Biological Calciu m Phosphat e Material s by Micro-Rama n Spectrometr y
571
G. Penel, G. Leroy, G. Cournot and E. Bres Biological Evaluatio n of Glas s Reinforce d Hydroxyapatit e by Flow Cytometr y M.A. Lopes, J.C. Knowles, K.A. Hing, J.D. Santos, F.J. Monteiro and L Olsen
575
Contentsxxiii
Evaluatio n of Macrophag e Respons e to Cerami c Particle s by Flow Cytometry : Analysis of Phagocytosi s and Cytotoxicit y
579
I. Catelas, R. Marchand, L’H. Yahia and O.L. Huk Stud y of Porou s Interconnection s of Biocerami c on Cellula r Rehabitatio n In Vitr o and In Vivo
583
J.X. Lu, B. Flautre, K. Anselme, A. Gallur, M. Descamps, B. Thierry and P. Hardouin Repai r of Osteochondra l Defect Using Artificia l Articula r Cartilag e
587
M. Hasegawa, A. Sudo, Y. Shikinami and A. Uchida Calciu m Phosphat e Ceramic s as Controlle d Release System s for FGF- 2
591
V. Midy , E. Hollande , C. Key and M. Dar d C-SRC Oncogen e mRNA Expressio n in Porou s Hydroxyapatit e Ceramic s K. Mishima, H. Ohgushi, T. Yoshikawa, H. Nakajima, E. Yamada, S. Tabata, Y. Dohi and K. Ichijima Autho r Inde x Keywor d Inde x
595
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CALCIUM PHOSPHATE IN VIVO FORMATION
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PRECIPITATIO N OF CALCIU M PHOSPHAT E ON TITANI A CERAMIC S K.-L. Eckert, S.-W. Ha, S. Ritter, E. Wintermantel Chair of Biocompatible Materials Science and Engineering, Department of Materials, ETH Zurich, Wagistrasse 23, CH-8952 Schlieren, Switzerland
ABSTRAC T Titania ceramics were prepared, pretreated with lOM NaOH solution and immersed in simulated body fluid (SBF) for up to 10 days. Morphological and chemical changes were analysed by SEM and EDX. NaOH-pretreatment lead to surface roughening of the titania grains and to formation of needle-like crystals. EDX analysis showed that Na was present at the surface. After immersion in SBF precipitation of spherical agglomerates occurred together with formation of a layer of needle› like crystals. Na in the surface disappeared and Ca became distinct after 10 days of immersion, suggesting a precipitation of calcium phosphate. Thus the reactions which were observed on titania surfaces after NaOH treatment are similar to those observed on titanium metal after identical treatment. KEYWORD S Titania ceramics, NaOH treatment, calcium phosphate INTRODUCTIO N The formation of biomimetic calcium phosphate layers on titanium surfaces immersed in simulated body fluid (SBF) was shown to occur after chemical treatment with alkaline solutions [1,2]. Due to the similarity of surface oxide layers on titanium metal and previously developed titania ceramics [3], it is assumed that the same procedure could also induce the formation of calcium phosphate layers on titania ceramics. The aim of the present study was to prove latter hypothesis that titania ceramics can be modified by sodium hydroxide (NaOH) treatment in a way that afterwards a calcium phosphate layer is formed if they are immersed in simulated body fluid. MATERIAL S AND METHOD S Ceramic Processing Titania ceramic discs of 10 mm diameter were prepared from titania powders by mixing 20 g of processed titanium dioxide powder (dso" 6.1 |Lim) with 12.0 g titania powder 1171 (Kronos, Germany), 1.0 g graphite powder KS6 (Lonza, Switzerland) and 2.8 g paraffin (MP 64 C, Fluka, Switzerland) into a thermoplastic body. The mixture was heated to 100 C and pressed into round discs of 15 mm diameter. For polymer burnout the samples were placed on an alumina refractory plate and heated to 300 C at a rate of 5 K/h. Final sintering was performed at 1350 C
4
Bioceramics Volume10
with a heating rate of 3 K/min and 25 min holding time. Finishing was done by ultrasonic cleaning for 15 seconds and flushing with deionized water. NaOH Treatmentand Immersionin SBF Alkaline treatment of the titania samples was carried out in lOM NaOH. The specimens were placed into a conical flask filled with 100 ml NaOH. Immersion was performed at 60 C for 2 hours in a laboratory shaker rotating at a speed of 70 rpm. After NaOH treatment the samples were gently flushed with deionized water for 1 minute. Immediately after soaking, the NaOH treated samples were placed into polypropylene vessels containing 25 ml simulated body fluid (SBF) prepared according to [5]. The pH of SBF was 7.4, pH control was performed at the end of every immersion period. The vessels were sealed and immersion in SBF was carried out at 37 C for 1, 4 and 10 days in a laboratory shaker rotating at 70 rpm. After the immersion in SBF the samples were thoroughly rinsed with deionized water and dried in ambient atmosphere at room temperature. Morphological and Chemical Characterization Morphology of the NaOH treated and of the immersed titania surfaces was analysed using scanning electron microscopy (SEM, Hitachi S-2500C). They were compared with control samples, not treated with NaOH, but immersed in SBF. Energy dispersive X-ray (EDX) analysis was performed at an acceleration voltage of 25 kV with an X-ray microanalysis system attached to the SEM (Voyager, Noran Instruments). EDX spectra were acquired with an acquisition time cf 100 seconds. The specimens were coated with platinum in a sputter coater prior to SEM and EDX analysis. RESULT S AND DISCUSSIO N NaOH Treatment Scanning electron microscopic evaluation of the surface of the untreated ceramic showed that the surface has a granular structure (figure 1) with growth steps on the individual grains which occur due to the crystalline nature of the material (figure 2). NaOH treated titania surfaces were similar to untreated samples with the difference that, at higher magnification, surface roughening on the titania crystallites and formation of needle-like structures was observed (figure 4). EDX analysis
Figure 1: Survey of untreated titania ceramic. The surface topography is defined by the granular structure of the material. The total porosity of the material is 25 %.
Figure 2: Untreated titania ceramic. The grains are structured by growth steps.
Precipitationof Calcium Phosphateon Titania Ceramics: K-L. Eckert et al.
5
(figure 7) of the untreated controls showed no other signals than titanium. After NaOH treatment Na was present at the surface. It is assumed that the newly formed, needle-like structures, which were observed with SEM (figure 4) contain sodium. Besides Na and Ti, no additional elements were found at the surface of NaOH treated specimens. Immersion in SBF After one day of immersion, the formation of globular precipitates and a fine layer of needle-like crystals were observed. The appearance of the precipitated layer did not change significantly during the following periods of immersion. After 10 days, the titania surface was completely covered with a layer consisting of needle-like crystals and spherulitic precipitations (figure 5). The thickness of the deposited layer as well as the amount of spherulites (figure 6) was markedly
I igure 3: I ilania ceramic after NaOHtreatment. No change in the topographical characteristics can be noticed.
Figure 4: Titania ceramic after NaOHtreatment. The surface was roughened, growth steps were partly etched away. Newly formed crystals occurred at the surface and in pores.
Figure 5: NaOH-pretreatment and 10 days of immersion in SBF. The surface is completely covered by a precipitated layer.
Figure 6: NaOH-pretreatment and 10 days of immersion in SBF. The precipitated layer consists of needle-like crystals and of spherical agglomerates.
6
Bioceramics Volume10
c
3 O
O
10 Energ y [keV]
Figure 7: EDX spectra of the sample surfaces at various processing stages. On untreated samples only titanium signals could be detected. Pt signals are caused by the platinum sputter coating of the samples. After immersion in NaOH solution (Ti02+NaOH), a Na peak occurs. After 10 days of immersion in SBF, Ca is present, indicating the precipitation of calcium phosphate at the surface. increased compared to day one. However, the original topography of the titania ceramic surface was still maintained. In the EDX spectra, Ca was identified after the first day of immersion and became more distinct with time (figure 7). In contrast, Na vanished, probably due to ion exchange processes [5]. On control samples which were not treated with NaOH, no precipitation was observed. CONCLUSIO N The current work showed that precipitation of calcium phosphates on titania ceramics occurs similarly to those on titanium metal. In conclusion, the investigation has shown the potential of biomimetic calcium phosphate deposition on titania ceramics after pretreatment with lOM NaOH solution at 60 C for 2 hours. REFERENCES 1. Kokubo, T., Thermochimica Acta, 1996, 280/281, 479-490 2. Kim, H.-M., Miyaji, F., Kokubo, T., Nakamura, T., Journal of Biomedical Materials Research, 1996, 32, 409-417 3. Blum, J., Eckert, K.-L., Schroeder, A., Petitmermet, M., Ha, S.-W. and Wintermantel, E. In: Bioceramics Volume 9, Elsevier Science Ltd., Oxford 1996, 89-92 4. Kokubo, T., Hata, K., Nakamura, T., Yamamuro, T. In: Bioceramics Volume 4, Butterworth-Heinemann, Guildford (1991), 113-120 5. Clearfield, A., Lehto, J., Journal of Solid State Chemistry, 73 (1988), 98-106
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
APATIT E FORMATIO N ON POLYMER S BY BIOMIMETI C PROCESS USING SODIU M SILICAT E SOLUTIO N F. Miyaji^, S. Handa\ T. Kokubo^ and T. Nakamura^ ^ Department of Material Chemistry, Faculty of Engineering, Kyoto University, Yoshida-honmachi, Sakyo-ku, Kyoto 606-01, Japan ^Department of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, ShogoinKawaharacho, Sakyo-ku, Kyoto 606-01, Japan
ABSTRAC T A variation of biomimetic process which aims at a bonehke apatite coating on organic polymers with complex shapes was attempted by using sodium silicate as a catalyst for the apatite nucleation, and a simulated body fluid as a medium for the apatite growth. An apatite-forming abihty was the highest when the Si concentration and Si02/Na2 0 ratio of the sodium siUcate solution were above 2.0 M and 1.0-1.5, respectively. Particular sihcate ohgomers were assumed to be most responsible for the apatite nucleation. The apatite layer was formed not only on the flat PET surfaces but also on curved surfaces of fine PET fibers constituting a fabric, where the apatite layer was interconnected each other. This method is expected to enable the bonelike apatite coating on various kinds of materials with complex shapes. KEYWORDS : Apatite, Biomimetic process, Sodium silicate, Simulated body fluid INTRODUCTIO N A biomimetic process has been developed for coating a dense and uniform bonelike apatite layer on organic polymers as follows [1,2]. First, in order to form apatite nuclei on the substrates of organic polymers, the polymer substrates are placed on CaO-Si02- or Na2 0-Si02based glass particles soaked in a simulated body fluid (SBF) with ion concentrations nearly equal to those of human blood plasma [3]. Next, in order to make the apatite nuclei grow, the polymer substrates are soaked in 1.5SBF with ion concentrations 1.5 times the SBF. The disadvantage of the above biomimetic process lies in the diflficulty of apatite coating on the materials with complex shapes, since in the first treatment the apatite nuclei are formed only on the material surface which is faced to the glass grains. In the present study, the apatite formation on organic polymers was attempted by using sodium silicate solution as a nucleating agent for the apatite formation instead of the glass particles, and 1.5 SBF as a medium for the apatite growth. MATERIAL S AND METHOD S Preparatio n of sodiu m silicat e solutio n Reagent grade sodium metasilicate (Na2 Si03) was dissolved into distilled water to prepare solutions with 0.5, 1.0, 2.0 and 3.0 M Si-concentration. As another series of sodium silicate
8
Bioceramies Volume10
solutions, the solutions with SiO./Na.O ratio of 0.5, 0.67, 0.8, 1.0, 1.5 and 2.0 were prepared by adding reagent grade NaOH or SiO ^ xH^ O into sodium metasihcate solution, where Si concentration was fixed at 3.0 M. Preparatio n of l.SSBF The 1.5SBF with ion concentrations (Na^ 213.0, K^ 7.5, Mg^^ 2.3, Ca^^ 3.8, CI’ 223.2, HCO3’ 6.3, HPO/ 1.5, s o / 0.8 mM) 1.5 times the SBF was prepared by dissolving reagent grade NaCl, NaHC03, KCl, K^HPO.^H^O, MgCl^^H^O, CaCl^ and Na^SO^ into distilled water. The pH of the solution was adjusted at 7.25 with NH^QCH^OH) and 1 M-HCl at 36.5T. Apatit e coatin g on polyme r Rectangular substrates (10 x 10 x 1 mm ) of poly (ethylene terephthalate) (PET) were abraded #400 and washed with ethanol. And then the substrates were subjected to a glow discharge treatment in O2 gas for 30 s [4] for producing polar groups on the polymer surfaces, which might contribute to the strong attachment between silicate ions and the substrate. After the treatment, the substrates were soaked in sodium sihcate solutions with various concentrations and Si02/Na20 ratios for 6 h at 36.5T. After removing from the solution, the substrates were dried at room temperature, rinsed with distilled water and then soaked in 20 ml of 1.5 SBF for various periods. A fabric (10 x 15 mm^ in area) woven with ultrafme PET fiber (2 \xm^)(Toray Co. Ltd., Otsu, Japan) was also used as a substrate. Surfac e analysi s After the soaking in 1.5 SBF, surface structural and morphological variations of the specimens were characterized by a thin-film X-ray diffractometer (TF-XRD; thin-film attachment CN2651A1, Rigaku-Denki Co., Tokyo, Japan), a Fourier transformed infrared (FT-IR) reflection spectrometer (System 2000 FT-IR, Perkin-Elmer Ltd., Buckinghamshire, England) and a scanning electron microscope (SEM; S-2500CX, Hitachi Co., Tokyo, Japan). RESULT S AND DISCUSSIO N Effect of concentratio n of sodiu m silicat e solutio n Figure 1(a) shows the TF-XRD patterns of the surfaces of PET substrates soaked in 1.5 SBF for 6 d after soaking in sodium metasihcate solutions with various concentrations for 6 h. The peaks ascribed to apatite were observed for the 2.0 and 3.0 M-treated specimens. Figure 2 shows the SEM photographs of the surfaces of PET substrates soaked in 1.5 SBF for 6 d after soaking in sodium metasihcate solutions with various concentrations for 6 h. In the case of 0.5 M, apatite particles were deposited only on a small part of the surface of the substrate. The number of apatite particles increased for the treatment with 1.0 M solution. Moreover, whole surfaces of the substrates were covered with apatite layer for the treatment with 2.0 and 3.0 M solutions. These results indicate that the apatite-forming tendency becomes higher with increasing concentration of sodium sihcate solution. This is explained by assuming that the number of sihcate ions attached to the substrates increased with increasing concentration of sodium silicate solution, forming more apatite nuclei. It should be, however, noted that the degree of the apatite formation is almost the same between 2.0 and 3.0 M. This suggests that the number of sihcate ions attached to the substrates are saturated at 2.0 M concentration. Effect of compositio n of sodiu m silicat e solutio n Figure 1(b) shows the TF-XRD patterns of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium silicate solutions with various Si02/Na2 0 ratios for 6 h. The peaks ascribed to apatite were observed for 1.0 and 1.5 of Si02/Na20 ratios. Figure 3 shows the SEM photographs of the surfaces of PET substrates soaked in 1.5 SBF for 6 d after
Apatite Formation on Polymers by BiomimeticProcessing: F. Miyaji F. et al. 9 O Apatite
O Apatite
SiOg / Na20
20 30 40 50 60 i^ 20 30 40 50 60 26 (CuKa) / degree 29 (CuKa) / degree Figure 1 TF-XRD patterns of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium silicate solutions with (a) various concentrations and (b) various Si02/Na20 ratios for 6 h. 10
fi 6 ' D3
go fi CD
mMm
0 fi O O O u D Rt
’0fiOQQuD
Figure 2 SEM photographs of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium metasilicate solutions with various concentrations for 6 h.
Hofi
o'
iofi tlfiCQQuD
’’Wm
Figure 3 SEM photograplis of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium sihcate solution with various Si02/Na2 0 ratios for 6 h.
10 Bioceramies Volume10
soaking in sodium silicate solution with various Si02/Na2 0 ratios for 6 h. No apatite was observed for 0.5 of ^\0^f^?i^Oratio. The whole surfaces of the substrates were covered with apatite for 1.0 and 1.5 of Si02/Na20 ratios. Apatite particles were deposited on the substrates in part for 2.0 of Si02/Na2 0 ratio. These indicate that the sodium silicate solutions with 1.0 and 1.5 of Si02/Na2 0 ratios are most adequate for apatite formation. It is well known that the structure of sihcate ions changes with the composition of sodium silicate solutions [5]: silicate ions are primarily present as monomer for SiO2/Na2O<1.0, dimer, linear trimer and/or cychc tetramer for 1.0<SiO2/Na2O<2.0, and polymer larger than oligomer for SiO2/Na2O>2.0. Taking into account that the substrates which were treated with sodium silicate solutions of 1.0-1.5 of Si02/Na20 ratio formed apatite most effectively, silicate oligomers such as dimer, linear trimer and/or cyclic tetramer might most contribute to the apatite nucleation. Apatite coating on PET fabric It was observed by SEM that the continuous apatite layer is formed on each fine PET fiber. The apatite layer was interconnected each other and it was not peeled off even by the bending of the fabric. Characteristics of apatite In the present work, even in the case where apatite-forming ability is the highest, the surface of the apatite layer was not smooth. Tliis indicates that the number of apatite nuclei formed on the substrates are not enough for forming dense and uniform apatite layer, which may be due to lower apatite-forming ability of adsorbed silicate ions compared with that of silicate ions released from CaO, Si02-based glass grains. The adhesiveness of the apatite to the substrate was also lower in the former than in the latter. This is explained by assuming that the bonding of sihcate ions with the substrate surface is not stable. However, the most distinct advantage of this method from other ones is a possible coating of the bonelike apatite on various materials with complex shapes. The bonehke apatite layer is formed not only on the flat surfaces but also on curved surfaces of fine PET fibers constituting a fabric. This structure is quite similar to that of natural bone in part. If this composite can be fabricated into three dimensional structure analogous to that of natural bone, the resultant composite could exhibit analogous mechanical properties to those of natural bone, in its elastic modulus as well as in its strength and fracture toughness. ACKNOWLEDGMENT This work was supported by a Grant-in-Aid for Scientific Research, The Ministry of Education, Science, Sports and Culture, Japan. REFERENCES 1. Tanahashi, M., Yao, T., Kokubo, T., Minoda, M, Miyamoto, T., Nakamura, T. and Yamamuro, T., 7. Am. Ceram, Soc. 1994, 77, 2805-2808. 2. Tanahashi, M., Yao, T, Kokubo, T., Minoda, M., Miyamoto, T., Nakamura, T. and Yamamuro T., 7. Ceram, Soc. Japan 1994, 102, 822-829. 3. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T. and Yamamuro, T., 7 Biomed.Mater.Res. 1990,24,721-734. 4. Tanahashi, M., Yao, T, Kokubo, T, Minoda, M., Miyamoto, M, Nakamura, T. and Yamamuro, T., 7 Biomed.Mater.Res.1995, 29, 349-357. 5. Her, R.K., In: Ue Chemistryof Silica,John Wiley & Sons, New York 1979, 130-145.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BONELIK E APATIT E FORMATIO N ON THE SURFACE OF CHEMICALL Y TREATE D TANTALUM SUBSTRATES : EFFEC T OF HEAT TREATMEN T T. Miyazaki\ H. M. Kiin\ F. Miyaji\ T. Kokubo^ and T. Nakamura^ ^Department of Material Chemistry, Faculty of Engineering, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan, ^Department of Orthopedic Surgery, Faculty of Medicine, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan
ABSTRAC T It was previously shown that the induction period for apatite formation on Ta metal in a simulated body fluid (SBF) is decreased from 4 to 1 weeks, when it was treated with a NaOH solution to form a sodium tantalate gel layer on it. The sodium tantalate gel layer as-formed on the surface of the NaOH-treated Ta metal is, however, mechanically unstable and easily peeled off from the substrate. In the present study, the effects of the heat treatment on the structure and apatiteforming ability of the surface layer of the NaOH-treated Ta metal were investigated. The sodium tantalate gel layer which was formed on Ta via NaOH-treatment was dehydrated and stabilized by the heat treatment above 300 C. The NaOH-treated Ta metal heat-treated at 300 C showed almost equal induction period for the apatite formation to that of the substrate without heat treatment. KEYWORDS : Tantalum, Heat treatment, Simulated body fluid (SBF), Bioactivity, Apatite, Sodium tantalate INTRODUCTIO N Metallic materials such as stainless steel and titanium alloys have a high strengths and hence have been widely used for orthopedic implants under highly loaded condition. They, however, do not bond directly to living bone. On the other hand, some ceramic materials are known to bond directly to living bone via formation of a bonelike apatite layer on their surfaces [1]. If such a bonelike apatite layer is formed on the surfaces of the metals in vivo,they can directly bond to living bone. Since Ta metal shows high malleability and ductility and hence can be fabricated in complex shapes [2], bioactive Ta metal is expected to be useful as orthopedic implants in complicated shapes. It was previously shown that a dense and uniform bonelike apatite layer was formed on Ta metal within 1 week when it was treated with NaOH aqueous solution and soaked in simulated bodyfluid(SBF) with ion concentrations nearly equal to human blood plasma [3]. This induction period for the apatite formation is much shorter than that of the untreated Ta metal, 4 weeks. In the process of the NaOH treatment, sodium tantalate gel layer is formed on the surface of Ta metal. The surface gel layer is, however, unstable and easily peeled off from the substrate. Heat treatment of the NaOH-treated Ta metal is expected to dehydrate and densify the surface gel layer as in the case of the NaOH-treated Ti metal [4]. In the present study, the effects of the heat treatment on the surface structure, stability of the surface layer, and apatiteforming ability in SBF were investigated for the NaOH-treated Ta metal.
12 Bioceramics Volume10
MATERIAL S AND METHOD S Tantalum plates of 10 X10 X1 mm^ were ground with #400 diamond paste, treated with 5 ml of 0.5 A/-NaOH aqueous solutions at 60 C for 24 h, washed with distilled water and dried at 40 C for 24 h. After the above treatments they were heated up to various temperatures, from 200 to 500 C, at a rate of 5 C/min in an electric ftinace, inunediately took out of the fiimace, and allowed to cool in room temperature. In order to examine the stability of the surface layer, the NaOH-treated Ta substrates with and without heat treatment were attached with Scotch tapefi and then it was detached. The NaOH- and heat-treated Ta substrates were soaked in 30 ml of simulated body fluid (SBF) with ion concentrations nearly equal to those of human blood plasma [5] and pH7.40 at 36.5 C for various periods until 4 w. The surface structural changes of the Ta due to the NaOH treatments and soaking in SBF were characterized by thin-film X-ray diffraction (Thin-film attachment CN2651A1, Rigaku, Co., Tokyo, Japan) and scanning electron microscopy (SEM; S-2500CX, Ifitachi Co., Tokyo, Japan). Changes in element concentrations and pH of SBF due to soaking of the specimens were measured by inductively coupled plasma (ICP) atomic emission spectroscopy (Model SPS-1500VR, Seiko Instrument Co., Japan) and a pH meter (Model D-14, Horiba Co., Japan). RESULT S AND DISCUSSIO N Figure 1 shows the SEM photographs of the detached surface of Ta substrates with and without heat treatment at 300 C after the NaOH-treatment. The gel layer on the surface of asNaOH-treated Ta substrate without heat treatment was mechanically so weak that it was easily peeled off the substrate. On the other hand, the surface layer of the substrate which was heattreated at 300 C was not peeled off from and the glue of the tape remained on the substrate. This indicates that the sodium tantalate gel layer on Ta metal was stabilized by the heat treatment as low as at 300 C. Figure 2(a) shows the thin-film X-ray diffraction patterns of the surfaces of Ta substrates with NaOH-treatment and heat treatment. The sodium tantalate gel layer remained as an amorphous phase up to 300 C, taking into account the broad halo patterns at about 30 in 29, and it converted into crystalline sodium tantalate at 500 C. In the former process below 300 C, the sodium tantalate gel layer might be dehydrated to form an amorphous sodium tantalate layer. Figure 2(b) shows the thin-film X-ray diffraction patterns of the surfaces of Ta substrates which were subjected to NaOH and heat treatments, and subsequent soaking in SBF for Iweek. The NaOH-treated Ta metal heat-treated at 300 C showed almost equal induction period for the As NaOH-treated
With heat treatment at 300 C
'8(9 ' (§QOQ'
m
Figure 1 SEM photographs of the detached surfaces of Ta substrates with (right) and without (left) heat treatment at 300 C after 0.5M-NaOH treatment.
Bonelike Apatite Formation on theSurface of TantalumSubstrates:T. Miyazaki et al. 13
10
20
30
40
50
60
10
2ft(CuKa) /deg.
20
30
40
50
60
2B(CuKa) /deg.
Figur e 2 Thin-fil m X-ra y diffractio n pattern s of the surface s of Ta substrate s which were subjecte d to 0.5A/-NaOH treatmen t and heat treatmen t (a) and subsequen t soakin g in SBF for 1 w (b). 155
None -*-
8 i4oJ r c o O 135-
Na 50
100
150
200
0
Soaking time /h
1 -1
E
^ 0.9 C
o
(0 C 0.8 -
(1)
50
100
150
Soaking time /h
200
o c o Oo.7-
50
100
150
200
Soaking time /h
L
^ V^500 C \ ^ \ "^ \ 300 C V. "^ None ^ 1
50
,
100
1
150
200
Soaking time /h
Figur e 3 Change s in element concentration s and pH with soakin g of NaOH-treate d Ta substrate s with and withou t heat treatment .
14 Bioceramies Volume10
apatite formation to that of the substrate without heat treatment, whereas that heat-treated at 500 C showed the induction period as long as 2 weeks. At any rate, these induction periods are much shorter than that of original Ta substrate without NaOH-treatment, 4 weeks. Figure 3 shows the changes in element concentrations and pH with soaking of NaOH-treated Ta substrate with and without heat treatments. The pH value and the concentration of Na steeply increased and those of Ca and P decreased with soaking of Ta substrates in SBF for all the treatments. The increase in pH at the initial soaking period is significantly smaller in heat treatment at 500 C than in heat treatment at 300 C and without heat treatment. The difference in the apatite-forming ability with the heat treatment conditions are explained as follows. The tantalum oxide hydrogel laytir is formed on the NaOH- and heat-treated Ta metals by ion exchange of Na^ ion in the amorphous or crystalline sodium tantalate layer with HsO^ ion in SBF when they were soaked in SBF. It is assumed that thus formed tantalum oxide hydrogel induces apatite nucleation and the released Na^ ion increases the degree of supersaturation with respect to apatite in SBF by increasing pH to accelerate the apatite nucleation. The apatite nuclei grow into a continuous layer by incorporating the Ca^^ and P04^" ions from the surroundingfluidwhich is already supersaturated with respect to the apatite [3]. In this case, the rate of apatite formation should be governed by the rate of the Na^ ion release fron the sodium tantalate layer, since it determines the rates of the tantalum oxide hydrogel formation and the increase in the degree of supersaturation with respect to apatite in SBF. The rate of the Na^ ion release decreased with increasing heat treatment temperature due to the dehydration and densification of the sodium tantalate gel layer. Therefore, the rates of the tantalum oxide hydrogel formation and the increase in the degree of supersaturation of SBF decreased and hence the induction period for the apatite formation increased in the order without heat treatments treatment at 300 C< treatment at 500 C. The increase of induction period for the apatite formation was especially remarkable for the treatment at 500 C (2 weeks) because the crystalline sodium tantalate gave only a small amount of Na^ ion release, resulting in the small pH increase (SeeFig.3) In conclusion, the Ta metal which was subjected to NaOH and 300 C-heat treatment gave the stable surface layer without retarded apatite formation. CONCLUSIO N The surface of the NaOH-treated Ta was stabilized by the heat treatment at 300 C. The induction period for the apatite-formation of it was still much shorter than that without NaOH treatment. Thus NaOH- and heat-treated Ta metal is believed to form apatite on its surface also in vivo in a short period and bond directly to living bone. REFERENCE S 1. Kokubo, T., J.Ceram. Soc. Japan,1991, 99, 965-973. 2. Stackpool, G. J., Kay, A. B., Morton, P., Harvey, E. J., Tanzer, M., Bobyn, J. D., In: Proceedingsof CombinedOrthopaedicResearchSocietiesMeeting,San Diego, California 1995, 45. 3. Miyazaki, T., Kim, H. M., Miyaji, F., Kokubo, T., and Nakamura, T., In: Bioceramics Volume 9, Elsevier, Oxford 1996, 317-320 4. Kim, H. M., Miyaji, F., Kokubo, T., and Nakamura, T., J. Mater.Sci.: Mater.Med, in press 5. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T. and Yamamuro, T., J. BiomedMater.Res., 1990, 24, 721-734.
Bioceramics, Volume10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposiumon Ceramics in Medicine, Paris, France, October1997) '1997 Elsevier Science Ltd
INFLUENC E OF MORPHOLOG Y ON
HYDROXYAPATIT E
PARTICL E
SIZE
AND
HAPEX^ M
M.Wang, R Joseph* and W.Bonfield IRC in Biomedical Materials, Queen Mary and Westfield College, Mile End Road, London El 4NS,U.K. Sree Chitra Tirunal Institute for Medical Sciences and Technology, Trivandrum-695 012, Kerala, India
ABSTRAC T Synthetic hydroxyapatite particles of two median sizes and different morphologies have been used to manufacture hydroxyapatite reinforced high density polyethylene composites (HAPEX^^) for medical applications. The effects of hydroxyapatite particle size on properties of the resultant composites were investigated using various techniques. It was found that HAPEX^^ with smaller hydroxyapatite particles possessed higher torsional modulus. Young’s modulus and tensile strength, but lower strain to failure. Examination of fracture surfaces revealed that a mechanical bond existed between the reinforcement and the matrix. particlesize,morphology,polyethylene, composite,property KEYWORDS : hydroxyapatite, INTRODUCTIO N Hydroxyapatite (HA) reinforced high density polyethylene (HOPE) composite, now designated as HAPEX’^^, has been developed since the early 1980’s as an analogue for bone replacement [1]. With 40vol% of HA, HAPEX^^ possesses a stiffness approaching that of cortical bone, a superior toughness and considerable bioactivity [2]. Implants made of HAPEX^^ encouraged bone apposition rather than fibrous encapsulation which is encountered with other implant materials [3]. The modulus and strength of HAPEX^^ are adequate for minor load bearing applications such as orbital floor reconstruction [4]. Recent progress in hydrostatic extrusion of HAPEX^^ indicated that composites with high stiffness and strength (within the bounds for cortical bone) can be manufactured for major load bearing skeletal implants [5]. Previous work on ceramic particle (or glass bead) reinforced polymers suggested that the particle size had effects on mechanical properties of composites [6,7]. Likewise, HA particle size and morphology are expected to influence the properties of HAPEX"^^. In this paper, results obtained from HAPEX’^^ composites with HA of different particle characteristics are reported. Such information is important for other investigations where techniques such as chemical coupling and hydrostatic extrusion are used for the enhancement of properties of HAPEX’^. MATERIAL S AND METHOD S Two grades of synthetic hydroxyapatite powder (P88 HA and P8IB HA, Plasma Biotal Ltd, UK) were incorporated, respectively, into a high density polyethylene (Rigidex HM4560XP, BP Chemicals Ltd, UK) to produce HAPEX^^ through an established processing route [8]. Prior to 15
16
Bioceramics Volume10
composit e processing , the particl e size of HA was measure d with a Malver n MasterSize r and the particl e morpholog y examine d by a JEOL6300 F scannin g electro n microscop e (SEM) . The BET surfac e are a of HA paiticle s was also analysed . Unfilled HDPE and HAPEX^ M with up to 45vol% of HA were produce d 1.75mm and 4mm thick composit e plate s were mad e by compressio n moulding . The molecula r mass of the HDPE matri x was measure d using high temperatur e gel permeatio n chromatograph y (GPC ) and the distributio n of HA particle s in HDPE was investigate d unde r SEM. Tensil e specimen s (ISO 527) were mad e from 1.75mm thick composit e plates . All tensil e tests were conducte d on an Instro n 6025 testin g machin e at a crosshea d speed of 0.5mm/min . Selected tensile fractur e surface s of the HAPEX^ ^ composite s were examine d unde r SEM. For torsiona l tests , rectangula r bar s (25mmx2mmx2mm ) were cut from 4mm thick composit e plates . Torsiona l modulu s of HAPEX^ ^ was determine d using a specially built testin g rig [9]. For dynami c mechanica l analysi s (DMA), specimen s of dimension s 23mmx3mm x 1.75mm were used. All tests were performe d on a Perkin-Elme r DMA 7 system in the temperatur e time scan mode, using the thre e point bendin g configuratio n at the frequenc y of IHz . The stati c stres s was 120% of the dynami c stres s require d in orde r to give a strai n value of 0.02% . A temperatur e rang e of -100 to 110*C was used for specimen s with 0, 20 and 40vol% of HA. A heatin g rat e of 4"C/minut e was maintaine d throughou t the analysis . RESULT S The media n particl e sizes of as-receive d hydroxyapatit e powder s were 4.14jim and 7.32|Lim for P88 HA and P81B HA, respectively , but the particl e size distributio n peake d at 3.8|Lim (P88 HA) or 6.5|a,m (P81B HA). Both grade s of HA had significan t amount s of sub-micro n size particles . Their specific surfac e area s were marginall y different , being 8.27m2/g for P88 HA and 7.61mVg for P81B HA. SEM examinatio n at high magnification s reveale d tha t particle s of both grade s of HA were compact s of crystallites , but these crystallite s appeare d mor e orderl y and were large r in P81B HA tha n in P88 HA. After compounding , HA particle s in both composite s were well dispersed , with a homogeneou s distributio n in the polymer matrix . Subsequen t composit e processin g did not alter these characteristics . Result s obtaine d by high temperatur e GPC showed tha t a slight reductio n in weight averag e molecula r mass of HDPE occurre d durin g compounding . With an increas e in HA volume fraction , the averag e molecula r mass of HDPE was decreased . Tabl e 1 Mechanica l propertie s of HAPEX"^ ^ composite s HA Volume
(%) 0 10 20 30 40 45
Young’ s Modulu s (GPa ) (P88) (P81B)
Tensile Strengt h (MPa ) (P81B) (P88)
0.65+0.02 0.98–0.02 1.60–0.02 2.73–0.10 4.29–0.I7 5.54–0.62
17.89–0.29 18.32+0.34 17.30+0.27 17.00+0.29 17.77+0.09 17.65+0.17 19.55+0.20 19.99–0.15 20.67+1.56 20.85+0.30 18.98+2.11 20.07+1.33
0.72–0.03 0.98+0.07 1.55+0.04 2.46+0.21 3.74+0.14 5.39–0.81
’ rorsiona l Modulu s (GPa ) (P88) (P81B) 0.28+0.09 0.39+0.16 0.48–0.06 0.71+0.16 1.18+0.07 1.46+0.26
0.28–0.09 0.27–0.04 0.39–0.05 0.62–0.07 1.03–0.16 1.23–0.21
Result s obtaine d from tensile and torsiona l testin g ar e tabulate d in Tabl e 1. Young’ s and torsiona l modul i as well as tensile strengt h of HAPEX^ ^ increase d significantl y with an increas e in HA volume fraction , while fractur e strai n decreased . Composite s with high HA volume s (up to 40% ) possessed considerabl e ductiht y (fractur e strai n = 4.5% with 40vol% of P81B HA). With 45vol% of HA, the composite s becam e brittl e and exhibite d a reduce d tensile strength .
Influenceof Hydroxyapatite Particle Size and Morphology on Hapex : M. Wang et al.
17
SEM examinatio n of tensile fractur e surface s of the composite s suggested tha t ther e was no chemica l bond between HA and HDPE . The polymer flowed aroun d the HA particle s and the particle s debonde d from the matri x durin g tensile testing . No residua l polyethylen e was found on HA particle s after tensile fracture. The DMA result s presente d in this pape r ar e obtaine d from HAPEX’^ ^ with P81B HA, which ar e used to show trend s rathe r tha n for comparison . For polymeri c materials , the following equation s exist: E = a/e = E’ + iE"
(1)
tan a = E"/E ’
(2)
wher e E is dynami c modulus , E’ is storag e modulu s and E" loss modulus , tan 3 quantifie s the interna l frictio n or damping . The storag e modulu s (E’ ) of HAPEX"^ ^ is plotte d agains t temperatur e in Figur e 1. As expected , it increase d with an increas e in HA conten t and decrease d with an increas e in temperature . Unlik e storag e modulu s which decrease d monotonicall y with an increas e in temperature , the loss modulu s of the composite s initiall y decrease d with a temperatur e rise from -100*C to -25*C, it then increase d to p e ^ at aroun d 40*C, with a subsequen t decreas e toward s higher temperatures . The tan 3 values were calculate d for HAPEX^ ^ between -100*C and lOO’C . At temperature s of interest , i.e. 20*C, 37 C and 60’C , obtaine d tan d values ar e plotte d agains t HA volum e fractio n in Figur e 2. The additio n of HA reduce d the dampin g of the polymer , with the degre e of reductio n dependin g on HA volume fraction . 10.00 n
S
o
8.00 H
Vp = 0 Vp = 20% Vp = 40%
0.30 n eo 0.2 0 4
0.10
T r -150 -100 -50 0 50 100 150 Temperatur e ( C) Fig.l Variatio n of storag e modulu s with temperatur e for HAPEX^ M with P81B HA
0.00- 1
0.0
r
0.1
0.2 0.3 0.4 HA Volume Fractio n Fig.2 Variatio n of tan d with HA volume fractio n for HAPEX with P81B HA
T
0.1 0.2 0.3 HA Volume Fractio n
Fig.3 Normalise d Young’ s modulu s as a functio n of HA volumefraction for HAPEX^ ^
r
0.2 0.3 0.4 HA Volume Fractio n Fig.4 Normalise d tensile strengt h as a functio n of HA volumefraction for HAPEX^ ^
18 BioceramiesVolume 10
DISCUSSIO N The two grades of particulate hydroxyapatite, P88 HA and P81B HA, do not differ significantly from each other except for their size and morphology. Difference in the properties of the composites can only be attributed to the difference in HA particle size and morphology. Examination of polished surfaces indicated that HAPEX^^ containing either grade of HA had a homogeneous distribution of HA particles in the HDPE matrix. As was discussed in a previous paper [8], the design of the compounding extruder is such that a high level of shear is produced during composite compounding, which results in the break-down of HA particle agglomerates. When composite properties are normalised to those of unfilled HDPE, Young’s and torsional moduli of the composites exhibit monotonic increases with the increase in HA content (Figure 3). The tensile strength experiences an initial decrease but increases to peak at 40vol% of HA. Further addition of HA causes reductions in composite strength (Figure 4). The Young’s and torsional moduli are consistenUy lower for HAPEX^^ with P81B HA than those of HAPEX^M with P88 HA. The difference in median particle size is 3.18|xm between P81B HA and P88 HA (but P81B HA nearly doubles the size of P88 HA), which has caused reductions of O.lSGPa and 0.55GPa in torsional modulus and Young’s modulus respectively. These decreases are relatively small but significant. If HA particles with much larger size differences were used, greater differences in properties would be expected, as was the case for other composites [7]. Since HDPE is a viscoelastic material, it has both the capability to store energy (measured as storage modulus) and the ability to dissipate energy (measured as loss modulus). The dissipation of energy manifests itself as internal friction or damping. In a polymeric composite, the energy dissipation may also comefromthe filler-matrix interface wherefrictionbetween the two phases can occur. It appears that the addition of HA has actually limited the mobility of the amorphous phase in the polymer and hence reduced the damping of the composites. An increase in HA content results in the decrease in polymer volume and thus the overall damping of the composite. CONCLUSION S An increase in hydroxyapatite volume fraction leads to increases of both strength and modulus of the HAPEX’^M composites, with a simultaneous reduction in strain to failure. HAPEX’^^ with smaller hydroxyapatite particles possesses higher torsional modulus. Young’s modulus and tensile strength, but lower strain to failure. It is demonstrated that DMA is useful for studying the viscoelastic behaviour of HAPEX’^ composites. ACKNOWLEDGEMENT S The authors would like to thank their colleagues for assistance and helpful discussions during the course of this work. Support from the UK EPSRC is gratefully acknowledged. R.Joseph is indebted to the British Council for sponsorship.
REFERENCES
1. Bonfield, W., Grynpas, M.D., TuUy, A.E., Bowman, J., Abram, J., Biomaterials(1981), 2, 185-186 2. Bonfield, W., Annals ofNew York AcademyofSciences(1988), 523, 173-177 3. Luklinska, Z.B., Bonfield, W., Bone-Bonding Biomaterials, Reed Healthcare Communications, 1992, 74-77 4. Downes, R.N., Vardy, S., Tanner, K.E., Bonfield, W., Bioceramics4 (1991), 239-245 5. Ladizesky, N.H., Wang, M., Miettinen, E.M., Tanner, K.E., Ward, I.M., Bonfield, W., Proceedingsof theFifthWorldBiomaterialsCongress,Toronto, Canada, 1996,442 6. Landon, G., Lewis, G., Boden, G.F., JournalofMaterialsScience(1977), 12, 1605-1613 7. Xavier, S.F., Schultz, J., Friedrich, K., JournalofMaterialsScience(1990), 25, lAW-lAlQ 8. Wang, M., Porter, D., Bonfield, W., British Ceramic Transactions(1994), 93,91-95 9. Somerton, M., Braden, M., Ward, I.M., Woods, D.W., Biomaterials(1991), 12, 13-16
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EFFEC T OF FLUORID E SUBSTITUTIO N ON THE BIOCOMPATIBILIT Y OF HYDROXYAPATIT E King K.A.,1 Di-SUvio L.,2 Gibson I. R.,1 Ohtsuki C.,^ Jha L. J.,1 Best S.M.,1 Bonfield W. 1 IRC in Biomedical Materials, Queen Mary and Westfield College, Mile End Road, London El 4NS, UK. 2 IRC in Biomedical Materials, Institute of Orthopaedics, Brockley Hill, Stanmore, HA7 4LP,UK. 3 Biomaterials Laboratory, Faculty of Engineering, Okayama University, Tsushima, Okayama 700, Japan ABSTRAC T Hydroxyapatite (HA) and 0.2 and 0.3 wt% fluoride-substituted apatite (F-HA) were studied using in vitrotesting by both inmiersion in simulated body-fluid and direct cell culture on ceramic substrates. X-ray diffraction and scanning electron microscopy revealed the formation of a bone› like ^atite layer on F-HA discs after only 3 days inunersion in simulated body-fluid, compared to 14 days on HA discs. The increased activity of the substituted ^atite was confirmed by in vitro assessment using a primary human osteoblast-like (HOB) cell line. HOB cells cultured on the substituted apatites demonstrated elevated levels of tritiated thymidine incorporation compared to pure HA. However, alkaline phosphatase expression after 7 days in vitrowas found to be independent of substrate composition, indicating that fluoride substitution did not affect the phenotypic expression of the HOB cells. Scanning electron microscopy of the cultures after 1 day in vitrodemonstrated formation of a bone-like apatite layer on the surface of the 0.3wt% F-HA, and a protein-like layer was identified on both 0.2 and 0.3 wt% F-HA discs after 3 days in vitro. INTRODUCTIO N De Jong Was the first to observe the similarities between the X-ray diffraction patterns of bone powder and hydroxyapatite in 1926 and until recently, bone mineral was widelyregardedas being analogous to hydroxyapatite. However, the apatite lattice accommodates many ionic substitutions^’^ and it is therefore not surprising that bone mineral, which is produced in solution in the presence of many complex and varied physiological fluids, is composed of a highly substituted non-stoichiometric hydroxyapatite. The principle substituent is carbonate which replaces phosphate groups^, but the calcium ions also undergo substitution by ions such as sodium, magnesium, and strontium^*^ and the hydroxyl groups are reported to be readily exchanged with fluoride and chloride^. It therefore follows that the substitution of controlled levels of the appropriate ions into the hydroxyapatite lattice could be advantageous as a result of closer mimicry to bone mineral. In this study an assessment of the effect of a single lattice substitution (fluoride for hydroxyl) has been made. Ruoride levels in natural bone mineral have been reported as between 0.02 and 0.08 wt % ^’9. 19
20
Bioceramies Volume10 14 Days
JL
J^
7 Days
14 Days 7 Days
*Ac^
a
AMUWA .
3 Days
-.gu^AJkLk. LI
ODays
1 I I I I I I I I I I I 1*1^*1 M f’\ i f I 1 ’ I* rt
20
25
30
35
40
45
50
20
25
30
35
40
45
rr^
50
(5) Diffraction Angle (26) (a) Diffraction Angle (26) Figure 1 XRD patterns for SBF immersed (a) HA and (b) 0.3 wt% F-HA disks. MATERIAL S AND METHOD S Hydroxy^atite was prepared by a precipitation reaction between stoichiometric proportions of calcium nitrate and diammonuium orthophosphate solutions^^. The substitution of 0.2 and 0.3 wt % fluoride was achieved by the addition of varying concentrations of ammonium fluoride during precipitation. The filter cake was dried and milled to pass through a <75 pm sieve. The sieved powder was uniaxially pressed into discs at 40 MPa and sintered at 1200 C for 2 hours to a final diameter of either 12 or 25 mm. The flat surfaces of the larger discs were polished to a 1 pm surface finish and placed in K-9 simulated body fluid (SBF) solution^ ^ for 3, 7 and 14 days at 36.5 C. After immersion, the surfaces of the discs were analysed for any spontaneous apatite precipitate using X-ray diffraction (XRD) and scanning electron microscopy (SEM). The as-sintered, flat surfaces of the smaller discs were seeded with human osteoblast-like (HOB) cells and incubated in Dulbecco’s Modified Eagles Medium, supplemented with 10 % foetal calf serum, at 37 *C in a humidified, 5% CO2 atmosphere for up to 7 days. As a control, cells were also cultured on Thermanox discs. Cellular activity and proliferation were assessed by measurement of the total DNA^^ and the ^H TdR uptake on 2 and 4 day cultures which had been exposed to ^H TdR for the last 24 hours of culture. Alkaline phosphatase (ALP), a predictive marker of osteoblastic phenotype and differentiation, was assayed on 7 day cultures using a COB AS-BIO (Roche, UK) centrifugal analyser. The cellular morphology and activity of 1 and 3 day cultures were studied qualitatively using scanning electron microscopy. 4 los
250-]
0.2 wt% 0.3 wt% F-HA F-HA
Released in Medium Produced in Cell Lysate
Control
HA
L
^^
0.2 wt% F-HA
0.3 wt% F-HA
Figure 2 Results of biochemical assays indicating (a) relative cellular activity (proUferation) and (b) expression of cellular phenotype (differentiation).
Effect of Fluoride Substitutionon the Biocompatibilityof Hydroxyapatite:K.A. Hing et al.
21
RESULT S AND DISCUSSIO N Change s in the intensit y and broadnes s of peak s obtaine d by surfac e XRD of SBF soaked discs indicate d the presenc e of precipitate d £^tit e after 14 days on HA discs compare d to only 3 days on F-HA discs (Fig. 1). The increase d activit y in the substitute d apatite s was confirme d by the result s of the biochemica l assay s performe d on the HOB cell cultures , in which cells incubate d on the substitute d 2q)atite s demonstrate d elevated levels of ^H TdR incorporatio n compare d to pur e HA (Fig. 2a). In contrast , it appeare d tha t substitutio n did not effect the phenotypi c expressio n of the HOB cells, with cells incubate d on both HA and F-HA substrate s expressin g similar levels of ALP after 7 days invitro(Fig. 2b). Scannin g electro n microscop y of culture s demonstrate d tha t HOB cells incubate d on the contro l and 0.3 wt% F-HA substrate s possessed a mor e flattene d morpholog y (Fig. 3a,d) in compariso n to the HOB cells incubate d on the 0.2 wt% F-HA and particularl y the HA substrate s (Fig. 3b,c). Thi s agree s well with the findings of the ^H TdR incorporation , as a flattene d morpholog y generall y indicate s a well attache d cell^^ , wher e cell attachmen t has been shown to affect the proliferatio n rat e of osteoblast s in vitro^^.Furthermore , a protein-lik e layer was
Figur e 3 Scannin g electro n micrograph s demonstratin g typica l HOB cell morphologie s when culture d for 1 day on (a) Thermanox , (b) HA, (c) 0.2 wt% and (d) 0.3 wt% F-HA substrates .
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Bioceramics Volume if)
Figur e 4 Scannin g electro n micrograp h demonstratin g a bone-lik e ^atit e layer on a 0.3 wt% F-HA substrat e after 1 day in HOB cell culture . identifie d on both the F-HA substrate s after 3 days invitroand the formatio n of a bone-lik e apatit e layer on 0.3 wt% F-HA substrate s was noted after 1 day in vitro(Fig. 4). This behaviou r has previousl y been reporte d to be indicativ e of a bioactiv e surface^^*^^ . CONCLUSIO N The substitutio n of low levels of fluorid e into the apatit e lattic e has resulte d in an elevate d level of bioactivit y when compare d to tha t of stoichiometri c hydroxy^atite . ACKNOWLEDGEMENT S The author s wish to acknowledg e thefinancial suppor t of EPSR C (UK) and JNIC T (Portugal) . REFERENCE S 1 De Jon g WJ. , Rec. Trav . Chim. , 45,445,1926 2 Posner A.S. Physiologica l Reviews. 49 (4), 760-792,1969 3 Le Gero s R.Z., Le Gero s J.P. , In: An Introductio n to Bioceramics , Eds: Hench LI.. , Wilson J. Worl d Scientific, Singapore , 1993 4 Le Gero s R.Z., Traut z O.R., Le Gero s J.P. , Klein E., Bull. Soc. Chim . Franc . 1712-1713,1968 5 Vaugha n J.M . The Physiology of Bone. Clarendo n Press , Oxford , 1981 6 Posner A.S., Eane s E.D., Harpe r R.A., Zipki n I., Arch . Ora l Biol., 8, 549-570,1963 7 Posner A.S., BuU. Hosp. Join t Dis., 39,126-144,1978 8 Aoki H. Science and Medica l Application s of Hydroxyapatite . Takayam a Press , Tokyo , 1991 9 Le Gero s R.Z., Proc . Inter . Sem. Orthop . Res. Ed: Niwa S., Springer-Verlag , Nagoya , 1992 10 Jh a LJ. , Best S.M., Santo s J.D., Bonfield W., Bioceramics , 9,165-168,1996 11 Kukub o T., Kushitan i H., Sakk a S., Kitsug i T., Yamamur o T., J. Biomed. Mater . Res. 24, 721-734,1990 12 Kapuscinsk i J., Skoczyla s B., Anal. Biochem., 83, 252-257,1977 13 Hunte r A., Arche r C.W., Walke r P.S., Blunn . Biomaterials , 16 (4), 287-295,1995 14 Folkma n J., Moscona A., Nature , 273, 345-349,1978 15 Davies J.D . The Bone-Biomateria l Interface . Toronto , Universit y of Toront o Press , 1991 16 de Bruiji n J. Calciu m Phosphat e Biomaterials : Bone Bondin g and Biodegredatio n Properties , PhD. Thesis , Leiden , 1993
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
APATIT E PRECIPITATIO N IN BIPHASI C CALCIU M PHOSPHAT E CERAMI C AFTE R INCUBATIO N IN RABBI T SERU M AND IONI C SIMULATE D BODY FLUI D (SBF) R. Rohanizadeh, J.M. Bouler, D. Couchourel, M. Padrines and G. Daculsi Centre de recherche interdisciplinaire sur les materiaux d’interet biologique - UPRES 2159, 1 Place A. Ricordeau, 44042 Nantes cedex 01, France
ABSTRAC T To evaluate the influence of organic factors in the precipitation process, biphasic calcium phosphate blocks were incubated in ionic simulated body fluid (SBF) and rabbit serum. Fragments of block surface were observed and analyzed by transmission electron microscopy after 2 and 6 weeks, and the compressive strength of the blocks was measured after 8 weeks. The results showed that i) apatitic microcrystals appeared around ceramic crystals in both solutions; ii) organic factors such as proteins and enzymes may control crystal growth and nucleation; iii) precipitated microcrystals were in continuity with the lattice planes of hydroxyapatite (HA) ceramic crystals; and iv) microcrystals were more aggregated around HA than p-tricalcium phosphate implants. KEYWORD S Calcium Phosphate, Precipitation, Ultrastructure INTRODUCTIO N A hydroxyapatite (HA)/p-tricalcium phosphate (p-TCP) mixture (60/40% by weight) produces biphasic calcium phosphate (BCP) with optimal properties for bone substitution [1]. Precipitation of carbonated apatite resulting fi-om the dissolution/precipitation process has been observed after implantation of BCP [2]. Precipitated microcrystals decrease the void space between crystals, enhancing the mechanical properties of the ceramic [3]. Although precipitation is a physicochemical reaction, organic factors such as proteins and enzymes may influence this process, as in the mineralization of biological tissues. To investigate the effects of organic factors on precipitation, we incubated BCP blocks in two different solutions: ionic simulated body fluid (SBF) and rabbit serum. Transmission electron microscopy (TEM) was used to investigate crystal morphology, the structure of the precipitates, and the influence of the seed species (HA and (3TCP). The relative effects of precipitation on the compressive strength of incubated BCP were also studied. MATERIAL S AND METHOD S Three blocks of macroporous BCP (5x5x10 mm, Triosite^^, Zimmer, France) composed of 60% HA and 40% P-TCP were incubated in 2 ml of rabbit serum and in SBF prepared according to a previously described method [4]. The blocks were kept in sterile conditions in an oven at 37 . After 2 and 6 weeks, fragments of block surface were embedded in LR White and cut into ultrathin sections (90 nm) for TEM analysis. To measure compressive strength after 2 months, the ends of blocks were removed and each block was cut horizontally into 3 equal parts. TEM observations at 23
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low and high magnification s were then carrie d out at 200 Kv (Jeol 200 CX). Electro n diffractio n was performe d to identif y the structur e of cerami c and precipitate d crystals , and the pattern s were indexed and identifie d by mean s of a simulato r (Electro n Diffractio n 2.4, Laboratoir e de Metallurgi e Physique , LSPES ) which used store d crystallographi c dat a for known calcium phosphat e phase s and generate d pattern s for direc t compariso n with observe d diffractio n patterns . RESULT S AND DISCUSSIO N TEM studie s of fragments incubate d in SBF reveale d tha t needle-shape d microcrystal s precipitate d aroun d cerami c crystal s (Fig. la). The shap e and size of precipitate d microcrystal s did not var y in differen t areas . The precipitate s were very well crystallized , and the lattic e plane s of the microcrystal s could be observe d in high-resolutio n TEM (HrTEM ) (Fig. lb). These plane s of 0.82 nm and 0.53 nm, with an angle of 72 between them , corresponde d respectivel y to apatit e lattic e plane s of (100) and (101). Figur e lb shows tha t precipitate s in SBF grew along the a-axis of the apatit e structur e (lattic e plane s of (100)). Electro n diffractio n confirme d tha t the structur e of precipitate d microcrystal s was apatiti c (Fig. Ic), showing a ring diffractio n patter n of precipitate d microcrystal s in SBF. D-spacin g of the first and second rings was 0.35 and 0.28 nm respectively , correspondin g to apatit e lattic e plane s of (002) and (112).
Figur e 1. TEM (a), HrTE M (b) observations , and electro n diffractio n of precipitate d microcrystal s (P) aroun d cerami c crystal s (C). In rabbi t serum , precipitate d microcrystal s were also observe d aroun d cerami c crystal s (Fig. 2a). As they were poorl y crystallize d and very fragile unde r irradiation , their lattic e plane s
Apatite Precipitationin BCP After Incubation in Rabbit Serum and SBF: R. Rohanizadehet al.
25
could not be observed in high-resolution TEM. As in SBF, the electron diffraction pattern of microcrystals in serum showed a ring pattern corresponding to the apatite structure (Fig. 2b). In contact with ceramic crystals, the precipitates in serum took the form of small round microcrystals, whose shape and size changed with increasing distance from the crystal surface until they ultimately became flattened crystals. Comparison of the precipitation between two different solutions indicated that organic factors in the serum changed the shape, size and crystallinity rate of the precipitates. These organic factors did not modify the crystal structure of the precipitates, which was apatitic in both solutions. TEM observations in serum showed that the size of precipitates in contact with the ceramic surface was smaller than in SBF. In fact, some proteins may have adhered to the ceramic surface [5]. In biological fluid, non-coUagenic proteins such as osteopontin and bone sialoprotein [6] can affect crystal nucleation and growth. In serum, organic factors also inhibited growth along the a-axis, with crystals assuming flattened shapes.
Figure 2. TEM observation (a) and electron diffraction pattern (b) of precipitated microcrystals (P) around ceramic crystals (C) in serum. Figure 3a and b presents high-resolution TEM observations of the junction between ceramic and precipitated microcrystal. Figure 3a shows an HA crystal with lattice planes of (100). Precipitates were in continuity with the lattice planes of the MA crystal (homogeneous epitaxic growth). Figure 3b shows a p-TCP crystal. The fiinge along the edge of the crystal indicates that precipitation proceeded in a direction different from that of the lattice planes of p-TCP. Electron diffraction showed that apatitic precipitated microcrystals were more abundant and were aggregated more around HA than p-TCP ceramic, since both HA and microcrystals have an apatite structure. In terms of potential energy for the system, the growth of apatitic microcrystals that prolong HA lattice planes is more favorable than that developing within the p-TCP structure. Table 1 shows the compressive strength of incubated and non-incubated blocks. Compressive strength increased after incubation in both solutions, though this increase was greater for rabbit serum than SBF, probably because of the effect of drying on the organic material in the former solution. In fact, one resuh of the presence of precipitated microcrystals was a decrease in ceramic microporosity which enhanced the mechanical properties of the ceramic.
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Figure 3. a) Precipitated microcrystals (P) were in continuity with the lattice plans of HA crystal; b) Precipitated microcrystals were in a direction different from that of the lattice planes of p-TCP. Table 1. Compressive strength of incubated and non-incubated blocks. Before incubation Compressive strength (MPa)
4.07 – 0.02
After incubation in SBF 5.73 –0.01
After incubation in rabbit serum 6.32 –0.07
REFERENCE S 1. Daculsi G., Passuti N., Martin S., Deudon C , LeGeros R.Z. and Raher S., J. Biomed. Mater. Res., 1990, 24, 379-396. 2. Daculsi G., LeGeros R.Z., Heughebaert M. and Barbieux I., Calcif. TissueInt., 1990, 46, 2027. 3. Trecant M., Delecrin J., Nguyen J.M., Passuti N. and Daculsi G., J. Mater. Sci., 1996, 7, 227229. 4. Yamada S., Nakamura T., Kokubo T., Oka M. and Yamamuro T., J. Biomed Mater. Res., 1994, 28, 1357-1363. 5. Fujisawa R. and Kuboki Y., Biochim. Biophys. Acta, 1991, 1075, 56-60. 6. Hunter G. K., Hauschka P.V., Poole A.R., Rosenberg L.C., and Goldberg H.A., Biochem. J., 1996, 317, 59-64.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
APATITE PRECIPITATION IN BBPHASIC CALCIUM PHOSPHATE CERAMIC AFTER IMPLANTATION: INFLUENCE OF IMPLANTATION SITE R. Rohanizadeh, M. Trecant-Viana, J. Delecrin, J.M. Bouler and G. Daculsi Centre de recherche interdiscipiinaire sur les materiaux d’interet biologique - UPRES 2159, 1 Place A. Ricordeau, 44042 Nantes cedex 01, France ABSTRACT Diphasic calcium phosphate blocks implanted into rabbit trabecular bone and muscle were recovered 18 weeks later and observed and analyzed by transmission electron microscopy, electron diffraction and electron microprobes. The resuhs showed that: i) apatitic microcrystals appeared by secondary nucleation at both bone and muscle sites; ii) precipitated microcrystals were found around ceramic crystals at bone sites but distributed randomly in micropores at the muscle site; iii) the ratio of calcium to phosphate was higher for microcrystals at bone than muscle sites; and iv) precipitated microcrystals around (J-tricalcium phosphate crystals were less aggregated to the ceramic surface than those around hydroxyapatite. These findings suggest that microenvironmental parameters such as fluid circulation and interaction of ceramics with proteins or cells affect the physicochemical dissolution/reprecipitation process. KEYWORDS Calcium Phosphate, Precipitation, Ultrastructure INTRODUCTION The most widely tested types of bone graft materials are hydroxyapatite (HA) and Ptricalcium phosphate (P-TCP). A mixture of HA and P-TCP produces biphasic calcium phosphate (BCP) which possesses the reactivity of 3-TCP and the stability of HA, providing optimal properties for resorption/bone substitution [1]. Calcium phosphate ceramics undergo physical and chemical changes in biological fluids, causing the precipitation of apatitic microcrystals between ceramic macrocrystals [2]. Precipitation is a physicochemical process which restores the balance between ceramic and fluid. In in vivo conditions calcium phosphates are also in contact with proteins, cells and enzymes, and these biological factors could influence the precipitation process, as in the mineralization of biological tissue. In this study, MBCP (macroporous biphasic calcium phosphate - 70% HA and 30% PTCP) blocks were implanted into rabbit bone and muscle to investigate the influence of the two different environments on precipitation process, the junction between ceramic crystals and precipitated microcrystals, and the influence of two ceramic compounds (HA and p-TCP crystals) on the precipitation process. MATERIALS AND METHODS Twenty cylinders form implants of macroporous BCP (5x5 mm) were implanted into five rabbits. Each animal received 4 implants: two in lefl: and right femurs and two in left and right thighs. The animals were sacrificed after 18 weeks, and the implants were recovered, fixed in 80% 27
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ethyl alcohol and embedde d in methylmethacrylate . A fragmen t of the implan t surfac e was cut with a knife into ultrathi n sections (90 nm thick ) and examine d in transmissio n electro n microscop y (TEM ; Jeol 200CX, 200 Kv) to stud y the morpholog y and spatia l distributio n of precipitate d microcrystal s at two implantatio n sites. Electro n diffractio n and electro n microprob e analysi s were also performe d to identif y the structura l and chemica l component s of cerami c and precipitate d crystal s and the influenc e of HA and P-TC P on precipitation . RESULT S AND DISCUSSIO N TEM studie s showed tha t precipitate d microcrystal s had completel y or partiall y filled the spaces between cerami c crystals . Variation s in the quantit y and densit y of these microcrystal s were observed , especially at muscle sites. A differenc e in the spatia l distributio n of precipitate d microcrystal s was noted at the two sites. The precipitate s were dense and surrounde d cerami c crystal s at bone sites (Fig. 1), while precipitatio n was generall y less dense and occurre d randoml y in the spaces between cerami c crystal s at muscle sites (Fig. 2). This differenc e could have been due to the circulatio n of biological fluids in the microporosity . Because of the natur e of angiogenesi s in bone implants , the circulatio n of biological fluid is greate r tha n in muscle implant s which ar e surrounde d by muscle fibers formin g a capsul e tha t canno t easily be penetrate d by fluids. Fluid circulatio n in the microporosit y of bone implant s could be a factor limitin g precipitatio n between cerami c crystals .
Figur e 1. TEM observatio n showing precipitate d microcrystal s (P) and cerami c crystal s (C) in a bone implan t
Figur e 2. TEM observatio n shovin g precipitate d microcrystal s (?) and cerami c crystal s (C) in a muscle implan t
Figur e 3 shows a rin g patter n of precipitate d microcrystal s at a muscle site in which the ring s correspon d to apatit e (002) and (112) lattic e planes . Electro n diffractio n demonstrate d tha t the precipitate d microcrystal s at both bone and muscle sites were apatite . However , the Ca/P weight rati o obtaine d by microprob e analysi s on precipitate d microcrystal s was higher at bone (2.1 –0.1) tha n muscle sites (2 – 0.05). The lower Ca/P rati o at muscle sites indicate d tha t the microcrystal s containe d less carbonat e or had a high rat e of additiona l non-apatiti c phase s [3]. The rin g electro n diffractio n patter n was not an adequat e mean s of distinguishin g between polycrystal s when additiona l minor phase s formed .
Apatite Precipitationin BCP After Implantation: R.Rohanizadehet al.
29
High-resolutio n TEM demonstrate d tha t epitaxi c growt h was limited at muscle sites, wher e microcrystal s were less associate d with cerami c crystal s and showed a rando m orientation . In physicochemica l terms , epitaxi c growt h can occur at both bone and muscle sites. As in the mineralizatio n of biological tissues, protein s influenc e the precipitatio n process . At a muscle site, some protein s may adher e to the cerami c surfac e and contro l or inhibi t crysta l growth[4-5] . At a bone site, absorbe d nonspecifi c protein s ar e regularl y eliminate d from the cerami c surfac e due to cellular degradation , and growt h is controlle d by specific protein s such as osteoponti n and osteonecti n [6-7]. At both sites, cerami c crystal s identifie d as HA by electro n diffractio n (Fig. 4a) were surrounde d by precipitate s mor e closely associate d with the crysta l surfac e tha n those identifie d as P-TC P (Fig. 4b). As the precipitate d microcrystal s were apatite , the structur e of HA may be mor e favorabl e tha n tha t of (J-TC P to the growt h of apatit e microcrystals .
Figur e 3. Electro n diffractio n patter n of precipitate d microcrystal s (P) at a muscle site; the ring s correspon d to apatiti c lattic e plane s of (002) and (112).
Figur e 4. Electro n diffractio n in a bone implant ; a: an HA crysta l identifie d by electro n diffractio n was surrounde d by precipitate s mor e associate d with the crysta l surface ; b: a P-TC P crysta l
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identified by electron diffraction was surrounded by precipitates less associated with the crystal surface. CONCLUSIO N The features of precipitated microcrystals in implanted ceramics differ according to the implantation site. In the mineralization of biological tissues, elements such as cells and proteins probably influence the local precipitation process. Ceramic/protein and ceramic/cell interactions and their effects on dissolution and reprecipitation have not been clearly identified and require further investigation, e.g. in vitrostudies in different simulated body fluids. REFERENCES 1. Daculsi G., Passuti N., Martin S., Deudon C , LeGeros R.Z. and Raher S, J. Biomed. Mater. Res., 1990, 24, 379-396. 2. Daculsi G., LeGeros R.Z., Heughebaert M. and Barbieux I., Calcif,TissueInt., 1990, 46, 2027. 3. LeGeros R.Z., Daculsi G, Gregoire M., Heughebaert M. and Gineste M., Bone-Bonding.Reed HealthcareCommunication,1992, 89, 201-212. 4. Hunter G. K., Hauschka P.V., Poole A.R., Rosenberg L.C. and Goldberg H.A., Biochem. J., 1996, 317, 59-64. 5. Hunter G.K., Kyle C.L. and Goldberg H.A., Biochem. J., 1994, 300, 723-728. 6. Nagata T., Todescan R., Goldberg H.A., Zhang Q. and Sodek J., Biochem. Biophys. Res. Commun., 1989, 165, 234-240. 7. Fisher L.W. and Termine J.D., CliN Orthop., 1985, 200, 362-385.
GLASS CERAMICS BIOACTIVITY
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BIOACTIVIT Y AND STRUCTUR E OF ORGANICALL Y MODIFIE D SILICAT E SYNTHESIZE D BY TB E SOL-GE L METHO D Kanji Tsum, Satoshi Hayakawa, Chikara Ohtsuki, and Akiyoshi Osaka Biomaterials Lab., Faculty of Engineering, Okayama University Tsushima, Okayama-shi 700 Japan
ABSTRAC T We investigated effects of H Q and calcium nitrate(Ca(N03)2»4H20) contents in gelling solutions on bioactivity and microstructure of Onnosils-type organically modified silicates. Analysis of 29si MAS NMR spectra indicated that amounts of H Q do not show any significant effects for local structure aroimd Si atoms. The SEM-EDX analysis of fracture-surface for Ormosils-type sample showed that the pore size and amount of Ca incorporated in their structure depended on HCl amounts. We concluded that greater bioactivity was attributed to incorporation of more calcium ions in the polymer due to addition of larger amounts of HCl and/or Ca(N03)2*4H20. INTRODUCTIO N Bioactive ceramics can directly bond to Uving bone via an apatite layer deposited on their surface when they are embedded in the body[l]. Although some of them are already clinically used, they still could not be used for soft-tissue replacements because of their high Young modulus. Thus they should be modified to more softness and toughness for clinical application to repairing sites that needs higher flexibility. Incidentally, Ohtsuki et al. reported that Ca(II) ions and silanol group(Si-OH) were the important components for formation of biologically active apatite layer in a body fluid[2]. One might expect that organically modified ceramics containing calcium and silicates shows the apatite-forming abihty, i.e. bioactivity. On this basis, we recently synthesized Ormosils-type organically modified silicate through sol-gel processing, and found that thus synthesized material formed apatite layer on its surface in a simulated body fluid(SBF)[3,4], This indicates thus synthesized Ormosils-type organically modified ceramic is bioactive, and has potential of serving basis for novel bioactive materials. However, effect of synthetic conditions on the bioactivity and microstructure are not revealed yet. In this study, we investigated effects of the amount of H Q or Ca(N03)2»4H20 added to the solutions for Ormosils-type samples on their bioactivity and microstructure. MATERIAL S AND METHOD S We synthesized caldum-containing Ormosils-type organically modified ceramics starting from ploy(dimethylsiloxane) (PDMS,-(SiO(CH3)2)n-)> tetraethoxysilane (TEOS), Si(OC2H5)4) and calcium nitrate (Ca(N03)2*4H20). One of the typical compositions for starting materials was: TEOS:PDMS:HCl:H20:Ca(N03)2*4H20 = 1 : 1.67
: 0.05 : 3 : 0.05. Detailed procedure was given in the previous reports[3,4]. In the present study, 33
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the ratio of HCl and Ca(N03)2«4H20 was varied in the range of Ha/TEOS=0.05-0.1(mol) and Ca/TEOS=0’-0.1(mol), respectively. The synthesized samples were cut to 15x10x1 mm^ in size, and the surface was polished with a #2000 emery paper. Then they were gently rinsed with ethanol and dried. Microstructure of the fracture-surface was observed with a scaiming electron microscope(SEM), JEOL JSM-6300, equipped with an energy dispersive X-ray analyzer (EDX, PhiUips EDX-4). The specimens were soaked in a simulated body fluid(Kokubo solution) at 36.5 C up to 14 days. The Kokubo solution was prepared as described by Cho et al\5\. It has been proved[6] that the in vivoapatite formation on implants can almost fully be reproduced in vitroexperiments in the Kokubo solution. Thin-film X-ray diffraction(TF-XRD) was used to examine formation of hydroxyapatite layer on the surface of the specimens after soaked in the Kokubo solution. To evaluate local structure around Si atoms in the samples, ^^Si Cross-polarization Magic Angle Spiiming (CP MAS) NMR spectra was measured for the samples after pulverized. In the NMR measurement, the sample spiiming speed was 5-6kHz at the magic angle(54.7 ) to the external field. 29si NMR spectra of the samples were measured at 9.4T on a JEOL JNMGX400 FT-NMR spectrometer, equipped with a TU-GSX400 MAS probe. 29si NMR spectra were acquired at 79.3MHz with 5.0-(xs pulses, 12.0s recycle delays, 15.0(xs dead time. The signals for an 80’-400 pulses were accumulated. The chemical shift (6 in ppm) of 29si was determined using tetramethylsilane(TMS)(6=0 ppm : ^^Si chemical shifts where 6 denoted the isotropic chemical shift) as an external reference substance. Poly dimethyl silane(6=-34.0 ppm) was used as the secondary external reference. RESULT S AND DISCUSSIO N Figure 1 shows the TF-XRD patterns of the samples with varied ratio of HCl after soaked in the Kokubo solution up to 14 days. They are denoted as "HCl X Ca Y " where X is the molar ratio HCl/TEOS(mol) and Y is the molar ratio Ca/TEOS (mol). Fig. 1 indicates that apatite formation on the Ormosils-type samples with a fixed ratio of Ca(N03)2*4H20 was enhanced by the increasing of the amount of HCl added. For the samples with varied ratio of Ca(N03)2*4H20 at H Q of Y=0.1, apatite formation was enhanced by the increasing of the amount of calcium nitrate added. Thus the bioactivity of the Ormosils-type materials was effected not only additional amounts of Ca(N03)2*4H20 but also the amounts of H Q solution. The results of 29si CP MAS NMR measurements shows that two groups of Si atoms were distinguished on the basis (rf the chemical shift: the first peak(-10 to -25 ppm) was due to 02SiMe2(Me=-CH3) units from PDMS while the other peak(-100 to -110 ppm) was due to Si02 units originated from TEOS. Both peaks were very broad because they were the envelopes of a few component peaks. Deconvolution of them on the basis of the chemical shift data in the literature[7] indicates that the former peak had three components and the latter had two. The former curve was fitted with Lorentz functions and the latter was fit with Gaussian functions. Each component in the former peak is assigned to PDMS chain(D) , cyclic ohgomers(E)cycUc) and copolymerized species(D(Q)), respectively. The Si04 units for the latter peaks are denoted as (y^ groups where n represents the number of bridging oxygen atoms around a Si atom in the siUca matrices. Thus obtained peak positions and the relative peak area (%) are given in Table 1. Table 1 indicates that amount of H Q does not show a significant effect for the local structure of Si atoms. Similar was found for the effect of Ca(N03)2*4H20. Fig. 2 shows the results of SEMEDX analysis of fractured surfaces for Ormosils-type samples. Size of pores in the samples seems to increase with increasing the ratio of HCl solution. In addition, the concentration of Ca
Bioactivity and Structureof Organically Modified Silicate: K. Tsuru et al.
25
25
30 35 2e/de g
30 35 2e/de g
25
35
30 35 2e/de g
Figur e 1 TF-XR D pattern s of sample s synthesize d with variou s amount s of HCl after soaked in the Kokub o solution up to 14 days. O : Apatit e
in the sample s also increase d with increasin g the rati o of H Q and Ca(N03)2*4H20 . Therefore , greate r apatite-formin g abilit y has been attribute d to incorporatio n of mor e calcium ions in the polymer due to large r amount s of HQ . Dissolution of Ca(II ) ions in the structur e not only favor s formatio n of a sihca hydroge l layer conductiv e to apatit e nucleatio n but increase s in degre e of siq)ersaturatio n for ^atit e in the Kokub o solution . Therefore , we conclude d tha t controllin g the incorporatio n of Ca(II ) ions is essentia l for preparatio n of bioactiv e Ormosils-typ e organicall y modified ceramics .
Tabl e 1. Peak position s and relativ e peak are a (%) of the 29si CP MAS NMR peaks . The left tabl e is due to 02SiMe2 units(Dgroup ) from PDMS. Theright tabl e is due to Si02 units(Q ^ group ) originate d from TEOS . Dg roups Sampl e nam e
0*^9 roups
6 (relative pea k area , % ) D(Q)
Dcydic
6 (relative pea k area , % )
D
Q3
Q4
HCl 0. 1 -CaO.OS
-17. 2 (28 )
-19. 6 (51 )
-21. 7 (21 )
-102. 3 (24% )
-108. 5 (76% )
HCl 0.0 9
C a 0.0 5
-17. 1 (29 )
-19. 7 (53 )
-21. 9 (18 )
-102. 7 (22 )
-108. 6 (78 )
HCl 0.0 7
C a 0.0 5
-16. 9 (30 )
-19. 7 (49 )
-22. 0 (21 )
-102. 9 (28 )
-108. 7 (72 )
HCl 0.0 5
C a 0.0 5
-17. 0 (34 )
-19. 6 (45 )
-21. 9 (21 )
-103. 0 (31 )
-108. 6 (69 )
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HCI 0.0 5 rCa 0.0 5
SiKa 99.8% I CaKaO.2 %
SIKa 984%; CaKal6%
Figure 2 SEM-EDX analysis of fracture-surfaces for Ormosils-type samples. Apatite : X = no apatite after 14 days Apatite 3 days = apatite deposited after 3 days
CONCLUSIO N Apatite-forming ability of Ormosil-type organically modified ceramic is enhanced by controlled amounts of HCl and Ca(N03)2*4H20. Analysis of ^^Si MAS NMR spectra indicates that additional H Q does not show any significant effects for local structure around Si atoms. The porosity and Ca content incorporated in the structures depends on amounts of HCl. We have concluded that controlling the incorporation of Ca(II) ions is essential for preparation of bioactive Ormosils-type organically modified ceramics.
REFERENCES
1. Hench, L. L. and Wilson, J. In: An introductionto bioceramics.World Scientific, Singapore, 1993,1-24. 2. Ohtsuki, C , Kokubo, T. and Yamamuro, T., J. Non-Cryst.Solids, 1992,143,84-92. 3. Tsuru, K., Ohtsuki, C. and Osaka, A. 7. Mat. Sci.,Mat. Med., 1997,8, 157-161. 4. Tsuru, K., Hayakawa, S., Ohtsuki, C. and Osaka, A. In: BioceramicsVolimie 9, Pergamon, Oxford, 1996,419-422. 5. Cho, S.B., Nakanishi, K., Kokubo, T., Soga, N., Ohtsuki, C , Nakamura, T., Kitsugi, T. and Yamamuro, T. J. Am. Ceram.Soc, 1995, 78, 1769-1774. 6. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T. and Yamamuro, T., J. Biomed.Mater. /?ej.,1990, 24,721-734. 7. Iwamoto, T., Morita, K., and Mackenzie, J. D.,J. Non-Cryst.Solids, 1993,159,65-72. ACKNOWLEDGEMENT S The authors thank Prof. T. Yoko of Kyoto University for his helpful advice and assistance in the NMR measurements. One of the authors(K. T.) gratefully acknowledges the Research Fellowship of the Japan Society for the Promotion of Science for Young Scientists.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
HYDROXYAPATIT E FORMATIO N ON BIOACTIV E GLASS COATE D TITANIU M Cheol Y. Kim and Sungmin Kwon Dept. of Ceramic Engineering, Inha Univ. 253, Yonghyun-dong, Nam-ku, Inchon,402-751. Korea.
ABSTRAC T Two different bioactive glasses were coated on a titanium to give it a bioactivity, and then the bonding characteristics between a titanium and a coated layer and the hydroxyapatite formation in a simulated body fluid(SBF) on the coated specimens were studied. TisSis was found at the interface between titanium and glass coat in the groundcoat layer, and this enhaced the titanium-glass bonding. The hydrox}^apatite formation was observed only on the covercoat fired under 800 C, which is in an amorphous phase, and was retarded for the sample fired over 850 C, which contains an oxyapatite. KEYWORDS : Bioactive Glasses, Titanium, Hydroxyapatite, Groundcoat, Covercoat INTRODUCTIO N It has been well known that bioactive glasses are one of the best candidates for an implant material because of their excellent bonding behavior to the living tissue. ^^ The major drawback of these materials to use as a practical implant is their poor mechanical strength. To solve this problem, several works on coating the bioactive glasses to a strong substrate such as alumina and metal have been carried out.^^ A metal substrate has an advantage over an alumina because the former has better mechanical properties. However, most of metals easily oxidize specially at high temperarure and release some harmful ions when implanted for a long period of time. A titanium is now known as one of the best biometals for an implant material because it is chemically inert in a body. In the present study, therefore, two kinds of bioactive glasses are coated on a titanium by using a double coating method. The primary objectives of this study are 1) to find the bonding behavior between glass and titanium, and 2) to find the hydroxyapatite formation on the glass coated layer in a simulate body fluid. 37
38
Bioceramies Volume10
MATERIAL S AND METHOD S Two types of glass powder, as shown in Table 1, were prepared by melting the glass batch in a Pt-crucible at 1500 C for 2 hours followed by pulverizing them to less than 44/mi in an agate mortar. The glass powder was dispersed in an aceton, and spray-coated on a titanium. Glass 55SF and 5 5 SB were used for a groundcoat and a covercoat, repectively. The glass coated titanium was fired in a tube furnace at the temperature ranging from 1150 Cto 1300 C. Ar gas was flowed into the tube furnace at the rate of 2.5 1/min to prevent the oxidation of the titanium. The bioactive glass coated titanium was reacted in a simulated body fluid (SBF), which has pH of 7.3, and the ratio of the bioactive glass coated area to volume of the reaction solution was set at 0.1 cm’^ and the reaction temperature at 37 C. The crystalline phases of the heat-treated bioactive glass coated layer and the hydroxyapatite crystalline phase after the reaction test were examined by a thin film x-ray diffractometer. Tabl e 1. Batch Compositions of Bioactive Glasses
(mol%)
sample
Si02
P2O5
Na.O
CaO
Cap.
B2O3
55SF
55.1
3.4
9.2
27.8
4.5
-
50SB
50.1
3.4
9.2
32.3
-
5
RESULT S AND DISCUSSIO N Bondin g Behavior between Glass and Titaniu m When 55SF glass was coated on a pure titanium and fired under Ar-gas condition at 1300 C for 2 minutes, the crystalline phases of the glass layer are shown in Fig.l. TisSia crystal was found at the interface, and this crystal promotes the bonding of the glass layer to the titanium. 55SF glass coated Ti-metal was fired between 1150 C and 1300 C, and the XRD patterns are shown in Fig.2. Afluorapatitecrystal was obtained at a lower firing temperature. Thefluorapatitewas mehed as increase in temperature and a -wollastonite was developed at 1300 C. It is believed that titanium ions penetrated into glass layer at a higher firing temperature and acted as a nucleating agent for an (7-wollastonite crystal formation. Generally an oxyapatite and fluorapatite crystal form at a lower temperature without nucleating agent. No hydrox>apatite formation was observed when these samples were reacted in the SBF.^^
Hydroxyapatite Formation on Bioactive Glass Coated Titanium:C. Y. Kim and S. Kwon 39
w a-Wollastonite
W a-Wollastonit e W
I
I
T Titanium
|i>ii MmA»nM^u,0umiiK^
i3ao c
50^m
wl w s gs ^
I w
^
1250X
w
c
F Fluorapatit e
83p.m
100^m
I t I i I 20
30
40
50
26 (degree)
60
70
10
20
30
40
50
60
20 (degree)
Fig.l XRD patterns of 55SF glass coated Fig.2 XRD patterns for the surface of 5SF with increase in depth, heat-treated coated titanium, heat-treated at variat 1300 C in Ar atmosphere ous temperature for 2 mins Hydroxyapatit e Formatio n on Overcoate d Glass Layer Because no hydroxyapatite was formed on the groundcoat, another low melting glass 50SB was overcoated on the 55SF, and fired at 750T- 900T for 2 min. The glass layer crystallized into an oxyapatite when heat-treated at 900 C, and an amorphous overcoat layer was obtained when heat-trested at 800 C as shown in Fig. 3. These double-coat samples were reacted in SBF for 48 hours, and their XRD patterns are shown in Fig.4. The samples heat-treated at 850 C and 900 C showed oxyapatite crystal which were formed during the heat treatment. The sample heat-treated at 800’’C, however, showed a typical X-ray diffraction pattern of a newly formed hydroxyapatite precipitated on the glass surface. The reason for the delay of the hydroxyapatite formation on the oxyapatite crystal containing samples is because the matrix phase txu’ns into a chemically durable phase, and most of phosphorus ions are trapped in the oxyapatite crystal. This indicates that the leaching of phosphorus ions from the bioactive glass plays a very important role for the easy formation of the hydroxyapatite even in the phosphorus ion containing solution like SBF.
40
Bioceramics Volume10 0 Oxyapatite O Oxyapatite
1
H Hydroxyapatite 1
01
1 I L*M**«M(
900 C
KaJLjimlJliil
L-H’*»W\.
850 C 1
i " L^NIWWV K 1 ""^
20
30
40
50
60
20 (degree)
Fig.3 XRD patterns for the surface of 50SB coated titanium. Heat treated at various temperatures for 2 minutes.
10
800 C
:Tr^
1
J .1 .i_L _
10
1 1
20
30
40
50
60
2e (degree)
Fig.4 XRD patterns for the surface of 50SB coated titanium after corrosion in trisbuffer solution for 48 hrs.
CONCLUSIONS Bioactive groundcoat glass has a good chemical bonding to titanium by producing TisSis crystal at the interface. No hydroxyaptite formation was observed on the groundcoat after reaction in SBF, while hydroxyapatite crystal was obtained on the covercoat glass, which was fired at the temperature lower than 800 C. REFERENCE S 1. P. Ducheyne, P. Bianco, S. Radin and E. Schepers, Bone-BondingBiomatehals,Reed Healthcare, conmiun., 1992, 1-12 2. J. K. Kim and C. Y. Kim , J.Kor.Ceram.Soc, 1990, Vol.27, No.7, 925-933 3. C. Ohtasuki. A. Osaka and T. Kokubo, Bioceramics,1994, Vol. 7, 73-78
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EFFEC T OF MULTIVALEN T CATION S IN CALCIU M SILICAT E GLASSE S ON BIOACTIVIT Y Naoki Imayoshi, Chikara Ohtsuki, Satoshi Hayakawa and Akiyoshi Osaka Biomaterials Lab., Faculty of Engineering, Okayama University, Tsushima, Okayama 700, Japan
ABSTRAC T We examined the apatite formation for CaOSiO^ glasses containing Vp^, WO3, Ta205or Cr03 after they were soaked in a simulated body fluid(SBF) for various periods. Surface reaction of glass with SBF and reconstruction of an Si-O-Si network in the surface layer were studied with ^^Si MAS-NMR spectroscopy which gave the average ratio of bridging/nonbridging oxygen for an Si atom(Q"). It was shown that Ta^O^ and Cr03 effectively depressed the apatite formation and that V^Og and WO3 up to 5mol% caused little influence while dissolution of Ca( II) and Si(IV) was enhanced. Analysis of the ^^Si MAS-NMR spectra showed that the distribution of Q" was different for each transition metal oxides, and thus suggested that the structure of silica hydrogel layer was affected by those cations. INTRODUCTIO N Bioactive glasses and glass-ceramics bond to living bone directly[l]. The condition for those to bond to living bone, i.e. bioactivity, is to form a bone-like apatite layer when embedded in the bony defect. A binary glass 50CaO-50SiO2(mol%) forms the apatite layer, hence it serves a basic system for bone substitutes. Addition of third components like AI2O3 and TiO, depresses the apatite formation [2]. Thus, those cations cause opposing effects on bioactivity of glasses as they modify chemical properties and structure of a silica gel layer that is assumed to favor the apatite nucleation. It is essential to understand what structural and chemical role the third cations have in bioactive glasses. Transition metal cations are worth examining such effects since their structural roles may depend on glass composition. In this study we examined the apatite formation of SOCaO-SOSiO^ glasses doped with 5mol% transition metal oxides as well as dopant-free glass soaking in a simulated body fluid(SBF) which are similar in inorganic composition to that of the human blood plasma. Moreover, ^^Si Magic Angle Spinning(MAS)-NMR spectra of pulverized glasses were measured before and after soaked in the SBF to determine glass structure. MATERIAL S AND METHOD S Glasses were prepared due to an ordinary melt-quench method in a series of composition: x (V2O3, WO3, Tap^ or CrO3)(50-x/2)CaO(50-x/2)SiO2 where the molar ratio CaO/SiO, was maintained to 1. The obtained glasses were cut into rectangular specimens 15 X 10 X Imm^ and polished with a diamond paste. They were soaked in the SBF with pH=7.25 at 36.5 C. The SBF was prepared as described by Cho eta/[31. After soaking the glasses into the SBF, their surface structure was examined by FT-IR spectroscopy and thinfilmX-ray diffractometry(TF-XRD). The concentrations 41
42
Bioceramics Volume10
of calcium, silicon and phosphorus in the SBF after each period of soaking were measured by inductively coupled plasma(ICP) emission spectroscopy. ^S i MAS-NM R measuremen t The glanular glasses with 150-300^im in diameter were served for the^^Si MAS-NMR measurement before and after soaking in the SBF for various periods. ^^Si MAS-NMR spectra were recorded at 9.4T on a JEOL JNM-GSX400 FT-NMR spectrometer, equipped with a TU-GSX400MAS probe. Samples were placed in a zirconia sample tube. The sample spinning speed at the magic angle to external field was 5-6kHz. ^^Si MAS-NMR spectra were measured at 79.3MHz with 4.0-^s pulses, 2.5s recycle delays. About 1000 pulses were accumulated. Poly dimethyl silane(PDMS) was used as secondary external reference substance to determine the -^Si chemical shifts. RESULT S AND DISCUSSIO N Figure 1 shows the TF-XRD patterns of the surfaces of the metal oxide-doped and oxide-free 50CaO-50SiO, glasses soaked in the SBF for 7 and 30 days. The diffraction peaks near 26 and 32 were assigned to (002) and an envelope of (211), (112) and (300) of apatite, respectively. Fig. 1 indicates that SOCaO-SOSiO^ glass, 5mol%V203 and 5mol%W03 doped-glasses formed apatite on their surfaces in the SBF within 7 days, whereas 5mol%Ta205 and 5mol%Cr03 doped-glasses did not form it even after 30 days. This difference on apatite-forming ability among glasses suggests changes of chemical property and structure of SOCaO-SOSiO^ glass due to the addition of the transition metal oxides. The concentrations of calcium and silicon of the SBF due to soaking the glasses are shown in Figs. 2 and 3. They increased with the period for 5mol%V205 and WO3 doped-glasses, whereas only a slight
(b)30d Q
X-ra y 50CaO’50SiO 2
’’^w V ^
5mol%V,O s
0
5mol%W0 3
CI
e
A 2 _A^A^ _ 5mol%Ta20 5 5mol%Cr0 3 ^ *f*l’*rrHrT i 1 1 1 1 1 Trt’^’i"^"’* "
20
25
30 35 2e/de g
40
20
25
30 35 2e/de g
40
Figure 1 TF -XRD patterns of 50CaO-50SiO 5mol% Vp^, WO3, Ta203 and Cr03 doped-glasses soaked in the simulated body fluid for 7 days (a) and 30days (b). O: Apatite
Effects of MultivalentCations in Calcium Silicate Glasses on Bioactivity:N. Imayoshi et al. 10 §2. 5 S ^2. 0 .0
(b)[SJ ]
-J
_
J
^,^
1 1.5
ii.o
-
s o
O.Or
10
15 20 Time/da y
25
43
-^ crS^m 15 20 i 25
10
6
30
135
Tim e / da y
Figure 2 Changes in Ca (a) and Si (b) concentrations of the simulated body fluid due to immersion of 50CaO-50SiO 5mol% V p ^ , WO3, Idifi^ and Cr03 doped-glasses in the simulated body fluid. D: 50CaO-50Sia, A: 5mol%V203, A: 5mol%W03, O: 5mol%Tap3 and FI: 5mol%Cr03. "
increase was noticed for 5mol% Cr03 doped-glasses. There appears no changes in the concentrations for 5mol% Ta203 -doped glass. This means the addition of Ta^O, improves chemical durability of 50CaO-50SiO^ glass and suppresses the dissolution of calcium and silicon. The increase in calcium favors the apatite formation due to increase in the degree of supersaturation with respect to hydroxyapatite, and the release of silicon from glasses shows formation of silanol groups which plays an important role on apatite nucleation. Similar analysis indicated release of vanadium, chromium and tungsten from the glasses in the SBF, although tantalum ion was not detected even after soaking for 30 days. Therefore, the dissolution of vanadium and tungsten ions from their glasses in the SBF did not affect the apatite formation. In contrast, 5mol%Cr03 doped-glasses did not form the apatite layer although release of calcium and silicon was detected. These results suggest that the dissolved chromium ions in the SBF remarkably suppresses the nucleation and/or growth of the apatite. Figure 3 shows ^^Si MAS-NMR spectra of 50CaO-50SiO,, 5mol%V,03, WO3, Ta,05 and Cr03 doped-glasses before soaking in the SBF. Each peaks were deconvoluted into two or three Gaussian functions on the basis of least square fitting. Each component was assigned to Q^ ( ? or Q^on the basis of the reference data[4], where Q" shows the number of bridging oxygen. Relative peak area of each Q" group (n=2-4) of 50CaO-50SiO, does not differ from that for 5mol% transition metal oxides. This means that the apatite formation is not associated the initial local structure around an Si atom. Fig. 4 shows 2^Si MAS-NMR spectra of 5mol% Vp^and WO3 doped-glasses after soaking in the SBF for various periods. Leaching of calcium ions from the glass surfaces forms the Si-OH bonds of Q2 and Q^ groups without changing much of their ratios. However, the fraction of Q^ decreased for 5mol%V205 and WO3 doped-glasses with longer periods, probably due to polymerization of the SiOH bonds of Q2 groups. The increase in the fraction of Qt is attributed to the formation of Q^ units on the glass surface in the SBF. From Figs. 1 and 4, it is concluded that Q"^ units are favorable to apatite formation. Similar analysis for 5mol%Ta205 and Cr03 doped-glasses suggested absence of such changes in Q^ and Q^ for those glasses. A high field chemical shift, found in Fig. 4, for Si in the Q^ groups suggests that some of the Q^ groups involve Si-O-P bonds.
44
Bioceramics Volume10 I I H | H I H I I I I | I I I I | I I I 1 | I I I I | I I H |IH I I | I I I
’S i MAS-NM R spectr a
I l l i | n i l 11I I i | I II I I n i l | i i i 1 1 I I 1 1 | i 1 1 1 1 I I I
5mol % V^O g
Q-
^4
50CaO-50SiO ,
5mol%Cr0 3 tHliliiiilmiLifH l
Figure 3 ^^Si MAS-NMR spectra of pulverised 50CaO-50SiO,, 5mol%V,03, WO3, Ta,03 and CrOj doped-glasses before soaking in the simulated body fluid.
Figure 4 ^^Si MAS-NMR spectra of pulverised 5mol%V20g and WO3 doped-glasses after soaking in the simulated body fluid.
SUMMAR Y The apatite formation for 50CaO-50SiO, glass respectively varied by addition of transition metal oxides. The addition of 5mol% V^O^ and WO3 hardly effected the apatite formation in the SBF, while the addition of 5mol% TaPg and CrOj remarkably suppressed the apatite formation. The addition of the former increased in calcium and silicon concentration and the structure of silica gel layer as a result of chemical reaction in the SBF favored the apatite formation. But the addition of the latter improved the chemical durability and suppressed the formation of silica gel layer. Moreover dissolution of chromium ion in the SBF suppressed the apatite formation. REFERENCES 1. Hench, L.. L., Splinter, R. J., Allen, W. C. and Greenlee, T K. J, Biomed.Mater.Res. Symp.,2, 117-141(1972). 2. Ohtsuki, C , T. Kokubo, T, Takatsuka and T. Yamamuro, J. Mater.Sci.: Mater Med.,3, 119125 (1992). 3. Cho, S.B., Nakanishi, K., Kokubo, T, Soga, N., Ohtsuki, C , Nakamura, T, and Yamamuro, T. J. Am. Ceram.Soc, 78, 1769-1774 (1995). 4. Selvaray, U., Rao, K. J., Rao, C. N. R., Klinowski, J. and Thomas, J. M. Chem.Phys. Lett.,114, 24(1985). ACKNOWLEDGMENT S The authors thank Professor F. Horii and Mrs. K. Omine of Kyoto University for their helpful advice and assistance in the NMR measurements. Fmancial support by the Asahi Glass Foundation is gratefully acknowledged.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TRANSFORMATIO N OF BIOACTIV E GLAS S GRANULE S INT O CA-P SHELL S IN VITR O S. Radin, P. Ducheyne, S. Falaize, A. Hammond Department of Bioengineering, University of Pennsylvania 3320 Smith Walk, Philadelphia, PA 19104
ABSTRAC T Bioactive glass (BG) reactions were modeled in physiological solutions in static or dynamic conditions using either integral (no solution exchange) or differential (with solution exchange at designated time periods) modes of immersion. BG granules (either 300-355 or 200-300 |im) were immersed in tris buffered solution complemented with plasma electrolytes (TE), or with both plasma electrolytes and 10% serum (TES-10). Post-immersion solutions were analyzed for changes in Siconcentrations. Granules were analyzed for compositional, morphological and structural changes resulting from immersion by using scanning electron microscopy (SEM), energy dispersive X-ray (EDX) analysis and Fourier Transform Infrared (FTIR) spectroscopy. The amount of Si released from BG granules in the differential mode of immersion in TE was double the one in the integral mode. The dissolution was further enhanced in serum-containing TES-10: total dissolution of silica was observed after 1 week of differential immersion in TES-10. The granules with totally dissolved silica were transformed into hollow Ca-P shells. KEYWORDS : bioactive glass, in vitro, surface modification, dissolution INTRODUCTIO N A unique phenomenon of excavation of bioactive glass granules of narrow size range (300-355 |im), concomittant differentiation of osteoprogenitor cells, and formation of new bone tissue inside the excavated granules was observed in vivo [1]. Notwithstanding the biological evidence for this resorption to be cell mediated [1], solution mediated effects can also play a role. Before, bioactive glass (BG) reactions were modeled in physiological solutions using static or integral conditions (no fluid flow or solution exchange) [2,3]. However, physiological fluids in the body are in constant circulation. Thus, in this study, we modeled the reactions of BG granules in dynamic conditions using a differential mode of immersion. This immersion methodology represents conditions in which the solution is continuously replenished. In this paper we document that only differentially performed immersion experiments simulates the internal resorption process. MATERIAL S AND METHOD S BG 45S5 granules with a composition 45% Si02, 24.5% Na20, 24,5% CaO and 6% P2O5 (Orthovita, Malvern, PA), were immersed, either integrally or differentially, in tris buffered solution complemented with either plasma electrolytes 45
46
Bioceramics Volume10
(TE) or complemented with both electrolytes and 10% newborn bovine serum (TES-10) at 370c for up to 1 week. Granules of either 300-355 or 200-300 |iim were used. The granules were immersed at a weight-to-solution ratio of 0.5 mg/ml. In the differential mode, the solutions were exchanged at 3, 6, 10, 24 hours and then every day. Post-immersion solutions were analyzed for changes in Siconcentrations using atomic absorption spectrophotometry (AAS, 5100, PerkinElmer, Norwalk, CT). Following immersion, the BG granules were analyzed for compositional and morphological changes using SEM/EDX (JEOL 6400) analyses. Morphological changes were viewed on granules attached to a carbon tape and coated with carbon. Some of the post-immersion granules were fractured under a light stereo microscope to expose the inner surface. Cross sectional analysis was performed on post-immersion granules embedded in epoxy and polished using diamond coated discs. After polishing the cross-sections were coated with carbon. FTIR analysis was also used to determine the structure and composition of the reaction surfaces. RESULT S AND DISCUSSIO N The Si-release from BG granules of 300-355 |Lim (expressed as weight % of the original Si content) vs. immersion time in the integral mode in TE, and the differential mode in TE and TES-10 is shown in Figures 1 a,b. Whereas 34.5% of the original Si-content was dissolved after 1 week of immersion in the integral mode in TE, the cumulative Si-dissolution in TE was equal to 69.6% in the differential mode. The dissolution was further enhanced by differential immersion in TES-10: total dissolution of the silica from granules was observed after 1 week of differential immersion. Similar results (i.e. total silica dissolution after differential immersion in TES-10) were obtained for BG granules of 200-300 ^m.
80- .
Le 40 -
1
20 ^
0 H f 3
-
- TE.inligr«
1^ 1 24
j
X
1 72
--
4
^
168 0
24
Time of immersion (hours )
Figure 1 a,b. Si-release (% of the original Si-content in BG) from BG granules (300-355 \xm)vs. immersion time in integral mode in TE (a) and differential mode in TE and TES-10 (b). Arrow indicates time to formation of a HA surface layer detectable by FTIR.
In Vitro Transformationof Bioactive Glass Granules (300-355 fiMj into Ca-P Shells: S. Radin 47
Figures 2 a, b. SEM micrographs of BG granules after differential immersion in TES-10 for 168 hours: (a) view of a fractured granule after silica dissolution; (b) cross-section of an internally dissolved granule (embedded in epoxy and polished). The SEM micrographs of granules after differential immersion in TES-10 (Fig. 2 a,b) show transformation of the granules into hollow shells. The EDX spectra (Fig. 3) corresponding to three different spots of the cross-section indicate that the hollow particle was mainly composed of a Ca-P phase which contained traces of CI, Na and Mg. The Ca/P ratio in the different spots on the shell varied from 1.1 to 1.3. No significant Si-content could be detected on either outer or inner surfaces of the shell. FTIR spectrum of the granules after differential immersion in TES-10 showed that the Ca-P phase was partly or fully crystalline, as indicated by the sharpening and splitting of the P-O band in the lower energy region (Fig. 4). The crystalline phase could not be identified because of extra absorption bands present in the energy region from 700 to 950 cm-1 due to spectral distorsion associated with a rough surface (i.e. small particles).
keV
Figure 3.Three EDX spectra (the spectra overlap) taken from three different spots on the hollow shell cross-section (shown on Fig. 2b): at the outer surface, at subsurface and on the inner surface.
48
Bioceramics Volume 10
Figure 4. FTIR spectrum of BG granules after differential immersion in TES-10 for 168 hours. Split of the P-0 band in the lower energy region indicates formation of a crystalline Ca-P phase. The observed effect of the mode of immersion and the presence of serum protems on the the degree of silica dissolution can be explained as follows. In the mtegral mode of immersion in serum-free TE silica dissolution was slowed down after 24 hours of immersion due to formation of a HA surface layer (detected by FTIR). The protective effect of the surface HA layer on dissolution of silica-based bioactive glass has been described elsewhere [4]. Our data indicate that the differential mode prompted continuous Si dissolution in TE at the immersion stages preceedmg the HA formation (Fig. la). As we previously reported [5], the presence of serum in TES-10 slows down the HA formation. The surface reaction layer which forms in TES-10 is composed of amorphous Ca-P accumulations in a silica matrix. This layer does not prevent continuous Si-dissolution. As a result all the sihca dissolved from the granules in these experiments. Granules of which all the sihca was dissolved in TES-10, were fully transformed to hollow Ca-P shells. Crystallization of the Ca-P phase occured upon Si-dissolution. CONCLUSION S 1. In vitro immersion in a serum solution using conditions reflecting the continuous fluid flow in vivo, leads to internal dissolution of glass granules. Static immersion experiments fail to achieve this phenomenon. 2. Transformation of bioactive glass granules into hollow Ca-P shells can be achieved in vitro. AKNOWLEDGMEN T This work was supported by grants NSF BSC 93-09053 and NIH DE-10693
REFERENCES
1. Schepers, E., De Clercq, M., Ducheyne, P. and Kempeneers, R., J.Oral. Rehab..,1991, 18, 439-452. 2. Hench, L.L., J. Am. Cer. Soc, 1991, 74(7), 1487-510. 3. Kokubo, T., J. Non-Cryst. Solids, 1990, 120, 138-151. 4. Hench, L.L., J. Non-Cryst. Solids, 1978, 28, 83-105 5. Radin, S., Ducheyne, P., Rothman, B., Conti, A., J. Biomed.Mat. Res., 1997, 35, in press.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
Multilayere d Coating s of Hydroxyapatite/Glas s Cerami c Composite s Plasm a Spraye d on Ti-6A1-4V Alloy P.L.Silva*^* J.D.Santos^*, F.J.Monteiro^* ^ISEP-CEIA-Instituto Superior de Engenharia, Institute Politecnico do Porto, R. de S.Tome, 4200 Porto, Portugal. INEB-Instituto de Engenharia Biomedica, Pra^a Coronel Pacheco 1,4050 Porto, Portugal. ’-’Departamento de Engenharia Metalurgica, FEUP, Universidade do Porto, R. dos Bragas, 4099 Porto Codex, Portugal. ABSTRACT Aiming at obtaining plasma sprayed coatings with enhanced bioactivity, when in contact with host tissues, double layered coatings were prepared. The top layer was composed of a hydroxyapatite (HA)/4% P2O5 glass composite and the undercoat was of pure simple HA onto Ti-6A1-4V alloy. X-Ray Diffraction Analysis (XRD) of the composite top layer showed that it was less crystalline than simple HA layer due to the glass addition, and was composed of a HA matrix with a small amount of p-TCP. Scanning Electronic Microscope (SEM) observations showed that it has been possible to obtain double-layered coatings with good bonding between them and well adherent to the metallic substrate. Coating to substrate adhesion was determined according to three standard methods: ASTM C 633-79, ASTM D 1002 and DIN 50161. Image analysis technique was used to determine the type of failure mechanism that took place and it was found that 85.6%–2.7 was adhesive and 13.4%–2.7 cohesive. The coating to substrate adhesion values are within the range commonly required for this type of plasma sprayed coatings for medical applications. KEYWORDS : Plasma-spraying, ceramic composites, adhesion. INTRODUCTIO N The interest in using HA for biomedical applications is commonly known [1-3]. However, it has been proved that for load bearing applications the mechanical properties of HA, mainly its strength and fracture toughness should be improved [4-5]. In a previous work [6], it has been shown that by applying a P2O5 based glass reinforcement to HA, fracture toughness and biaxial bending strength were enhanced. In this work, HA reinforced with 4%P205 based glass was used as a surface layer, i.e., the material that primarily establishes contact with tissues and organic fluids. In order to adjust glass chemical composition as closely as possible to inorganic part of bone chemical composition [7-9], Na20, MgO and K2O oxides were incorporated into P2O5 based glass. This composite was the top layer of a mixed coating containing as an undercoat a pure HA layer, adherent to the Ti-6Al-4V alloy. One of the main concerns when using plasma sprayed coatings is to obtain adequate coating to substrate adhesion [10]. The way by which a coating adheres to a substrate is very complex and not fully understood [11-13]. Several theories and mechanisms of adhesion have been proposed but there is no single one that might explain all adhesion behavior [11-14]. 49
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P205
35
Table 1 - Glass chemical composition (mol%). CaO Na20 K2O 35 10 10
MgO 10
Due to the difficulty in quantifying this property, three standard methods have been chosen allowing for crossed results evaluation. MATERIAL S AND METHOD S P205-CaO based glass was prepared from reagent grade chemicals and its chemical composition may be seen in Table 1. The glass was wet mixed and milled with HA powder using methanol. 4% glass addition was used. Mixed powders were dried and isostatically pressed at 200MPa, The method used for composite preparation has been fully described elsewhere [6]. Cylindrical shaped samples were sintered at 1300^C and once again milled, using an agate ball mill pot. The powders were then sieved to obtain a grain size distribution suitable for plasma spraying deposition. Two kinds of HA were used for the base coatings: HA-P, supplied by Plasma-Technik with the reference code, AMDRY AM 6021, specifically prepared for plasma spraying and used "as received"; HA-I, supplied by Plasma Biotal and conditioned to be used for plasma spraying. Plasma spraying was performed under atmospheric conditions using a Plasma Technik automated equipment. After grit blasting the Ti-6A1-4V alloy with AI2O3 spheres, and chemical degreasing with trichlorethylene, a 50\im HA coating was sprayed followed by the deposition of SOjxm HA/glass composite coating. Samples were characterized by XRD and compared with an XRD spectrum obtained for a 100|Lim commercial HA-P coating. Ten samples were used for each of the three adhesion experiments. From ASTM D 1002 tests, comparative shear strength values were obtained. With the ASTM C 633-79 it was possible to obtain coating adhesion under stresses normal to the surface, and with ASTM C 633-79 DIN 50161
0 1,5+0,2 cm
l . t cm
ASTM D1002
1,6mn n 1
1 Z.TTnm
sttear
"^.^afcr^ ^
Fig.l - Test specimens according to ASTM D 1002, ASTM C 633-79 and DIN 50160 standards.
Coatings of HA/Glass Ceramic CompositesPlasma Sprayed on Ti-6al-4v Alloy: P.L. Silva et al.
v._^L^U--^X^w ;
51
M- MU AMkli&«ii^.^.v^^^^ ^
a) b) Fig.2-XRD spectra for HA-I+composite samples a) and HA plasma sprayed samples b). DIN 1061 testing compressive stresses were applied. The shape and dimensions of the test specimens may be seen in Fig.l. The glue used to assemble the test specimens in both, ASTM D 1002 and ASTM C 633-79 tests, was Plasmatex Klebbi from Plasma Technik. In order to determine what type of failure mechanism took place, cohesive or adhesive, fracture surfaces were observed by SEM according to image analysis technique. RESULT S AND DISCUSSIO N XRD spectra may be observed in Fig.2. Both coatings had an HA matrix with small amounts of P-TCP. Comparing with the HA + composite samples, the simple commercial HA coating presents higher cristallinity. Adhesion results are presented in Table 2. Slightly higher adhesion values were obtained for the HA-I+composite samples. Both coatings have shown higher resistance to uniaxial tensile stress than to shear stresses, applied in ASTM D 1002 experiments or compressive stresses used in the DIN 60161 standard. Examples of image analysis applied to the samples tested according to ASTM C 633-79 are presented in Fig.3. The areas where some coating remained attached to the substrate are yellow colored and the gray color represents Ti-6A1-4V alloy substrate. Using this technique it was possible to determine that HA-I+composite failure was 85.6 – 2.7% adhesive and 13.4% – 2.7 cohesive. HA-P+composite samples have shown 57.2%–4.6 adhesive and 42.8%–4.6 cohesive failure. In samples subjected to ASTM D 1002 the coating was completely transferred to the counterpart so that image analysis determinations were not performed. Fracture surfaces of samples subjected to DIN 50161 have showed that failure occurred by the coating - cohesive failure. Magnification of the failure area have shown that fracture surfaces were irregular and some failure planes may be observed, which indicates that brittle failure took place. SUMMAR Y In this work an attempt was made to obtain new coatings designed to tailor the needs of bonding to neighbor tissues and whitstanding for a reasonable amount of time the contact with physiological environment. A surface coating capable of inducing a fast response from the host tissue in the early moments of implantation was created, showing a more reduced dissolution kinetics in the long term, thus allowing for the newly formed bone tissue to be fully established Coating HA-I+composite HA-P+composite 1 Glue
ASTM D 1002 (MPa) 14.9–4.3 13.1–3.8 22.7–0.1
ASTM C 633-79 (MPa) 35.4–6.5 33.8–5.6 70.0–0.1
DIN 50161 (MPa) 40.0–9.8 36.1 –8.6 2
1
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Bioceramics Volume10
Fig.3 - Failure surface image analysis photomicrograph of an HA-I+composite sample a) and HAP+composite sample b) and adequately fixed to the implants. As it was clearly detected in Fig.2, the upper layer coating was more amorphous than HA, probably due to the glass incorporation. Results presented in Table 2 show that coating to substrate adhesion was slightly improved when HA-I + composite double layers were used. These results may be explained by the fact that HA powders used to produce these coatings had higher strength and density than those used to prepare the HA-P+composite coatings since they were isostatically pressed and sintered. Both kinds of double layered coatings present higher adhesion values for ASTM C 633-79 - tensile stresses than for ASTM D 1002- shear stresses or DIN 50161 - compressive stresses. Comparing these results with the needs for the implanted materials in load bearing applications, 30MPa, it may be concluded that both double layers seem to entirely satisfy stress requirements. Image analysis, seen in Fig.3 and performed on samples subjected to ASTM C 633-79 tests, have shown that adhesion failure was 85% adhesive for HA-I + composite samples but, for HAP+composite coatings this value decreases to 57%. This behavior may be explained by the fact that when using HA-I+composite samples both HA - the one from the composite layer and the "pure" HA layer - were prepared by the same procedure. HA-P+composite samples were prepared with two different kinds of HA powder, the HA powder from the composite that was isostatically pressed and sintered, and the commercial HA from the bottom layer that was applied "as received". The fact that in the ASTM D1002, the double layer was completely transferred to the counterpart have showed that this was a mainly adhesive failure. The DIN 50161 failure surface observations presented essentially cohesive kind of failure.
Acknowledgment s The authors which to thank INEB for the provision of laboratory facilities, financial support of JNICT trough Ref.PBICT/CTM/1890/95 and Marta Sa and Barbara Silva for their immense collaboration in the execution of the adhesion experiments.
REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14.
RDucheyne , J.BiomedMater.Res.App.Biomater., 21 (1987), 219. H.Aoki, in ScienceandMedical Applications of Hydroxy apatite, Takayam a Pres s System Center , JAAS, Tokyo, (1991), 165. R.Z.Legeros , Adv.Dent.Res.,2, (1988), 164. G.With,H.Va n Dijk, N.Hattu , K.Prijs , J. Mater.Sci., 16, (1981), 1592. M.Akao , H.Aoki, K. Kato , J. Mater.Sci., 16, (1981), 809. J.D. Santos , P.L. Silva, J.C . Knowles, F.J.Monteiro , /. Mater.Sci., Mater.Med.,1, (1996), 187. D. Williams , Medical& Dental Materials, Pergamo n Press , (1986). C.Rey, Biomaterials, 11,(1990), 14. G.Evans , J.Behiri , J.Currey , B.Bonfield, /. Mater.Sci., Mater.Med.,1, (1990), 38. D.Matejka , B.Benko, Plasma Spraying ofMetallic Materials, Joh n Wiley & Sons, (1989). S.D.Brown , ThinSolidFilms,119, (1984), 127. K.L.Mittal , Adhesion Me^LSurement of ThinSolidFilms,Thick Films and Bulk Coatings , ASTM STP 840. R. Dixon, Surface Engineering, 11,(1993), 4. W.Lian , Novel Plasm a Sprayin g Processin g for Enhance d Surfac e Engineering , Matetrial s World , 4,
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
THE BONY REACTIO N TO RAPIDL Y DEGRADABL E GLASS-CERAMIC S BASED ON THE NEW PHAS E Ca2KNa(P04) 2 C. Muller-Mail, G. Berger^, C. Voigt^ B. Bakki^, U. Gross^ Department of Traumatology and Reconstructive Surgery^ and Institute of Pathology^, Universitatsklinikum Benjamin Franklin, Freie Universitat Berlin, Hindenburgdamm 30, 12200 Berlin, Germany, Federal Institute for Materials Research and Testing, Berlin-^, Germany. ABSTRAC T Glass-ceramics in the composition field of CaO-K20-MgO-Na20-P205 can be produced based on the main crystalline phase Ca2KNa(P04)2. The degradation rate depends on the whole composition, i.e. the crystalline phases as well as the residual glass-phase. The degradation rate of the two investigated glass-ceramics was higher as in B-tricalciumphosphate (B-TCP) and was mainly due to passive leaching processes, especially in GB18. Both materials exhibited bonebonding properties and allowed at least for partial guided bone regeneration during degradation. KEYWORDS : Bone/interface/glass-ceramics/degradation/guided bone regeneration/ultrastructure. INTRODUCTIO N Bone defects, for example after trauma, need reconstruction. Such defects can be reconstructed by using autologous bone and homologous bone, both having their drawbacks and limitations. An alternative method uses bioactive bone-bonding implant materials which display different degradation rates. In most of these materials the interface is stabilized by bone-bonding and therefore, the bone bonding material is not further subject to degradation and replacement by bone [1]. On the other hand, the higher degradable materials should allow for guided bone regeneration while being degraded. Thus, the degradation rate should not be too high. The goal of this study was to evaluate two new synthetic implant materials based on a Ca2KNa(P04)2-phase as a glass/glassceramic composite. The dissolution rate of these materials in comparison to 13-TCP is enhanced by exchanging Ca against other ions such as K, Na and by adding Mg [2]. MATERIAL S AND METHOD S Two particulate implant materials were tested with the following constituents according to [2]: GB14 with a content of CaO 30.67 %, MgO 2.45 %, Na20 9.42 %, K2O 14.32 % and P2O5 43.14 % (particle dimentions were measured to be 370 – 70 |nm width and 610 – 100 jam length, n = 45), and GB18 containing CaO 17.72 % and MgO 12.75 %, Na20 9.79 %, K2O 14,88 %, P2O5 44.86 % (380 – 70 |im width, 630 – 100 |Lim length, n = 27). Both compositions are meltable and crystallize spontaneously fi-om the melt. Thus, they can be easily produced. Both particulates consist of a main crystal phase and other amorphous and crystal moieties. The dissolution rate was measured in vitro (0.2 M TRIS-HCl) and yielded rates of 292 – 40 and 1278 – 80 mg/1 for GB14 and GB18, respectively. Therefore, the materials were much more soluble as B-TCP in the same testing system (30,6 – 10 mg/1)[3]. The dissolution rate is controllable by the amount of added ions, such as Na, K, or Mg. Particulates were sterilized by dry heat at 180 C for 30 min.. 100 mg 53
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Bioceramics Volume10
were implanted into each femur below the patella sliding plane into trabecular bone in Chinchilla rabbits after arthrotomy of the knee joint. The defect containing the particulates was closed using a chondrocortical bone slice produced by the hollow cylinder drill after removing the adherent trabecular bone below the cartilage. After implantation, this slice was fixed in the anatomical position by the pressure of the revised patella. Thus, no particles were observed in the knee-joint. A total of 9 animals was operated bilaterally per material and each 3 animals were sacrificed at 7, 28, and 84 days and the femora were prepared for scanning (SEM) and transmission (TEM) electron microscopy as described [4] except that the specimens for SEM-investigation were air dried after rinsing in hexamethyldisilasane. RESULT S AND DISCUSSIO N Specimens prior to implantation In the SEM the GB14-particles showed three different appearences. Two particle types consistend mainly of crystalline spheres attached to each other by a thin film of amorphous matrix. The crystalline spheres had diameters of either 10 – 1 |im (n = 30 measured) or 45 – 6 )Lim (n = 30). Between the crystalline spheres representing the ceramic moiety and the amorphous glass-phase there were some cracks up to 2 |Lim width and more than 20 |im length. At the neck of separate spheres, where cracks contacted each other, roundish pores were created of approximately 3.5 |Lim diameter. The third kind of particles was the most dense, since there were no cracks or pores and crystalline spheres were not detectable. The GB18 particles were also dense and homogeneous with only few deepenings and elevations. There were no apparent substructures (Figure 1). Specimens after implantation All cases showed healing without special events. During the time course between 7 and 84 days of implantation the surface morphology of the three GB 14 particle types changed considerably. The crystalline spheres either of 45 or 10 jam in diameter were degraded, leading in the case of 45 )Lim spheres to a microporosity of their surfaces with pores and cracks with diameters of 1 jiim or 2 \xm width, respectively. Later on, this process led to a destruction and formation of smaller substructures (Figure 2). Thus, irregular porous substructures of some |Lim in length were produced indicating a rapid degradation of the spheres. The 10 \im crystalline spheres decreased contineously in size by keeping their spherical outline. In the third kind of GB14 particles the amorphous moiety was leached at already 7 days liberating roundish crystalline substructures with mean diameters of 13 – 4 |Lim (n = 10). Between the amorphous moiety and the spheres cracks developed which increased in width at longer time intervals (Figure 3). Already at 7 days pores of about 0.5 fim developed in the spheres. The pore diameters grew with time leading to a complete destruction in some cases. In vitro experiments have shown, that this degradation process led to high concentrations of Mg and P in the medium as well as to precipitations rich in these ions on the substrate. The osteoblast growth was slightly inhibited as compared to Thermanox^[5]. In the GB18 particles at 7 days the ceramic moiety consisting of rod-like structures of up to 20 fim length was liberated, possibly by leaching of the surrounding amorphous glass moiety. Later on, these rod-like structures were also degraded as in the GB14 particles and smaller irregular substructures were formed (Figure 2). At 7 days there was almost exclusively fibrous material surrounding the different particle types. At 28 days bone was observed in contact to the GB14 particles. The bone contacted the dense particles and the particles with 45 \xm crystalline spheres directly, whereas the particles with small 10 \xmspheres were separated from the bone by clefts of about 1 |um width. The different tissue
Bony Reaction to Degradable Glass-Ceramics Based on the Phase Ca2KNa(P04)2: C. Muller-Mai et al.
55
reactions indicate that the chemical composition of the different moieties within single particles as well as in particle types was varying. In general, GB18 developed more bone contact as GB14, probably related to the higher reactivity. There was bone at 28 days in contact to the particles and at 84 days there was almost complete bone contact of the GB18 particles. In the TEM, all particulates exhibited crystalline and amorphous parts. There seemed to be empty spaces between both phases suggesting some microporosity prior to implantation. At each time interval there was soft tissue with fibroblasts and macrophages as well as bone in the interface with increasing bone amounts at longer time intervals especially in contact to GB18 and the dense GB14 (Figure 3). Mineralization occured in the cracks between spheres. Both particulates were degraded with a higher rate of the GB18. In GB14 this was due to a destruction of the spheres as well as to leaching of the amorphous moiety. In GB18 there was first leaching of the amorphous phase leading to the liberation of rod-like structures which degraded also later on. REFERENCE S 1. Muller-Mai, CM., Voigt, C , Baier, R.E. and Gross, U. Cells & Mater. 1992, 4, 309-327. 2. Berger, G., Gildenhaar, R. and Ploska, U. In: Bioceramics Volume 8, Wilson, J., Hench, L.L., Greenspan, D. (eds), Pergamon, Oxford, UK, 1995, 453-456. 3. Berger, G., Gildenhaar, R., Gross, U., Knabe, C, Loginow-Spitzer, A., Miiller-Mai, C , Ploska, U. and Radlanski, R. In: Werkstoffe fiir die Medizintechnik, Symposium 4, Breme, J. (ed), DGM-Informationsgesellschaft Verlag, Frankfurt, Germany, 1996, 59-65. 4. Miiller-Mai, C , Voigt, C. and Gross, U. Scanning Microsc. 1990, 4, 613-624. 5. Knabe, C , Gildenhaar, R., Berger, G., Ostapowicz, W., Fitzner, R., Radlanski, R., Gross, U. and Siebert, G.K. In: Transactions, Vol. I, 5th World Biomat. Congress, International Liaison Committee (ed), University of Toronto Press, Toronto, Canada, 1990, 890.
Figure 1: Surfaces of the particle types prior to implantation. A: GB14 with 45 |Lim spheres, cracks and a pore; B: GB14 with 10 |Lim spheres; C: Dense GB14; D: GB18. SEM, bar 20 jim.
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Figure 2: A: Spheres (45 jim-type) at 7 days with newly developed pores and cracks. B: GB18 at 84 days surrounded by bone with rod-like crystalline moiety. SEM, bars A:50 |Lim, B: 100 |Lim.
Figure 3: A: Dense GB14 at 84 days with cracks between amorphous and crystalline moiety with pores in separate spheres surrounded by bone. B: Bone bonding to ceramic moiety of GB18 at 84 days (grey, white areas as artefacts due to loss of glass-ceramic). TEM, bars A: 100 |im, B: 1 jim.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
RESORBABL E POROU S PHOSPHAT E INVER T GLASSE S - FIRS T IN VITR O AND IN VIVO RESULT S J. Voger, K.-J. Schulze^ D. Reif, P. Hartmann^ U. Platzbecker^ B. Leuner^ ’Otto-Schott-Institut, Friedrich-Schiller-Universitat, FraunhoferstraBe 6, 07743 Jena, Germany, ^Klinik fiir Orthopadie, Technische Universitat, FetscherstraBe74, 01370 Dresden Germany, ^Biovision GmbH, Am Vogelherd 1, 98693 Ilmenau, Germany, ’’Institut fiir Optik und Quantenelektronik, Friedrich-Schiller-Universitat, Max-Wien-Platz 1, 07743 Jena, Germany
ABSTRAC T A resorbable phosphate invert glass was investigated in view of its solubility behavior "in vitro". First animal experiments were carried out using a porous shape of the phosphate invert glass and of the corresponding glass ceramic. Both the solubility tests and the animal experiments confirm an excellent biocompatibility. The living bone grows into the pores of the implants and the materials are reabsorbed in course of time. The resorption rates meet the expectations. KEYWORD S Phosphate glasses, resorbable, porous, solubility tests, animal experiments INTRODUCTIO N The regeneration of bony defects can be supported by resorbable implants. Sintered tricalcium phosphate is mainly used for these clinical applications today. Pure dense sintered TCP shows a good biocompatibility [1,2]. However, in dependence on size and place of the filled bony defect, the resorption term of this material can reach up to some years and rests of the material can cause biomechanical problems. Additionally, the different solubilitys of the crystalline phase and the amorphous grain bounderys can result in a disintegration of the material in grainy particles. In these cases, a biocompatible resorbable homogeneous material possessing a clearly increased resorption rate is desirable. Some calcium phosphate invert glasses are suitable for the development of resorba› ble implants showing well defined resorption rates [3,4]. Because these glasses don’t show phase separation, they represent single phase systems with an uniform solubility. Using a salt-sintering process, the glasses can also be produced in a porous shape with a texture similar to that of spongious bone. A selected glass of this group and the corresponding glass ceramic were used for animal experiments and solubility tests. 57
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MATERIAL S AND METHOD S The composition of the investigated glass/glass ceramic was determined by usual wet chemical methods: 32.6 mol% P2O5, 27.6 mol% CaO, 27.6 mol% Na2 0, 12.2 mol% MgO. Mixed sodi› um/calcium diphosphates and a B-TCP phase, stabilized by magnesium, are the main crystal phases of the glass ceramic. ^^P MAS NMR experiments confirmed the invert glass structure of the materials. They contain not or low condensed phosphate groups exclusively. In their porous form both materials show an total open porosity of 65%. The diameters of the pores vary between 150|Lim and 400^m (figure 1). Smaller pores in the walls between the main pores are caused by the sintering process and range from about l|im and 60^m (15%). Informations about the solubility behavior in dependence on time were obtained using a simple soaking test: Two grams of the substances (grain size 315|Lim to 400 |im) in 200 ml destilled water were shaked by a shaking machine at 37 C for 24 hours. In the eluate, the pH-value and the quantity of dissolved ions were determined. The soaked glass grains were washed and the same procedure was repeated eight times. For the animal experiments guinea pigs were used. Cubes (2x2x2 mm^) of the porous materials were implanted into the tibiae of 72 animals, both sides. After 2, 4, 8, 16, 32 and 64 weeks bone speci› mens containing the rests of the implants were resected. Besides the histological investigation, also histomorphometrical methodes were used. The determination of the total area of the implants by a light microscope combined with a drawing unit and a PC gave informations about the resorption rate in vivo. RESULT S AND DISCUSSIO N Table 1 gives the total quantity of dissolved ions (E Ca^^Mg^^,Na^,P04 ^in mg/1) in dependence of soaking period. The values of the glass are five to seven times higher than the values of TCP. Because of the strongly increased surface of the porous glass also the quantity of dissolved ions is
Figure 1. Resorbable porous phosphate invert glass with a total open porosity of 65%.
Resorbable,Porous PhosphateInvert Glasses: J. Vogel et al.
59
Table 1. Total quantity of dissolved ions (E Ca^^,Mg^^,Na’^,P04 ^’in mg/1) in dependence on soaking period eluatel
eluate2
eluateS
eluate4
eluateS
eluate6
eluate7
eluate8
glass
115.5
126.3
117.3
85.6
89.1
84.8
83.8
78.5
porous glass
290.1
190.1
144.8
120.0
112.6
153.5
160.7
159.0
TCP
16.0
17.8
13.8
15.8
16.5
increased. The reason for the minimum at the fourth and fifth period is unknown in this time. Possibly, an unstable intermediate layer is formed because of the higher reactivity of the porous material. The delayed resorption in vivo some weeks after implantation (see below) could be caused by such a layer. However, the amounts of the several dissolved ions never exceed the physiological tolerance. After an initial period, the values of both glass and porous glass are nearly constant. This is also valide for the pH-values of the eluates (figure 2). Similar to TCP the pH-values come close to the physiological pH-value. Therefore, cell toxicological problems are not to be expected. The lower pH-values of the porous glass in the first eluates is caused by the production process of the porous texture. The histological investigations of both glass and glass ceramic didn’t show any symptoms of inflammation. Within the first two weeks post operation the formation of thick osteoid beams was observed. In the following time the pores act as a guide rail for the young bony cells growing in and the implants are completely incorporated by osteoid after three to four month.
10,Q|
physiologica l pH TCP porous phosphat e glas s phosphat e glas s
9,59.0 i 8.5-
X
Q.
8,07,5 ^
7.oH
,,>^J ;
6,56,0 -
4
5
6
eluate number
Figure 2. pH-values of eluates in dependence on soaking period
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Bioceramics Volume10
10Oi 4
V
\
^
80-
^ ^ ?
60-
^
40-
A
0\ \ ’^^
20-
()
glass cerami c phosphat e glass j
A>.....___ ^
1
10
20
30
1
40
1
50
1
60
weeks
Figure 3. Total area of resorbable implants in vivo in dependence on time From that time on the osteoblasts of the osteoid are changed to osteocytes. After 64 weeks the spongiosa framework represents a mixture of ripe bone, osteoid and small amounts of incorporated glass/glass ceramic particles. Simultaneous to the formation of bone the porous phosphate glass and glass ceramic are reabsorbed. Figure 3 gives the course of the total area of the implants in dependen› ce on time. 64 weeks post operation only small rests of the materials are detectable (glass: around 2%, glass ceramic: around 10 %). A mechanical instability caused by the dissolution of the implants was not observed. Obvoisly, it is compensated by the increasing bony integration of the implants. CONCLUSION S Both the phosphate invert glass and the glass ceramic meet the requirements of resorbable implant materials. Simultaneous to the resorption process the implants are bony integrated. The resorption rate of the materials is adjusted to the growth and the mineralization of bone. REFERENCE S 1.
Koster, K., Heide, H., Konig, R., Langenbecks Arch. Chir. 343 (1977), 174
2.
Klein, C , Driessen, A., de Groot, K., van den Hooff, A., J. Biomed. Mater. Res. 17(1983), 769784
3.
Hartmann, P., Vogel, J., Schnabel, B., J. Non-Cryst. Solids, 176 (1994), 157-163
4.
Vogel, J., Hartmann, P., Schulze, K.-J. In: Advances in Science and Technology 12: Materials in Clinical Applications, Techna Sri., Faenza 1995, 59-66
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
IMPLANTATIO N OF BIOACTIV E AND INER T GLAS S FIBRE S IN RAT S TISSU E RESPONS E AND SHORT-TER M REACTION S OF THE GLAS S
SOFT
M. Brink\ P. Laine^ K. Narva^ and A. Yli-Urpo^ ^Department of Chemical Engineering, Abo Akademi University, Biskopsgatan 8, FIN20500 Abo/Turku, Finland ^Institute of Dentistry, University of Turku, Abo/Turku, Finland ABSTRAC T The purpose of this work was to develop a bioactive glass fibre that resorbs in soft tissue without causing inflammatory reactions. In addition, the glass should bond to bone and be easily manufactured. Two different biocompatible glasses were chosen for implantation, and glass surface reactions as well as soft tissue response were evaluated. An inert commercial glass fibre was used as reference. After implantation, all glasses were in good contact with the surrounding tissue. The biocompatible glasses were severely resorbed after 28 days in soft tissue indicating that these glasses are suitable for membranes in orthopaedic and maxillofacial surgery, and for reinforcement of resorbable biopolymers. The reference glass fibre did not show any signs of reaction. KEYWORD S glass fibre, glass reactions, resorption, soft tissue, tissue response INTRODUCTIO N Resorbable bioactive glass fibres may be used as membranes for tissue guiding and as carriers for growth factors in orthopaedic and maxillofacial surgery. The glass fibres may also be used for reinforcement of resorbable biopolymers. For these applications, the fibres should preferably resorb within weeks after implantation without causing any inflammatory reactions. In addition, the fibres must be easily manufactured without risk for devitrification (crystallisation) of the glass melt or fibres. Pazzaglia et al. [1] have developed bioactive glass fibres for substrates for bone apposition, but implanted into rat soft tissue, an intensive inflammation reaction occured. Other bioactive glass fibres, but intended for bone implants, have been presented by Vita Finzi Zalman et al. [2] while Graves and Kumar [3] have developed a bioabsorbable glass fibre for reinforcement of bioabsorbable polymers. The latter fibres are based on the system CaO-P205.
Two different glasses, glass 20-92 and 13-93, were selected for implantation into rat soft tissue. Glass 20-92 is biocompatible but since it contains only 50 wt % SiOi, it has a narrow working range [4,5]. Glass 13-93, containing 53 wt % Si02, is bioactive and has a large working range [4,5]. An inert commercial glass fibre, the E-glass, was used as reference material. After implantation in rat soft tissue, initial glass reactions as well as soft tissue response were evaluated. 61
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MATERIAL S AN D METHOD S The denotations and compositions of the fibres used are given in Table 1. The glasses 20-92 and 13-93 were melted for 1.5 h at 1360’’C in a ceramic crucible (Hackman-Arabia, Finland). The batch size was about 300 g, and all raw materials except the sand were of analytical grade. Glass fibres were obtained by pouring the glass melt onto a rotating plate (about 30 cm in diameter) of stainless steel. The thickness of the fibres could be controlled by changing the spinning velocity. Fibres of glass 20-92 are shown in Figure 1. The E-glass fibres were obtained from Ahlstrom Glass Fibres, Karhula, Finland. The sizing was removed by heating for 40 minutes at 650^C. Prior to implantation, all glass fibres were cut into a length of 5-7 mm, rinsed with ethanol and sterilised in hot air. Two or three different glass fibres were implanted subcutaneously into the soft tissue of 18 Long Evans male rats weighting 290-440 g. The rats were anaesthetised with 0.6-1.0 ml Hypnormfi/Dormicumfi. After the implantation times of 7, 14 and 28 days, the rats were killed with CO2. Tissue samples were fixed in 70% alcohol and embedded into plastics. Histological sections were made using a cutting-grinding technique and stained with toluidine blue. Tissue reactions were analysed with light microscopy and glass surface reactions by scanning electron microscopy (SEM) and energy dispersive X-ray analysis (EDXA). RESULT S AN D DISCUSSION After 7 days, no inflammatory cells were detected around glass 13-93 and 20-92. In general, the number of inflammatory cells around the glass fibres was minimal after all implantation times, and most of the surrounding tissue was inflammation-free. The number of inflammatory cells around E-glass was minimal after 7 days. All glass fibres were in good contact with surrounding tissue, and connective tissue grew in tight contact with the glass surfaces. Figure 2 and 3 present glass 20-92 and 13-93, respectively, after 7 days in rat soft tissue. For fibres of glass 20-92 and 13-93, it was found that the resorpfion had started already after 7 days in soft tissue. The resorption was detected as a silica rich layer at the fibre surface, with sporadic formation of calcium phosphate on top. This result is in accordance with previous studies on rods of glass 13-93 in soft dssue [6]. For glass 20-92, with a durability significantly lower than that of glass 13-93 [7], only a core of original glass surrounded by a silica layer was left after 7 days in vivo. The resorption of this glass was thus more pronounced than that for glass 13-93. However, the surrounding tissue did not show any signs of inflammation. The E-glass showed neither any inflammatory reaction in soft tissue, nor did it resorb. E-glass after 14 days in rat soft tissue is shown in Figure 4. After 28 days in rat soft dssue, fibres of glasses 20-92 and 13-93 were severely resorbed. Glass surface reactions after implantation are presented in Table 2.
Table 1. Denotations and composifions (in wt %) of the investigated glass fibres, and fibre diameter. Fibre 0 Glass NaiO K2O MgO CaO B2O3 AI2O3 P2O5 Si02 70-300 ^m 20-92 15 15 2 15 3 50 70-300 \im 13-93 6 12 5 20 4 53 10 |Lim** E-glass 1* 0.7 22.5 6.4 14.7 54 *Na20+K20 **The E-glass was implanted as a bunch of fibres.
Implantationof Bioactive and Inert Glass Fibres in Rats: M. Brink et al.
63
Table 2. Glass surface reactions after implantation into rat soft tissue for several observation times. (Ca,P = calcium phosphate) Glass 7 days 14 days 28 days 20-92 Silica gel with Ca,P Silica gel with Ca,P Silica gel with Ca,P 13-93 Silica gel with Ca,P Silica gel with Ca,P Silica gel with Ca,P No reaction No reaction No reaction E-glass* *without sizing CONCLUSION S This study indicates that glasses 20-92 and 13-93 are suitable for biomedical use as resorbable materials. Which of these two is the more suitable depends on the application. It was shown that fibres of these two glasses resorbed while the E-glass fibres did not. The resorption was initiated almost immediately but no inflammation reaction was detected. ACKNOWLEDGEMENT S This work was financially supported by the Finnish Technology Development Centre (TEKES) and the Academy of Finland (FA).
K-i^
Figure 1. Fibres of glass 20-92.
Figure 2. Glass 20-92 after 7 days in rat soft tissue, (magnification 125x)
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Figure 3. Glass 13-93 after 7 days in rat soft tissue, (magnification 125x)
Figure 4. E-glass after 14 days in rat soft tissue, (magnification 125x)
REFERENCES 1. Pazzaglia U.E., Gabbi, C , Locardi, B., Di Nucci, A., Zatti, G. and Cherubino, P. J Biomed.Mater.Res. 1989, 23, 1289-1297. 2. Vita Finzi Zalman, E., Locardi, B., Gabbi, C. and Tranquilli Leali, P. WO 91/12032 i. Graves (Jr.), G.A. and Kumar, B. United States Patent 4,604,097. 4. Brink, M. /. Biomed.Mater.Res. (to appear). ^’ appe^)^" ^"’’"’^"’ ^ ’ "’’PP’"’^’’’ ^-^- ^ ^ Yli-Urpo, A. J. Biomed.Mater.Res. (to
’
TTL^r.’^It
’
" ’ ^’’’""’^"^ ^- ’ " ’’^’’
^’’’’
BiomaterialsCongress,
7. Brink, M., Karlsson, K.H. and Yli-Urpo, A. WO 96/21628 (pending).
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedingsof the10th InternationalSymposiumon Ceramics in Medicine, Paris, France, October1997) '1997 Elsevier Science Ltd
Spinal Fusion using Titanium Spacers with Bioglassfi and Autogenous Bone: A Comparative study in Sheep June Wilson\ Gary Lowery^, Stephen Courtney^ 1. Imperial College, Dept. of Materials, Prince Consort Rd. London SW7 2BP Tel/Fax 44(0)171 594 6745 2. Research Institute International, 6400 W. Newberry Road, Ste. 206, Gainesville, Florida 32605-4391 USA
ABSTRAC T Spinal fusion, or the surgical accomplishment of a bony union of deficient vertebral segments, is required in cases where long-term immobilization of these segments is indicated. In the anterior lumbar spine, this procedure requires the use of an intervertebral graft construct which is able to support the axial forces until confluent healing of the graft material occurs. In this study titanium surgical mesh spacers were filled with autologous bone, Bioglassfi, or a mixture of the two in the adult sheep spine. After three and six months the new bone was assessed both qualitatively and quantitatively and results show that although at three months autologous bone is to be preferred, by six months there is no significant difference between the three graft materials. This is particularly important since provision of sufficient autologous bone is always difficult and is associated with a high morbidity rate.
INTRODUCTIO N Spinal fusion is indicated in a variety of disorders of the lumbar spine. The goal of this surgery is to achieve bony immobilization of the affected levels. Anterior interbody fusions, usually relying on the use of autologous bone, are indicated in many cases of discal deficiency [1]. AAV ceramic monoliths [2] and titanium mesh spacers packed with autologous cancellous bone [3] have been used in this procedure. Availability and morbidity issues limit the use of harvested autologous bone [4], while the use of allograft involves inherent concerns regarding disease transmission. The ideal bone graft supplement to autologous bone should be both osteoinductive and osteoconductive, allowing bone to grow in a confluent manner, and resorbable, allowing complete replacement by new bone and restoration of near-normal host conditions. The mixing of bioactive glass particles with autologous bone is expected to provide an osteoinductive and osteoconductive material in which the osteogenic proteins derived from the host bone act with the bioactive material to promote rapid bone growth throughout the spacer used for stabilization, which must eventually depend n bony union rather than metallic fixation.
MATERIAL S Titanium surgical mesh intervertebral spacers (Depuy, Inc., Warsaw, Indiana) Particulate 45S5 Bioglassfi 100-410|im diameter. (USBiomaterials Corporation, Alachua, Florida) Autologous bone chips harvested at surgery
METHOD S Cages were placed in four adjacent vertebral (LI -L4) fenestrations created in adult sheep. Six animals were studied after three months and five after six months. The death of one animal in the second group was unrelated to the experiment. Of the four cages, one was filled with the autologous bone chips, one with Bioglass particulate and one with an approximately 50-50 mixture of the two, before insertion in the defect. The fourth cage was left unfilled as a control. The allocation of cages to sites was rotated through the animals. The cage is shown in fig. 1. In the first two animals the larger cage was used, in the remainder the cage was 10mm in diameter. Postoperativecare: Sutures were removed after 7 days. Postoperatively, the animals were monitored for signs of pain or distress until fully recovered (7-10 days). Intramuscular anti-inflammatory pain medication was administered as needed for pain. Food and water intake was monitored during the experimental period. 65
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Fig. 1. Titanium mesh cage At autopsy the cages and surrounding tissues were removed and fixed in formalin. They were embedded in plastic and sawn sections cut transversely. The sections were stained with Paragon and Sanderson’s stains and the nature of the bone assessed. Sections with a complete and well-oriented cage cross-section were used to measure the amount of bone-fill within the cage using an Olympus Image Analysis system. The measurements were made by taking the total cross-sectional area of the cage, measured around the outer circumference and subtracting the area of the titanium mesh to give the area available for infill. The area which was not bone was then identified and subtracted from the area available and the bone infill was expressed as a percentage of the area available. In this way we hope to minimize any differences resulting from the use of 14mm cages in the first two animals and 10mm cages in the rest.
RESULT S The animals recovered well from the surgery and regained their mobility. After three months the bone infill was seen to contain particles of the Bioglassfi with normal bone (fig. 2). Space not filled by bone contained variable amounts of fibrous tissue, but no significant inflammation. After six months the appearance of the bone was similar but space not filled by bone contained little fibrous tissue and seemed to be empty.
Fig 2. Trabecular bone containing Bioglass after six months {Sanderson stain
Spinal Fusion Using TitaniumSpacers With Bioglass^and AutogenousBone: J. Wilson et al. 67 Table 1 - Bone infill after 3 months Cage alone: Bioglassfi alone: Bone alone: Mixture
46.2% 56.7% 93.3% 64.2%
(20.7%-58.1%) (23.5% - 100%) (85.4% - 100%) (57.6% - 78%)
n=5 n=5 n=4 n=5
Table 2 - Bone infill after 6 months Cage alone: Bioglassfi alone: Bone alone: Mixture
47% 82.3% 98% 81.7%
(32.4% - 83%) (67.6% - 90.4%) (94.1%-100%) (69.9%-88.1%)
n=4 n=3 n=3 n=4
After three months the percentage of bone infill was as shown in Table 1. At six months however differences between test groups were less, Table 2. There was considerable variation between animals and thus a spread in the measurements but it is clear that autologous bone remains the best choice in the shorter term, although under optimal conditions complete bone fill was achieved using Bioglassfi alone in one animal. Between 3 and 6 months no change occurred in the amount of bone in the control cages and little change was possible in those filled with bone, since the fill was so high by 3 months. The infill in those filled with Bioglassfi alone and the particulate had increased to where there was no significant difference between them. The results are shown in fig. 3. Students t-test showed that there are no statistical differences between the test groups at six months.
10 0
tt
40
Cage Alone D Bioglass(R ) Alone [H Bone Alone 3 Months
(A)
Fig. 3. (A) Bone fill at 3m. (B) at 6m.
6 Months (B)
Mixture
68 Bioceramics Volume 10
CONCLUSION S We believe that we have shown in this experiment that particulate 45S5 Bioglassfi can be used to replace or dilute autologous bone used to assist in bony repair, thus reducing the immobility associated with harvesting of such bone. It appears also from these results that the contribution of autologous bone and associated osteoinductive factors may be effective even when only small amounts, such as those derived from the bleeding subchondral endplates and generated during surgery, are mixed with the Bioglassfi. The regeneration is slower than that achieved with autologous bone but the clinical advantage lies in the reduction to a minimum of the need for harvesting of bone from the patient.
POS T SCRIP T In November 1996 an 11 month-old bulldog puppy was brought to the University of Florida’s Veterinary Medical Teaching Hospital suddenly paralyzed as a result of a congenital spinal disorder, hemivertebra. She was successfully treated by stabilizing the spine with titanium mesh filled with particulate Bioglassfi and is now completely recovered and mobile. The surgeon. Dr. Roger Clemmons, explained that there is normally no treatment for this condition other than euthanasia.
ACKNOWLEDGEMENT
S
The authors thank Dr. Y. Fujishura of Tahaka University, Japan for the image analysis. The work was supported by USBiomaterials Corporation, Alachua, Florida.
REFERENCE S
1. Fraser RD, Spine,1995 20 (suppl.) pp. 167-177S. 2. Yamamuro T; Bioceramics8, 1995 Pergamon Press, Oxford England., pp.123-127. 3. Lowery GL and Harims J., Manual of InternalFixationof Spine,Raven Press, 1996 Lippencott-Raven Publishers, Philadelphia, PA, pp. 127-146. 4. Fernyhough JC, Schimandle JH, Weigel MC, Edwards CC, Levine AM, Spine,1992 17 pp. 1474-1480.
DENSE AND POROUS BIOACTIVE CERAMICS
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MACROPOROU S DIPHASI C CALCIU M PHOSPHAT E CERAMICS : INFLUENC E OF MACROPOR E DL\METE R AND MACROPOROSIT Y PERCENTAG E ON BON E INGROWT H 1,2
O. Gauthier, ^J-M. Bouler, ^’^E. Aguado, ^P. Filet and ^G. Daculsi
^ Laboratoire de Chimrgie, Ecole Nationale Veterinaire de Nantes, Route de Gachet, BP 40706, 44307 Nantes cedex 03, France, ^ Centre de Recherche sur les materiaux d’interet biologique, Faculte de Chimrgie Dentaire, 1 place Alexis Ricordeau, 44042 Nantes Cedex 01, France ABSTRAC T A total of 60 cylindrical 6 x 6 mm samples of a macroporous biphasic calcium phosphate (MBCP) ceramic was implanted into a distal femoral site in 30 rabbits. These samples represented 6 kinds of implants with 2 different macropore diameters and 3 different macroporosity percentages. Eight weeks after implantation, analysis of backscattered electron images of implant surfaces analysed by a factorial design method showed that implants with 565 ^im pore size provided more abundant newly-formed bone both in peripheral and deep pores than those with 300 jam pore size. No significant differences were found between implants with 40% and 50% macroporosity, suggesting that the influence of macropore size on bone ingrowth was greater than that of macroporosity percentage. MBCP implants with 565 ^im pore diameter and 40% macroporosity represented the optimal association for homogeneous and abundant bone ingrowth. KEYWORD S : bone substitute, ceramic, calcium phosphate, porosity INTRODUCTIO N Macroporosity is conducive to osteoconduction of BCP ceramics but also has many effects on their mechanical behaviour [1, 2]. Cell colonisation and bone ingrowth apparently occur when macropore size is greater than 100 |.im [3], and a reduction in macroporosity may have negative results for the biological properties of macroporous biphasic calcium phosphate (MBCP) ceramics, in that optimal macroporosity parameters have not yet been defined. The purpose of this study was to evaluate the influence of macroporosity on the osteoconduction of BCP ceramics. Bone ingrowth was quantified in several kinds of MBCP implants to determine the most desirable pore size and porosity percentage for osteoconduction. MATER[AL S AND METHOD S We used a MBCP ceramic with a 60/40 HA/pTCP weight ratio. Two main parameters were tested : macropore size (Fl) and macroporosity percentage (F2). Two different macropore diameters, 300 and 565 jam, and two different macroporosity percentages, 40% and 50%, were studied. These four values defined an experimental domain that was studied with a factorial design method (FDM), based on a first-order polynomial mathematical model, to analyse the quantitative resuhs and determine the influence of each factor on bone ingrowth [4, 5]. Ten samples each of 4 kinds of MBCP implants with different macroporosity levels (table 1) were prepared for purposes of statistical evaluation (I = 300 jam and 40% , II = 565 \xmand 40%, III = 300 |Lim and 50%, IV = 565 |im and 50%). 71
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Table 1. Experimental matrix
1
^
1
I II III IV
Fl : macropore diameter (jLim) -1 +1 -1 +1
F2 : macroporosity percentage (%) -1 -1 +1 +1
level -1 level +1
300 –33.3 565121.7
40+1.8 50 –2.0
11-2 : interaction between Fl andF2 +1 -1 -1 +1
Ten samples of 2 other kinds of implants (V = 50% and 565 jam, VI= 30% and 300 \xm)were used on ten other rabbits to compare the experimental data for newly-formed bone with the values calculated with the FDM. Ceramic implantations (randomised distribution) were performed on 30 New Zealand white rabbits. A cylindrical defect was created at the distal end of rabbit femurs at the epiphysometaphyseal junction. A MBCP implant 6 mm long and 6 mm in diameter was positioned to fill the defect. All rabbits were killed 8 weeks after implantation. Femoral extremities were excised, fixed in glutaraldehyde solution, dehydrated in graded ethanol and embedded in glycolmethylmethacrylate. Sections of the femur from each group were analysed by undecalcified histological examination. For each sample, serial sections were cut perpendicular to the long axis of the implant. Qualitative observations were performed in light microscopy on solochrome-cyanine stained sections and with polarised light on unstained ones. The block was then sputtered with Au-Pd for scanning electron microscopy (SEM) observations. Quantitative evaluation was performed by image analysis of the SEM observations of implant surfaces using backscattered electrons (BSE). The whole surface of implants was divided in 12 contiguous fields and recorded on SEM with magnification X50. Threshold was determined by the operator on image analyser and the newlyformed bone surfaces were then automatically calculated and expressed as the percentage of the whole surface. RESULTS All implants showed extensive osteoconduction. In light microscopy, most peripheral pores were completely filled with well-mineralised lamellar bone. This new bone often showed a haversian structure. Only implants II and IV showed evidence of bone colonisation in deep pores where lamellar bone was found on the surface of almost every macropore. Measurements from SEM and image analysis observations based on graylevel distribution allowed to calculate a mean percentage of newly-formed bone for the samples of each kind of implant (table 2). According to the experimental matrix of the FDM, the influences of each factor and the interaction between Fl and F2 (I1.2) were determined for implants I, II, III and IV. The equation describing the percentage of newly-formed bone in implants can be formulated as follows : (1) newly-formed bone % = CM% + Si(fi.Fi), where CM% is the calculated mean for implants I, II, III and IV, and li(fi.Fi) is a first-order polynomial fiinction depending on significant influences and interaction between factors Fl and F2.
Macroporous Biphasic Calcium Phosphate Ceramics: O. Gauthieret al.
73
Table 2. Newly-formed bone percentages for the different kinds of MBCP implants. Implants
I
U
II I
IV
newly-formed bone % Implants newly-formed
16.7 – 3 . 9 1
20.6 – 5 . 5 0
16.8 – 3 . 0 7
22.0 – 6 . 9 3
V 22.0 – 5 . 3 0
VT 8.7 –3.28
Calculated mean (CM ) 19.0
bone % Using values from the experimental matrix, the equation became : (2) newly-formed bone % = C M % + (2.28 x F l ) + (0.39 x F2) + (0.35 x F l F2) Only the influence of macropore diameter seemed to have a significant impact on bone ingrowth. The equation can thus be simplified : (3) newly-formed bone % = C M % + 2.28 X I , XI was a coded value related to macropore diameter D by XI = (D - 432.5)/l32.5 The equation then became : (4) newly-formed bone % = 11.6 + (0.017 D) For implants V, F D M gave a percentage of newly-formed bone of 21.2% – 5.07, whereas the experimental results for 10 MBCP samples of implant V gave a percentage of newly-formed bone of 22.0% – 5 . 3 0 These two values were not significantly different. The experimental results for implants VI were not predicted by the FDM (table 2). DISCUSSIO N A FDM mathematical model was used to investigate in vivo mechanisms relative to the influence of macropore size and macroporosity percentage on bone ingrowth. Our precise experimental conditions showed that newly-formed bone in MBCP ceramics can be regarded as a simple and linear function of macropore size (11.6 + 0.017 D). FDM could not account for the behaviour of VI implants whose macroporosity parameters were chosen outside of the experimental area. This indicates that interpretation of FDM data cannot be extended beyond the experimental limits defined by implants I, II, III and IV but can predict bone ingrowth if macroporosity parameters are still chosen inside the experimental domain. Our in vivo conditions led to great variability. However, our work concerned a large number of BCP samples and provided very precise quantitative evaluation using SEM with BSE [6]. This original image analysis method based on the recording of contiguous images seems to be applicable to the study of biomaterials with good reproducibilty. Macroporosity confers osteoconductive properties on bone substitutes and it is generally admitted that 80-100 )Lim is the minimal pore size for osteoconduction. In our study, after 8 weeks of implantation, better bone ingrowth was achieved for macropores of 565 than 300 |nm. Newly-formed bone was not only significantly more abundant in BCP implants with 565 fim macropores but was also observed in both peripheral and central macropores. The BCP implants with 30% macroporosity and 300 |im pore diameter gave a very low rate of newlyformed bone. Their macropore size and macroporosity percentage were both inadequate with our
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model. In fact, the FDM and experimental results indicated that macroporosity percentage was a less important influence than macropore size. For a similar macropore size, there was no significant difference in newly-formed bone for implants with 40% or 50% macroporosity. Our results would appear to have important mechanical implications. The mechanical properties of BCP ceramics improve in vivo due to bone ingrowth in macropores and reprecipitation of biological apatites in micropores [7, 8, 9]. In our study, a macropore size of 565 jLim provided a higher rate of newly-formed bone than one of 300 \xm.The presence of this new bone over the entire implant surface can have a favourable influence on BCP mechanical behaviour. The more macroporous implants are before implantation, the less mechanical resistance they offer [10] but it has been demonstrated that macroporosity percentage has a greater influence than macropore diameter on the compressive strength of BCP implants [5]. From this study, we could consider that implants with 565 j.im macropore size and 40% macroporosity could have a 67% higher compressive strength (24.3 MPa) compared to the same implants with 50% macroporosity (14.5 MPa). A reduction in macroporosity without major effects on bone ingrowth seems possible with 40% rather than 50% macroporosity. CONCLUSIO N This in vivo study of bone ingrowth in MBCP ceramics indicates that the influence of macropore size is greater than that of macroporosity percentage. For the same macropore size, no significant difference in newly-formed bone was noted for implants of 40% and 50% macroporosity. Osteoconduction was more efficient for MBCP implants with a 565 than a 300 jiim macropore m pore diameter and a 40%) macroporosity percentage should provide mechanical diameter. A 565 |Li improvement and preserve optimal bone ingrowth in MBCP ceramics. REFERENCE S 1. Daculsi, G. and Passuti, N., Biomaterials 1990, 11, 86-87. 2. De Groot, K., Ann.NY Acad. Sci. 1988, 253, 227-233. 3 Shimazaki, K. and Mooney, V., J. Orthop. Res. 1985, 3, 301-310. 4. Goupy, J. In: La methode des plans d’experience, Dunod, Paris, 1988. 5. Bouler, J.M., Trecant, M., Delecrin, J., Royer, J., Passuti, N. and Daculsi, G., J. Biomed. Mater. Res. 1996, 32, 603-609. 6. Skedros, J.G., Bloebaum, R.D., Bachus, K.N., Boyce, T.M. and Constantz, B., J. Biomed. Mater. Res. 1993, 27, 47-56. 7. Martin, R.B., Chapman, M.W., Holmes, R.E., Sartoris, D.J., Shors, E C , Gordon, J.E., Heitter, DO., Sharkey, N.A. and Zissimos, AG., Biomaterials 1989, 10, 481-488. 8. Daculsi, G., LeGeros, R.Z., Heughebaert, M. and Barbieux, I., Calcif Tissue Int. 1990, 46, 20-27. 9. Trecant, M., Delecrin, J., Royer, J. and Daculsi, G., Clin. Mater. 1994, 15, 233-240. 10. Le Huec, J.C, Schaeverbeke, T., Clement, D., Faber, J., Le Rebeller, A., Biomaterials 1995, 16, 113-118.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MECHANICA L FATIGU E OF HO T PRESSE D HYDROX Y APATIT E S. Raynaud, E. Champion, D. Bernache-Assollant Laboratoire de Materiaux Ceramiques et Traitements de Surface, ESA CNRS 6015, 123, Avenue Albert Thomas, 87060 Limoges, France
ABSTRAC T Polycrystalline hydroxyapatite (HAP) was densified by hot pressing. Dynamic fatigue resistance of the resulting ceramics and degradation process in aqueous solution were investigated. Inmiediate fracture strength in air decreases from 90 MPa to 40 MPa when residual porosity ratio increases from 2% to 6%. The crack propagation exponent n, characteristic of subcritical crack growth, decreases from 22.5–2 in air to 10=^4 in Ringer’s solution for materials densified at 98% of the theoretical value. A value of only n = 14–4 is obtained in air at 94% of relative density. The degradation in solution proceeds by dissolution of crystalline HAP which leads to the decohesion of grains located around residual pores at the surface of the material. KEYWORD S : Hydroxyapatite, Strength, Fatigue, Dissolution. INTRODUCTIO N Hydroxyapatite (Caio(P04)6(OH)2) is a ceramic material of interest for biological applications [1,2]. Although the mechanical properties of dense HAP, fracture strength and toughness, have been widely reported [3-7], few studies concerning mechanical fatigue are available yet [6-8]. The long time application of stresses, even at low level, can induce delayed fracture, depending on environmental conditions. For HAP, it is important to evaluate its behaviour under mechanical loads because this bioceramic is known to be chemically affected by physiological environment [9]. Fatigue phenomenon is analysed in term of subcritical crack growth with the relationship [10]: V=
da dt
= AKf A
(1)
where V is the crack velocity, a is the crack length, Ki is the stress intensity factor at the crack tip, A is a constant, and n is the propagation exponent. The value of n is characteristic of the resistance to mechanical fatigue of a material under a given environment. The fracture strength depends on the stressing rate according to the relation [11]: 75
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Lna f
-Ln B(n + l)an-2 n+1
(2)
-Lna n+l
where a^ is the fracture strength, B is a constant, G\ is the inert fast fracture strength and d is the stressing rate. Thus, the measurement of fracture strength at different stressing rates allows the determination of the propagation exponent n from the slope of the straight line on the graph Lna^ = F(Lna). This work consisted in determining fatigue data for hot pressed HAP ceramics and evaluating the influence of the environment on materials degradation. MATERIAL S AND METHOD S A commercial stoichiometric hydroxyapatite powder was hot pressed under a constant compressive stress of 10 MPa either at 1165 C during 1 hour or at 1100 C during 30 minutes. Sintered blocks were cut into bars of 4*3*25 mm^ and each bar was polished with a 3 jim diamond paste. Quasi static, or immediate fracture strength in air was determined by three-point bending with a 16 mm span and a crosshead speed of 0.2 mm.min-l. The dissolution of HAP ceramics was investigated in Ringer’s solution at 37 C [12]. The degradation of samples was evaluated by measurements of surface roughness for immersion times ranging from one day to three weeks. Mechanical testing of HAP samples by dynamic fatigue in solution was investigated in a device which permits the control of liquid environment (constant temperature of 37 C and constant liquid flowing), the crosshead speed varied in the range 3.10-^ nmi.min-l to 2 mm.min-l. Since the determination of fracture strength measured by three-point bending can be biased by the location of the flaw which initiates crack propagation, an artificial defect was generated on the surface of samples by Vickers indentation under 4.9 N load. RESULT S AND DISCUSSIO N Immediate fracture strength in air decreases from 90 MPa to 40 MPa when the residual porosity ratio increases from 2% to 6%. Dynamic fatigue experiments allow to calculate the crack propagation exponent n (from equation 2). The results obtained from linear regressions of measured data are given in table 1. Values of n = 22.5–2 and n = 14–4 are obtained for HAP ceramics tested in ambient air and densified at 98% and 94% of the theoretical density, respectively. The propagation exponent is n = 10–2 for materials at 98% of relative density, tested in Ringer’s solution. In air, both fracture strength and resistance to subcritical crack growth decrease as the volume fraction of residual pores increases. In the same way, the liquid environment induces a drastic drop of the resistance to fatigue for HAP ceramics densified at 98%. These results are close to those found by G. De With who showed an important sensitivity of HAP Table 1. Analysis of dynamic fatigue plots. Experimental conditions HAP 98 % in air
Linear regression Lna = 4.43 + 4.23 10’^ L n a
HAP 98 % in Ringer’s solution Lna = 4.48 + 9.05 1 0 ’ L n a HAP 94 % in air HAP 94 % in Ringer’s solution
Lna = 4.04 + 6.59 lO’^Lna Not significant - too low confidence level
Meclumical Fatigue of Hot Pressed Hydroxyapatite:S. Raynaud Qi al.
77
180-
^
160140-
e
/
120-
HAP 94% HAP 98%
1 0 0^ 8 06 0-
20 30 40 50 60 70 APPLIE D STRES S (MPa)
Figure 1. Simulated lifetime versus fictive applied stress.
40^
nt-* 10 15 TIM E (days)
20
25
Figure 2. Average surface roughness versus immersion time in Ringer’s solution.
to slow crack growth in air (n = 26), mechanical degradation which is enhanced by water environment (n = 12) [7]. The results on mechanical fatigue characteristics can be clearly illustrated by the evaluation of time-to-failure under a fixed stress. Indeed, dynamic fatigue experiments allow to calculate an estimation of the lifetime of a material. The lifetime under constant applied stress is given by the following relationship : Lntf=LnBa|
n-2
- nLn Ga
(3)
where tf is the time-to-failure and a^ is the applied stress, n and B a " ^ are calculated from the linear regressions constants of dynamic fatigue data. Simulated plots of time-to-failure under constant applied stress are given in Figure 1. They show that the lifetime under mechanical loads is much shorter when the material is subjected to liquid environment. For example, in the case of HAP densified at 98 %, the lifetime under a tensile stress of 30 MPa would be of about 100 hours in solution whereas it would be of more than 100 years in air. An expected lifetime of 20 years in solution would require that stresses do not exceed 15 MPa, which means that HAP ceramics cannot be used in stressed regions of the body. This behaviour also indicates that stress enhanced chemical reaction proceeds at the crack tip, resulting in a very low resistance of HAP to subcritical crack propagation in liquid solution. The influence of liquid environment on HAP degradation can be evaluated by surface observations. Figure 2 shows the plots of average roughness (Ra) of samples surfaces versus immersion time in Ringer’s solution. The difference in initial Ra values between HAP densified at 98% and 94% is due to residual pores at the surface of materials. In any case, the average roughness increases with the immersion time in Ringer’s solution. For HAP densified at 94%, a doubling of Ra value is noticed after three weeks of immersion. A typical SEM micrograph of HAP surfaces after 3 weeks of immersion is given in figure 3a. Compared with the initial surfaces (fig. 3b), the degradation of materials surface after immersion is not uniform. Rings like grooves with dimensions close to the grains size appear in only some regions and dense regions do not seem to be degraded. It can be assessed that the degraded regions are preferentially located around residual pores presents at the surface of initial material.
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Figure 3. SEM micrographs of HAP surfaces (material densified at 94%). (a) after 3 weeks of immersion - (b) initial. This shows that dissolution of HAP is accompanied by the decohesion of some grains. Different interpretations can be found to explain the degradation mechanism of calcium phosphates in liquid. In our case, it can be hypothesised that the degradation proceeds through HAP surface dissolution. This phenomenon would lead to a preferential decohesion of grains located around pores. CONCLUSIO N The mechanical behaviour of HAP depends strongly on the presence of residual pores. In solution, it is subjected to surface dissolution accompanied with grains decohesion around these pores. Slow crack growth is enhanced by the dissolution at the crack tip. Thus, HAP ceramics appear too brittle and sensitive to liquid environments to be used under stresses. Providing a good control of the microstructural design of HAP matrix may be obtained to prevent the detrimental effect of residual pores, composite technology seems to be a way to improve the mechanical reliability and decrease the motion of subcritical crack growth in HAP based materials and finally extend their potential applications.
REFERENCES
Oonishi, H., Biomaterials,1991, 12, 171-178. Hench, L.L., J, Amer. Ceram.Soc, 1991, 74 [7], 1487-1510. Jarcho, M., Bolen, C.H., Thomas, M.B., Bobick, J., Kay, J.F. and Doremus, R.H., J. Mater.ScLA916, 11, 2027-2035. 4. Akao, M., Aoki, H. and Kato, K., ibid.,1981, 16, 809-812. 5. Halouani, R., Bernache-Assollant, D., Champion, E. and Ababou, A., 7. Mater. Sci. Mater. Med., 1994,5,563-567. 6. Thomas, M.B., Doremus, R.H., Jarcho, M. and Salsbury, R.L., J. Mater. Sci, 1980, 15, 891-894. 7. De With, G., Van Dijk, H.J.A., Hattu., N. and Prijs, K., ibid.,1981, 16, 1592-1598. 8. Nonami, T. and Wakai, P., J. Ceram.Soc. Jpn.,1995, 103 [6], 648-652. 9. De Groot, K., In: Bioceramics,Annals New-York Acad. Sci. 1988, 227-233. 10. Evans, A.G., Int.Journ.of Fracture,1974, 10 [2], 251-259. 11. Fett, T. and Munz, D., J. Eur. Ceram.Soc, 1990, 6, 67-72. 12. Barbosa, M.A., In: Biomaterialsdegradation,edited by M. A. Barbosa, 1991, 227-252. 1. 2. 3.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MECHANICA L PROPERT Y CHANGE S IN MACROPOROU S CERAMI C AFTE R IMPLANTATIO N INT O BON E AND MUSCL E M. Trecant-Viana\ J. Delecrin\ J.M. Nguyen", J. Royer\ G. Daculsi’ ^ Centre de recherche interdisciplinaire sur les tissus calcifies et les biomateriaux, Facuhe de chirurgie dentaire, 1 place A. Ricordeau, 44042 NANTES, France ^ Unite statistique et informatique medicales, CHRU, 44042 NANTES, France Laboratoire de mecanique des structures, Ecole Centrale de Nantes, 1 rue de la Noe, 44000 NANTES, France ABSTRAC T Compressive strength and stiffness of MBCP were investigated after 1 to 18 weeks of implantation in rabbit bone and muscle. It was shown that in the two sites the mechanical properties of the implants increased with the implantation duration. Nevertheless, these changes occured to a different degree or followed a different law, suggesting site-dependant structural, physico-chemical and histological modifications. This hypothesis was confirmed by a stepwise multiple linear regression analysis relative to the two mechanical characteristics and four variables (macroporosity, bone, ceramic and microporosity). The role of newly formed bone was confirmed : it filled the macropores and confered a composite stmcture to the implant. In addition the influence of physico-chemical exchanges (dissolution/reprecipitation process) leading to a decrease of the microporosity was revealed. KEYWORD S : calcium phosphates, mechanical properties, in vivo INTRODUCTIO N Clinical applications of calcium phosphate bioceramics as bone grafts substitutes are limited by their poor mechanical properties [1-4]. Macropores are necessary to promote bone formation inside the ceramic [5], but obviously they decrease the mechanical characteristics of the implant. Yet this initial strength has been shown to change when the biomaterial was placed in contact not only with bone [6-9] but also with muscle [9]. Phenomenons taking place in implanted calcium phosphate ceramics (dissolution/reprecipitation process and bone formation) have been the subject of numerous investigations [10-19] but the mechanisms which determine the mechanical properties modifications are still unidentified. The objective of this study w^as to provide quantitative understanding of the effects of physico-chemical changes and bone formation on the compressive strength and Young modulus of implanted Macroporous Biphasic Calcium Phosphate (MBCP). MATERIAL S AND METHOD S Experiment MBCP cylinders were implanted (1, 2, 3, 6, 12, 15 and 18 weeks) into the femoral epiphysis (6x6 mm) and muscle (5x5 mm) of mature male New Zealand rabbits. Samples intended for mechanical investigation were prepared as described in a previous w^ork [8]. After compressive strength and stiffness measurements, specimens were embedded in methyl methacrylate for visual characterization of implant stmcture ; macropore, ceramic and bone 79
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percentages were measured on microradiographs and microporosity was determined using backscattered electron (BSE) imaging [20]. Analysis The means and standard deviations for com.pressive strength and the Young modulus w^ere determined and a stepwise multiple linear regression analysis w^as performed relative to the tv/o m.echanical characteristics and the four variables (macroporosity, bone, ceram.ic and microporosity) to determ,ine the best predictor of a and E of implanted MBCP. RESULT S Visual characterizatio n of implant structure Figure 1 shows the changes in bone and macroporosity after implantation into bone. It can be observed that macroporosity decreased as bone formation occured during the first 12 weeks. Then bone percentage became stable and macroporosity remained unchanged For MBCP implanted in muscle, only a slight tendancy to degradation of the ceramic was obser\^ed at 18 weeks. SEM image analysis showed that microporosity decreased similarly in bone and muscle sites during the implantation period studied (Figure 2). Compressive strength and stiffness Compressive strength increased linearly in both sites during the first 15 weeks (Figure 3). Comparison of the slopes revealed that the increase in compressive strength was greater when N4BCP v/as implanted into bone (Student /-test, n=72, p<0.0001). Concerning the stiffness, it w^as shown that in bone site the Young modulus increased linearly until w^eek 15 and then stabilized, w^hereas in muscle it obeyed an hyperbolic law (Figure 4). Multiple linear regression Microporosity and ceramic were the best em,pirical predictors of compressive strength and stiffriess (Table 1). Only the com^pressive strength of MBCP implanted in muscle v/as less correlated but the significance of microporositv was confirmed by simple linear regression (n=17, p=0.0017). DISCUSSIO N Results of this study showed that mechanical properties of MBCP changed after implantation in both bone and muscle. Nevertheless, according to the implantation site, modifications occured to a different degree (lower slope for compressive strength of MBCP in muscle) or followed a different law (nonlinear development of the Young modulus in muscle), revealing site-dependant structural, physico-chemical and histological modifications. MBC P implanted into bone Visual characterization of implant structure revealed thai bone formation occured inside the m^acropores during the first 12 weeks of implantation followed by stabihzation. On the other hand mechanical strength still increased until the 15^^’ week. Then two phases in the m.echanical properties evolution of implanted MBCP could be distinguished : i) during the first 12 weeks, MBCP mechanical characteristics seemed to be related to partial filling of macropores with new bone and ii) after 12 weeks, modification of the mechanical properties depended only on the composite behavior of the implant. The strengthening between 12 and 15 weeks could result from the maturation of newiy-fbrmed bone (collagen fiber density and orientation, mineralization, trabecular organization and connectivity, as observed with normal bone [2123]) or from a consolidation of the ceramic/new bone interface. Conversely, iht decrease after 15 w^eeks could be due to bone remodelling (increasing number of resorption sites and w^oven bone transformation into lamellar bone [21-23]). This study also revealed the significance of microporosity, w^hich v/ould appear to be the best predictor of compressive strength and elastic modulus. The decrease in m^icroporosity following apatite microcr>’Stals precipitation [20] could have partly accounted for the increase in the mechanical strength o^ MBCP im^planted in bone.
Mechanical Property Changes in Macroporous Ceramic After Implantation:M. Trecant-Vianaet al.
81
MBC P implanted into muscle Regression analysis showed that microporosity was the best predictor of compressive strength and elastic modulus. Thus strengthening of MBCP im.planted into m.uscle appeared to be the result of apatite niicrocr>’stals precipitation in ceramic micropores. mJcroporosit y (%) 7 01
50-i
\ \ 40-]
1
H^
10
12
14
16
10
13
12
14
16
18
vwseks
Figure 1. Evolution of macroporosity ( )and bone ( ) percentages in MBCP implanted into bone.
Figure 2. Evolution of microporosity percentage of MBCP implanted into bone ( ) and muscle ( )
E (MPa)
CT (MPa)
15001 i25Cri
I
t
’ I*
^%\U. 6
8
10
12
10
14
12
14
16
18
20
w e e ks
weeks
Figure 3. Com.pressive strength of MBCP implanted into bone ( ) and muscle ( } as a fijnction of implantation duration. CT (MPa )
BoRc .site Muscle site
Figure 4. Young m.odulus of MBCP imiplanted into bone ( ) and muscle ( ) as a function of implantation duration. E (MPa) i: - 1125.3042 - 98.9118M + 34.8183C (r-0.83) f-: = 837.2275 - 60.8271K4 + 21.4962C (r=0.82)
* CT = 16.042 - 0.34534M (r=U.57)
Table 1. Results of multiple linear regression (M : Microporosity, C : Ceramic). * Simple linear regression confirming the significance of microporosity (p<0.05).
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CONCLUSION The physico-chemical, histological and stmctural modifications result of bioactivity and osteoconduction processes. This study formed bone which not only filled macropores but also gave implant, and the significant role of physico-chemical exchanges biological apatites in ceramic microporosity.
of im.planted MBCP were the confirmed the effect of newly a composite structure to the leading to the precipitation of
REFERENCES 1. Bemache-Assolant, D. L ’IndustrieCeramkjueet Verriere1993, 883, 421-434 2. De Groot, K. Arm.NY Acad.Sci. 1988, 253, 227-233. 3. Metsger, D.S., Rieger, M.R., Foreman, D.W. Fourth WorldBiomaterialsCougres 1992, April 24-28, Berlin, Federal Republic of Germany. 4. Rao, R.W., Boehm, R.F. J. Dent.Res 1974, 53, 1351 -1354 5. Daculsi, G., Passuti, N. Biomat. 1990, 11, 86-87. 6. Martin, R.B., Chapman, M.W., Holmes, R.E., Sartoris, D J., Shors, E C , Gordon, J.E., Heitter, DO^ Sharkey, N.A., Zissimos, AG. Biomat.1989, 10, 481-488. 7. Martin, R.B., Chapman, M.W., Sharkey, N.A., Zissimos, S.I., Bay, B., Shors, E C Biomat. 1993, 14,341-348. 8. Trecant, M., Delecrin, J., Royer, J., Daculsi, G., Clin.Mat. 1994, 15, 233-240. 9 Trecant, M., Delecrin, J., Nguyen, J.M., Royer, J., Passuti, N., Daculsi, G. /. Mater. Sci. : Mater. In Med. 1996,^ 7, 227-229. 10. Daculsi, G., LeGeros, R.Z., Ner>’, E., Lynch, K., Kerebel, B. J. Biomed.Mater. Res. 1989,23,883-894. 11. Daculsi, G., Passuti, N., Martin, S.,. Le Nihouannen, J.C, Brulliard, V., Delecrin, J., Kerebel, B. Rev. Chir. Orthop.1989, 75, 65-71. 12. Daculsi, G., Passuti, N., Martin, S., Deudon, C , LeGeros, R.Z., Raher, S. J. Z?/o/wet/. Mater Res. 1990, 24, 379-396. 13. Daculsi, G., LeGeros, R.Z., Heughebaert, M., Barbieux, 1. Ca/c// Tissue Int. 1990,46, 20-27. 14. Daculsi, G., LeGeros, R.Z., Deudon, C Scanning Microscopy. 1990, 4(2), 309-314. 15. Hardouin, P., Choppin, D., Devyver, B., Flautre, B., Blary, M.C., Guigui, P., Anselme, K. J. Mater Sci. : Mater In Med. 1992, 3, 212-218 16. Heughebaert, M., LeGeros, R.Z., Gineste, M., Guilhem, A. J. Biomed.Mater Res. 1988,22,257-268. 17 LeGeros, R.Z., Parsons, R., Daculsi, G., Driessens, F., Lee, D., Metsger, S. In: Bioceramics: Material characterizationvs. in vivo behavior, 1988, Ducheyne, P . ; Lemons, J. (eds.). New York Acad. Sci. 253, 268-271. Moore, D C , Chapman, M.W., Manske, D J. Orthop Res. 1987, 5, 356-365. 19 Passuti, N., Daculsi, G., Rogez, J.M., Martin, S., Bsamtl J.V. Clin. Orthop. andRel. Res. 1989,248, 169-176. 20 Trecant, M. Ph.D. Thesis, Nantes University, France, 1996. 21 Katz, J.L., Yoon, H.S., Lipson, S., Maharidge, R., Meunier, A., Christel, P. Calcif. TissueInt. 1984, 36, S31-S36. 22 Martin, B. Calcif. Tissue Int. 1993, 53, S34-S40. Martin, R.B., Ishida, J../ Biomech.1989, 22, 419-426
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
DIFFERENC E OF BONDIN G BEHAVIO R BETWEE N FOUR DIFFEREN T KIND S OF HYDROXYAPATIT E PLAT E AND BONE . S.S. Chung\ C. K. Lee^ K. S. Hong^ H. J. Yoon^ 1. Department of Orthopaedic Surgery, Samsung Medical Center. 50 ILWON-dong, Kangnam- Ku, Seoul, 135-710, Korea. 2. Department of Orthopaedic Surgery, Seoul National University Hospital. 3. School of Materials Science and Engineering, Seoul National University. ABSTRAC T The interface between four different kinds of hydroxyapatite(HAp: HA 1, HA 5, HA 6, and HA 9) and bone and the surface of the HAps were examined. The HAps were made with different starting Ca/P ratios (1.5, 1.67, and 1.83) and different maturation temperatures (30 and 90 C). Sintered HAp plates were implanted in rabbits’ tibiae, femora, and muscles of thigh. The XRD analysis, light microscopy, scanning electron microscopy, and Instron were used to examine the formation of hydroxy apatite, new bone formation, bonding behavior and tensile strength. Tensile strength was greatest between HA 9(Ca/P 1.67, 30 C) and bone, though not statistically significant. We also observed more significant new bone formation on the surface of HA 9 using light microscopy. Scanning electron microscopic examination showed partial resorption of the surface of HAp plates and mechanical as well as direct bonding between HAps and bones. [KEYWORDS : Hydroxy apatite, rabbit, tensile strength, bonding behavior] INTRODUCTIO N The autogenous bone graft has many advantages, but there are many complications and problems in harvesting autogenous bone from iliac crest [1]. The allograft as well as heterograft have many problems to be used routinely as graft material [2,3]. Hydroxyapatite has been widely studied and used in clinical field as a bone graft substitute [4,5]. There are many reports comparing biologic responses using different kinds of ceramics, such as hydroxyapatite, tricalcium phosphate, calcite, bioactive glasses, etc. [5,6]. It was our investigation that different kinds of hydroxyapatite as well as different kinds of ceramics could show different biologic responses. Four different kinds of hydroxy apatites have been made and examined for their biologic responses. MATERIAL S AND METHOD S Four kinds of hydroxyapatite powder were selected among 9 kinds of powder and compacted into plate shapes, which were then sintered at a temperature of 1300 C. The HAps were named as HA 1, HA 5, HA 6, and HA 9 and their synthetic conditions were Ca/P 1.5 (maturation temperature of 90 C), Ca/P 1.5 (30 C), Ca/P 1.83 (30 C) and Ca/P 1.67 (30 C), respectively. Eighty-four white rabbits around 3.5Kg were divided into 4 groups according to HAps used. HAp plates were inserted into proximal tibiae of all rabbits(n==21/group) through a slit of medial and lateral cortex (Fig. 1 a), into distal femora 83
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Bioceramics Volume10
(a) (b) Figure 1. Photos of rabbit’s tibia implanted with hydroxyapatite plate across the proximal tibial metaphysis (a) and prepared for biomechanical test (b). (n=6/group) through a slit of lateral cortex and into lateral thigh muscle (n=6/group). Seven rabbits of each group were sacrificed at 2, 4, and 8 weeks and tibiae, femora and HAps in the thigh muscles were harvested. The tibiae were wrapped with saline gauze and deeply freezed at a temperature of-70 C. One femur each sacrificed was immersed in 2.5% glutaraldehyde and the other femur of each sacrificed was preserved in formalin. HAps in the muscle was deeply freezed. After harvesting all specimens, tibiae were thawed for about 4 hours at room temperature, and then segments of proximal tibial metaphyses containing HAp plate were excised and prepared for the biomechanical test to measure the tensile strength between HAp and bone(Fig. 1 b). Traction was applied vertically to the interface between the HAp plate and the bone at a crosshead speed of 5mm/min using an Instron (Instron 8500, Instron corporation, USA). Thin segment (200 \\.mthickness) of distal femur containing HAp plate was obtained using diamond saw for the scanning electron microscopic (SEM, S2460N, Hitachi, Japan) examination to examine the interface between HAp and bone. Another thin segment of distal femur was obtained after decalcification for light microscopic observation. The surface of the HAp plates was examined with scanning electron microscopy before and after embedding of the HAp into thigh muscle. RESULT S AND DISCUSSIO N Table 1 shows average failure load between hydroxyapatite plates and rabbits’ tibiae. At two weeks after insertion of the HAp plates into rabbits’ tibiae and femora, bone and HAp did not bond together. Table 1. Failure load between hydroxyapatite plates and rabbits’ tibiae measured by Instron. HA 9 HA type HA l HA 6 HA S (1.83, 30 C) (Ca/P, temp.) (1.67, 30 C) (1.5,90 C) (1.5, 30 C) PO 2 weeks non-bonding (7) non-bonding (7) non-bonding (7) non-bonding (7) PO 4 weeks 2.54–1.48(7) 2.08 –1.04 (7) 2.25 –1.36 (7) 3.94–1.23 (7) PO 8 weeks 2.50–1.22(7) 2.12 –1.49 (7) 2.34–1.09(7) 4.01 –0.75 (7) HA : Hydroxyapatite, PO : postoperative Ca/P, temp. : Starting Ca/P ratio, maturation temperature Data : Average failure load – standard deviation (n=number of rabbits)
Bonding Behavior BetweenFour DifferentKinds of HA Plate and Bone: S. S. Chung et al.
85
(a) (b) Figure 2. Light microscopic findings 8 weeks after insertion of HA 5 (a) and HA 9 (b). Hydroxyapatite cannot be seen because of decalcification. Black dusts on (a) are remnant of hydroxy apatite. There is more prominent new bone formation on HA 9 than on HA 5. (H&E staining, x 100) At four and eight weeks, the average tensile strength was greatest between HA 9 and bone, though there was no statistically significant difference regarding the types of HAp and postoperative periods. This finding well corresponded to morphological observation using light microscopy and scanning electron microscopy. New bone formation on the HAp plates was observed on light microscopic examination and the new bone formation was more evident on HA 9 (Fig. 2 a-b). On scanning electron microscopic examination, there was very prominent new bone formation on HA 9. Irregular resorption of the surface of HAp plate and bone ingrowth into the irregularity were also more prominent on HA 9 than on other HAps (Fig. 3 a-d).
(c) (d) Figure 3. Scanning electron microscopic findings of the interface between hydroxyapatite plates and bone at 8 weeks after insertion. (x40) (a) HA 1, (b) HA 5, (c) HA 6, (d) HA 9.
Bioceramics Volume10 Di-y Pow^d-er
1300*^0, S i n t e r i n g C a / P=
1 &
.83. SO^C A
^
^ A. .
67, 30 C
c a y p=
il
jiyix^_
j^
./ C a / P= 1 5, 90 C
/
C a / P= 1.5. 30 C ’N
^-^ ; i/
^
A_^_^
._,_^-~^
(a) (b) Figure 4. XRD patterns of hydroxyapatite powder made for this study (a) reveal single phase of the powder and partial decomposition at high temperature (1300 C) to tricalcium phosphate (b). After 8 weeks in the muscle, the surface of the ceramic became granular because of partial resorption of the surface. XRD examination shows partial decomposition of HAp into tricalcium phosphate at high temperature(1300 C) and the amount of decomposition was different according to the synthetic conditions of the powder(Fig. 4 a - b). The partial resorption might reflect the partial decomposition of the HAps at high temperature, because the tricalcium phosphate was known to be subject to partial bioresorption in the biological environment [7]. This resorption resulted granular surface and this seemed to help mechanical interlocking between HAps and bone. Hydroxyapatites showed different biological responses according to the synthetic conditions. Hydroxyapatite made with Ca/P ratio 1.67 and maturation temperature 30 C showed most favorable responses in the rabbits’ tibiae and femora. Further investigation will be performed to produce porous hydroxyapatite using this biocompatible synthetic condition to fmd pore size and configuration, which can show more favorable biological responses. ACKNOWLEDGMENT The present work was supported by the grant of the ministry of Health and Welfare of Republic of Korea.
REFERENCES
1. Arrington E.D., Smith W.J., Chambers H.G., Bucknell A.L. and Davino N.A. Clin. Orthop. 1996,329,300-309. 2. Bolano L. and Kopta J.A. Orthopedics.1991, 14, 987-996. 3. Buck B.E., Malinin T.I. and Brown M.D. Clin. Orthop.1989, 240, 129-136. 4. Emery S.E., Fuller D.A., Bensusan J.S. and Stevenson S. Transactionsof the 40th annual meeting,OrthopaedicResearchSociety,1994, New Orleans, Louisiana. 156-27. 5. Jarcho M. Clin. Orthop.1981, 157, 259 - 78. 6. Neo M., Kotani S., Fujita Y., Nakamura T. and Yamamuro T. / Biomed.Mater.Res. 1992, 26, 255-267. 7. Renooij W., Hoogendoom A., Visser W.J., Lentferink R.H.F., Schmitz M.G.J., van leperen H., Oldenburg S.J., Janssen W.M., Akkermans L.M.A. and Wittebol P. Clin. Orthop.1985,197,272-285.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TREATMEN T OF OSTEOMYELITI GLAS S CERAMI C BLOC K
S BY ANTIBIOTIC-SOAKE
D POROU S A-W
K. Kawanabe, Y. Okada, H. lida, and T. Nakamura Department of Orthopedic Surgeiy, Faculty of Medicine, Kyoto University, 54 Shogoin-Kawaharacho, Sakyo-ku, Kyoto 606, Japan. ABSTRAC T A new dnig delivery system was developedforosteomyelitis using apatite-wollastonite containing glass ceramic (A-W GC) that had been soaked with antibiotics under high vacuum. An 8-mm^ porous A-W GC block (porosity; 70% and 20-30%) and hydroxyapatite (HA) block (porosity; 35-48%) were placed in a bone cement mixer, and mixed with an antibiotic solution. The slow release activity of two antibiotics, isepamicin sulfite and cefinetazole sodium,fromthe porous blocks was tested. An evaluation was made of the slow-release capabilities of the isepamicin sulfitefromthe porous A-W GC block (porosity; 70%) which was maintained at more than 0.5 ^ig/ml after 28 days. However, that from the porous HA block was less than 0.5 ng /ml after 14 days. In a clinical study, two patients with osteomyelitis, including one with infected hip arthroplasty and osteomyelitis of the tibia, were treated and thefecihad completely healed by the end of the follow-up period. INTRODUCTIO N Chronic osteomyelitis is difficult to treat due to the characteristics ofbone, and the object of treatment is to maintain the bactericidal concentration of antibiotic at the infection focus long enough for the healing process to begin. Various antibiotic carrier systems have been developed, including one in which bone cement is mixed with antibiotic-impregnated polymethylmethacrylate (PMMA) beads [1.2] . However, the problem with the use of PMMA beads inserted locally is that subsequent surgery is required for replacement with an autograft. Recently, dmg delivery systems (DDSs) using resorbable materials, collagen [3] , fibrinogen [4] and polylactic acid [5] have been developed. Although it is not necessary to remove them, they cannot be used tofillthe infection site with new bone without bone grafting. We have developed a new DDS using antibiotic-soaked porous A-W GC block, which was demonstrated previously to forma chemical bond with living bone and to have a mechanical strength nearly equal to that of cancellous bone [6] . MATERIAL S AND METHOD S In vitrostudy Two types of porous apatite-wollastonite containing glass ceramic (A-W GC: Nippon Electric Glass Co., Ltd, Otsu, J^an) werefibricatedin 8-mm^ blocks of porosity is 70% (A-W GC 70) and 2030% (A-W GC 20-30), with pore sizes of 200 ^m and 10-50 ^m, respectively. A porous hydroxy^atite (HA) block of the same size, porosity 3548%, pore size 50-300 [im (Bioceram: Sumitomo Pharmaceutical Co.. Ltd, Tokyo, J^an) was used as a control (Fig.l). Two kinds of antibiotic, isepamicin sulfate (ISP: C22H43N5O12 x H2SO4, (X ^2), MW: 569.61) andcefmetazole 87
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^iia.-a
liiiiiili^f c
1 ^^is^^W-^m^’W^MmMmB^, (a) (b) Figure 1. SEM appearance of (a) surface of A-W GC ( porosity 70%) and (b) HA (porosity 35-48%) (CMZ: Ci5Hi6N705S3Na, MW: 493.51) were used. An experimental study of DDS with antibiotic-soaked porous blocks was carried out follows. The three kinds of porous ceramic block were placed in a bone cement mixer (Mixevac 11 High Vacuum System, Stiyker, MI, USA) and mixed with solutions of the two antibiotics, ISP and CMZ (100 mg/ml), and vacuumed at about 500 mmHg for 10 min to allow the antibiotics to soak into the pores. The antibiotic absorption rates of A-W GC 70, A-W GC 20-30 and HA determined by this method were 76.1%, 21.8% and 25.3%/ 8 mm^ volume, respectively. To estimate the concentration of slow-release antibiotic, the blocks were stored in PBS (3 ml) at 37 C. andthePBS was replaced every two days. Preserved PBS containing the released antibiotic was stored at -20t;. An in vitro elution study was then performed using antibiotic assay by high-perfonnance liquid chiomatogi^hy (HPLC), Clinical case Two patients were treated using this method. A 35.year-old man was operated on for osteomyelitis of the right proximal tibia Abscessformationwas observ^ed in the sameregionnine years after the primary operation, and Salmonella was cultured from this specimen. Afier undergoing curettage, appropriate A-W GC blocks were placed in a cement vacuum mixer and soaked with the antibiotics CMZ and ISP. The A-W GC blocks were trimmed and inserted into the osteomyelitis focus. The other patient was a 55-year-old hemophilia man, who was sufieringfix)minfeaed arthroplasty. He underwent revision surgery using antibiotic-soaked A-W GC blocks. The follow-up terms were lyr 6mo and lyr, respectively. RESULT S An evaluation was made of the slow-release capabilities of ISP, and the level was maintained at at least 0.5 ng/ml after 28 days in both A-W GC 70 and 20-30 blocks. However, the level from the porous HA block was less than 0.5 ^ig/ml after 28 days. In the case of CMZ, the three kinds of porous block showed a level of less thanO.5 [ig/ml even after 14 days (Fig. 2). The mean release ratio (antibiotic released in PBS / antibiotic soaked in block) of A-W GC 70, A-W GC 20-30 and HA were 91%, 100% and 100% for ISP, and 48.7%, 37.9% and 42.3% for CMZ, respectively.
Treatmentof Osteomyelitisby Antibiotic Soaked Porous A-W Glass Ceramic: K. Kawanabe et al. CMZ(ug/mi)
ISP(ug/ml) 1 0 0 0 0 0 (] D A-W GC (70% )
D A-W GC (70% ) @ A-W GC (20-30% )
0 A-W GC (20-30% )
HA
2 d 4 d 6 c l 8 d
14 d
18 d
HA
0,1
L
2d 4 d 6 d 8d
14 d
(a) (b) Figure 2. A gr^h showing therateofreleaseof ISP (a) and CMZ (b)fromA-W GC 70, A-W GC 20-30 and HA. hi the dinical cases, both the feci had healed at the end of the follow-up period without complications. The border between the A-W GC blocks and bone became unclear lyr 6mo afer surgery in the case of osteomyelitis in the right proximal tibia (Fig. 3).
(a) (b) Figure 3. A 3 5-year-old-man, osteomyehtis of the right proximal tibia was recurred after nine years afier primaiy operation, (a) Radiogr^hs made one week after second operation. After curettage the infection focus, porous A-W GC blocks were soaked with CMZ and ISP, and implanted, (b) Radiograph made lyr 6mo after second operation. The border between the porous A-W GC and bone became unclear compared with one week.
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DISCUSSIO N Antibiotic-loaded acrylic bone cement beads has been studied in detail and used clinically. However, it must be removed by a further operation, and its long-term implantation is difficult. Some authors have reported that PMM A bone cement has a disadvantage of thermal damage to the antibiotic [1.2] Recently, biodegradable materials have been developed as DDSs for antibiotics. However, these materials can not fill the dead space and may be afocusforrecurrent infection over a long period. Bioactive ceramic is an ideal DDS fiom this view point. Although HA was used as a DDS for antibiotics in a several reports [7.8] , A-W GC has been demonstrated to have higher mechanical strength and bioactivity than HA, and A-W GC 70 absorbed andreleasedmore antibiotic than the porous HA block during a one month period in this study. The high-vacuum system using a cement mixer was efective for soaking the antibiotics into the ceramic pores REFERENCE S 1. Whaling, H., Dingelden, E., Bergmann, R., Reuss, K. The release of gentamicine fix)m polymethylmethacrylate beads. J. Bone Joint Surg., 1975. 60-B. 270-275. 2. Baker, AS., Greenham, L.W., Release ofgentamicinfix)macrylic bone cement. J. Bone Joint Surg., 1988. 70-A.1551-1557. 3. Ascherl, R., Stemberger, A., Lechner, F. Behandelung der chronischen osteomyelitis mit einem koUagen-antibiotika-verbund-vorlaufrge mitteilung. Umfall Chirurg., 1986. 12. 125-127. 4. Zilch, H., Lambiris, E. The sustained release ofcefotaxin fiom afibrin-cefotaxincompound in treatment of osteitis. Arch Orthop Trauma Surg. 1986. 106. 36-41. 5. Wei, G., Kotoura, Y., Oka, M., Yamamuro, T., Wada, R., Hyon, S.H., Ikada, Y., A bioabsorbable deliveriy systemforantibiotic treatment of osteomyelitis. J Bone Joint Surg., 1991. 73-B. 246-252. 6. Nakamura, T., Yamamuro, T., Higashi, S., Kokubo, T., Itoo, S. A new glass-ceramicforbone replacement: Evaluation of its bonding to bone tissue. J Biomed Mater Res., 1985. 19. 71-84. 7. Shinto, Y., Uchida, A.,Korkusuz, F., Araki, N., Ono,K. Calcium hydroxyapatite ceramicused as a delivery system for antibiotics. J Bone Joint Surg., 1992. 74-B. 600-604. 8. Itokazu, M., Matsunaga, T., Kumazawa, S., Oka, M. Treatment of osteomyelitis by antibiotic impregnated porous hydroxyapatite block. Clin. Mater., 1994. 17. 173-179.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CALCIU M HYDROXYAPATIT E CERAMI C IMPLANT S IMPREGNATE D WIT H ANTIBIOTI C FOR THE TREATMEN T OF CHRONI C OSTEOMYELITI S YasuoYamashita,Toru Yamakawa ,Kou Kato Yoshitaka Shinto*, Nobuhito Araki*, Atsumasa Uchida Department of Orthopaedic Surgery, Mie University Faculty of Medicine,Ed)basi 2-174, Tsu-shi, Mie-ken 514, Japan *Department of Orthopaedic Surgery, Osaka University Medical School, Yamada-oka, suita-shi 2-2, Osaka-Fu, 565, Japan. ABSTRAC T Twenty patients with chronic osteomyelitis were treated by implanting calcium hydroxyapatite ceramic with antibiotic into a cavity produced after through surgical excision of necrotic tissue. Within 3 months all of the infected sites had healed. During the period of follow-up ranging from 3 to 75 months we have never experienced a recurrence of infection. There were 3 of those patients had infected prostheses and were successfully revised One patients underwent one stage revision surgery, and another two patients underwent two stage operation. Not only was infection controlled, but there was incorporation of the ceramic material into host bone as judged by radiography. We recommend the use of porous pieces of calcium hydroxyapatite impregnated with antibiotic as a new drug delivery system for the treatment of chronic osteomyelitis. KE Y W O R D S Antibiotics, Hydoroxyapatite, Drug delivery system. Osteomyelitis INTRODUCTIO N Chronic osteomyelitis is known to have difficult surgical problem, particularly in the developing world, despite advances in surgery and more than fifty years experience with antibiotic therapy. Two principles of treatment are paramount: necrotic tissue which has a blood supply unnable to promote normal healing process must be removed, and appropriate antibiotic drugs must be administered [1]. Porous calcium hydroxyapatite (CHA) which is similar to bone mineral composition has excellent biocompatibility, can resist mechanical forces, and is effective in filling cavities and defects in bone [2]. We have already reported porous CHA is very effective as a slow release system for antibiotics in an animal model. [3,4] We have now used porous CHA impregnated with antibiotic clinically and report our experiences, and we believe that this new system is simple, can be performed safely in some few stage, and offers satisfactory results. MATERIAL S AN D METHOD S We treated 11 men and 9 women with chronic osteomyelitis using the principles of surgical debridement, local implantation of CHA impregnated with antibiotic, and systemic 91
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antibiotic administration. The mean age of the patients at the time of treatment was 39.1 years (range: 14 to 77 ). The duration of disease was judged to be from 12 to 132 months (mean: 33 months). At the time of initial presentation all patients had clinical and^or radiological evidence of chronic osteomyelitis. The chronic infection occurred after acute hematogenous osteomyelitis in 12 patients, following open fracture in 3, after closed fracture in 1, and after joint replacement in 3. Nine patients had one or more draining sinuses. Each site of infection was initially aspirated in order to detect a causative organism. The pathogens cultivated were Staphylococcus aureus in 8 cases, Staphylococcus epidermidis in 3, Pseudomonas aeruginosa in two, and Streptococcus pyogenes. Streptococcus pneumoniae, and Klebsiella pneumoniae in each one. The choice of antibiotic for impregnation into the CHA ceramic material was determined by the sensitivity of the cultured organism to drugs. In the 4 patients in whom no organisms were grown broad spectrum antibiotics were selected. We assessed healing by the clinical picture, laboratory findings, and radiological evidence of incorporation of the CHA implant and remodeling of surrounding bone. The duration of follow-up was from 3 to 75 months (average: 47.9 months). Preparation of CH A ceramic impregnated with antibiotic. CHA ceramic blocks were sintered at 1200 C for two hours and had a porosity of 30% to 40% with diameter of the micropores between 40 and 150 micrometers. There was an interconnecting pore structure open to the external surface of the blocks. Operative procedures. The bone cortex was fenestrated to a size permitting removal of all necrotic bone, sequestra, and pathological granulation tissue. During the necrectomy, the chosen antibiotic was packed into a central cylindrical cavity in each porous block and the cavity then sealed with a CHA plug. (Figure 1) The volume of antibiotics packed into the cavity depended on the size of the cavity within the different blocks. The usual range of antibiotic dose in each ceramic block was between 100 and 400 mg. The antibiotics were used, either alone, or in combination. The excavated defect in the bone was then packed with the CHA ceramic pieces which had been each impregnated with the chosen antibiotic. Various sizes and number of ceramic block were used so that the excavated bone defect could be completely filled.
Figure 1 Illustration of CHA impregnated with antibiotics.
Figure 2 Case 1 Chronic osteomyelitis of the proximal region of the tibia.
Calcium Hydroxyapatite Ceramic Implants Impregnated With Antibiotic: Y. Yamashitaet al.
93
CASE R E P O R T S Case 1 An 18-year old man had complained of dull pain in the proximal leg and around the knee. The radiographs showed a sclerotic thickening of the cortex in the proximal tibia. The lesion was opened, anddebrided, and then packed with number of CHA ceramic blocks which had been impregnated with fosfomycin sodium anddbekacin sulfate, ( as no causative organism was cultured ).Four months after the operation, the lesion had been completely healed (Figure 2). Case 2 A 40-year old man with osteosarcoma of the distal femur had infection after wide resection with reconstruction by a tumor knee prothesis. The prosthesis was removed and pathological granulation tissues were debrided and the antibiotic-impregnated CHA blocks were placed in dead space. After 3 months the revision surgery with another tumor knee prosthesis was performed The patients had no recurrence of infection and maintain excellent function after 3 years (Figure 3). Case 3 A 67-year old woman had infection of a knee prosthesis inserted for the treatment of osteoarthritis. She was treated firstly with antibiotic-impregnated Polymethylmetacrylate (PMMA) beads. She had persistent pain and swelling of the knee. A biopsy indicated that Staphylococcus epidermidis was the pathogenic organism. The prosthesis was removed and all necrotic tissue were carefully debrided The antibiotic-impregnated CHA ceramic blocks and another prosthesis was inserted in one stage. Twelve weeks after the revision surgery there had been no recurrence of infection (Figure 4).
Figure 3 Case 2 Chronic osteomyelitis of the mega prosthesis for osteosarcoma in the treatment of distal femur.
Figure 4 Case3 Inplantation of CFIA drug delivery system for the infectedTKA.
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RESULTS At the latest follow-up examination for each of the 20 patients all foci were completely healed Fifteen had pain relief and improvement of laboratory abnormalities within 4 weeks after surgery. In the rest there was resolution of infection within 12 weeks. Recurrence of infection has never occurred after this treatment for follow-up period Radiolucent zones around the ceramic implant gradually disappeared over six months, and in some cases homogenous intramedullary radiodensity surrounded the CHA blocks. There was no radiological evidence of degradation of the ceramic, but in those patients in whom we performed ’second look’ surgery after healing there was some histological evidence of ceramic degradation. DISCUSSIO N It is essential to maintain a high concentration of an appropriate drug at the affected sites for a sufficiently long time, in order to obtain complete eradication of infection of bone and soft tissues. Because of the altered structure of the tissues surrounding an infected site the diffusion of antibiotic drugs into the central part of the infection may require high serum concentration of the drugs. This may cause side effects such as myelosuppression, renal failure, and hepatitis. It is possible to increase the local concentration of antibiotics by impregnating them into carrier vehicles which are implanted into the infected site. PMMA used generally as a bone cement has been the most widely evaluated[5,6,7]. The disadvantages include reduced biocompatibility with bone, short duration of antibiotic release, very low release rate, thermal damage to the antibiotic, and the requirement to remove the PMMA at the end of therapy. Nevertheless this method has been widely used for the surgical treatment of chronic osteomyelitis. Drug delivery system with porous hydroxyapatite ceramic may be effective to apply an appropriate drug for various di sease such as pyogenic osteomyelis, tuberculous osteomyelitis. We are of the opinion that antibiotic-impregnated CHA ceramic is superior to acrylic bone cement systems. Many antibiotic can be placed in a CHA as there is no thermal damage to the drug. All of the impregnated antibiotic can be released over a long period and none is trapped in the ceramic. Biomechanical properties of CHA is similar to those of bone, and the composite of ceramic with newly-formed ingrowth of bone into the pore is almost same as the original bone. As a consequence the antibiotic-CHA ceramic composites both control the infection, restore mechanical strength, encourage osteoconduction into their pores, and avoid the need for further surgery. From these findings, we believe that this new system is simple, can be performed safely in one stage, and offers satisfactory results.
REFERENCES
1. Norden,C.W., Gillespie,W. J. and Nade,S. Infectionin Bones andJoints Boston,Blackwell, 1994, 3-418. 2. Uchida,A., Araki,N., Shinto,Y., Yoshikavva,H., Ono,K. and Kurisaki,E., J .BoneJoint Surg [Br] 1990, 72-B, 298-302. 3. Shinto,Y., Uchida,A., KorukusuzJF., Araki,N. andOno,K., y^on^/om/^wrg 1992, 74-B, 600-604. 4. KorukusuzJF., Uchida,A., Inoue,K., Shinto,Y., Araki,N. andOno,K. J Bone Joint Surg [Br] 1993, 75-B,111-114. 5. Buchholz,H.W., Elson,R.A., andHeinert,K. Clin Orthop1984, 190,96-108. 6. Bayston,R., and Milner,R.D. J Bone Joint Surg [Br] 1982; 64-B, 460-464. 7. Boda,R. Arc/i OrthopTraumaSurg 1982, 101:39-45.
BONE CELLS ONTO BIOACTIVE CERAMICS
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MEASUREMEN T OF INTAC T OSTEOCALCI N CONTENT S IN THE COMPOSIT E OF POROU S HYDROXUAPATIT E CERAMI C AND ALLOGENEI C MARRO W CELL S M. Akahane, H. Ohgushi, T. Yoshikawa, S. Tamai, Y. Dohi,* K. Hosoda** and T. Ohta** Department of Orthopedics, and Public health*, Nara Medical University, Kashihara city, Nara 634, Japan ; Teijin Institute for Bio-Medical Reserch**, Teijin Ltd., Hino city, Tokyo 191, Japan
ABSTRAC T Osteocalcin is synthesized particularly by osteoblast as an extracellular matrix protein. We measured intact osteocalcin contents in allogeneic rat marrow cells/hydroxyapatite (HA) composites implanted at rat subcutaneous sites. At 4 weeks after implantation, bone formation was not detected and only a trace of the osteocalcin was detected in the composite. However, under the immunosupression with FK506, bone formation together with abundant osteocalcin was detected in the composite, and the osteocalcin content was comparable to that of isogenic marrow/HA composites. These results indicate that under the immunosupression, allogeneic bone marrow cells can differentiate into active osteoblasts of which activity is comparable to that of isogenic marrow cells. KEYWORD S Osteocalcin, Hydroxy apatite.
Allogeneic bone marrow
INTRODUCTIO N We have reported that subcutaneous implantation of HA ceramics combined with marrow cells show new bone formation [1]. Osteocalcin (bone Gla protein) is a major noncoUageneous protein in bone matrix and exclusively synthesized by osteoblast. Biochemical analysis of the marrow/HA composites showed that the osteocalcin begun to appear at about 3 weeks when the obvious bone formation initiated, then the osteocalcin contents and bone area increased as time passed [2]. Therefore, the osteocalcin is a useftil biochemical parameter to identify bone tissue. For measuring the osteocalcin, the harvested composites were immediately crushed, homogenized and then measured by using radioimmunoassay (RIA). For some cases, the ceramics were frozen until the assay of osteocalcin [2,3]. Recently, we reported that not only isogenic cells but allogeneic [4,5] cells show bone formation under immunosuppression with FK506. We also established the method of measuring intact rat osteocalcin using anti N- and anti C-terminal rat osteocalcin antibodies raised against a Nterminal 20 residues peptide and a C-temiinal 10 residues peptide of rat osteocalcin [6]. In 97
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this experiment, we measured intact osteocalcin in the cryopreserved composite which was combined with either allogeneic or isogenic marrow cells, and focused on the effect of freezing and immunologic barrier. MATERIAL S AND METHOD S Marrow cellpreparationand Implantationof ceramic Male 6-week-old ACI and 7-week-old Fischer 344 rats were used for donor. Syngeneic 7week-old Fischer rats were used for recipient. The magnitude of the immunological mismatch between ACI and Fischer rat is major. Marrow cell suspensions (5 x 10 ^ nucleated cells/ml) from ACI and Fischer rats were prepared as reported previously [4]. Disk shaped (5mm diameter, 2mm thickness) HA ceramics (Interpore 200, Interpore International, Irvine, Cahfomia) were immersed in each cell suspension from ACI or Fischer rats, then implanted on the back of a recipient Fischer rat. AdministrationofFK506 and Harvest of ceramic FK506 (Fujisawa Pharmaceutical Co., Ltd., Osaka, Japan) was suspended in saline and administered intramuscularly to the recipient rats (Img/Kg/day). As a control, saline was administered. All rats received FK506 or saline every day for 2 weeks and every second days for additional 2 weeks. The ceramics were harvested at 4 weeks after the implantation and stored at -80 C for about 1 month.
Peroxidase | anti N-20 Ab ^^/^N^
N terminal
I
20 peptide s
10 peptide s
1
3
C terminal
^N"" ^
anti C-10 Ab
Figure 1. Schema of the intact osteocalcin measurement by sandwich immunoassay.
IntactOsteocalcinContentsin theCompositeof HA and AllogeneicMarrow Cells: M. Akahaneet al. 99 Measurement of intact osteocalcin
The frozen ceramics were crushed, homogenized in 0.2 % Nonidet P40 containing 1 mM MgCl2 and centrifuged. Osteocalcin was extracted from the sediment by shaking in 2 ml of 20 % fonnic acid for 2 weeks at 4 C. An aliquot (500 \i\) of the fonnic acid extract was then applied to a colunm of Sephadex G-25 and eluted with 10 % fonnic acid. Protein fractions were collected, lyophilized and used for measurement of intact osteocalcin. The principle of the measurement is based on the sandwich iminunoassay which recognizes both N and C tenninal peptides of osteocalcin molecule (Fig. 1). A peroxidase conjugated with a rabbit F (ab’)2 fragment of the anti-N-tenninal 20 residues peptide antiserum (anti-N-20) was prepared [7], Polystyrene balls were dipped in anti-C-10 IgG in phosphate buffered saline (PBS) and incubated at 4 C overnight. Immobilization was tenninated by rinsing the balls with PBS, followed by coating with 1 % bovine serum albumin-PBS at 4 C for 2 days. Standard solutions of purified rat osteocalcin were prepared at concentrations of 0-5 ng/ml. Two hundreds fil of standard solution and 200 \A of peroxidase-labeled anti-N-20 IgG solution with the anti-C-10 IgG-fixed balls were placed in glass tubes. After incubation for 1.5 h at 37 C, each ball was washed three times with saline, then 0.4 ml of tetramethylbenzidine and 0.017 % hydrogen peroxide were added to the tube. The mixture was incubated at 37 C for 30 minutes and the enzyme substrate reaction was tenninated by adding 1 ml of IN H2SO4. The enzyme reaction product was measured by the absorbance at 450 nm. RESULT S AND DISCUSSIO N In this experiment, the harvested ceramics (marrow/HA composites) were immediately immersed into liquid nitrogen and stored at -80 C for about 1 month. Then the ceramics were crushed and maintained at 4 C to extract osteocalcin for about 2 weeks in 20% formic acid. As shown in Table 1 (without FK506), mean intact osteocalcin content in the frozen and stored isogenic marrow/HA composite was 0.68 jug/implant. The content was comparable to 0.59(ag/implant in non-frozen composite which was immediately crushed at the time of harvesting and followed by osteocalcin extraction in 20% formic acid. The osteocalcin
Table 1. Bone fonnation and osteocalcin contents (|ig/implant) in marrow/HA composite, (data are mean – SEM).
Allografts^ Isografts ^ Isografts ^ 1) 2)
With FK506 Osteocalcin Bone fonnation contents + 1.017 –0.224 + 0.854 –0.179 "
" - -
-
-
_
_
Without FK506 Bone fonnation Osteocalcin contents 0.036 –0.004 + 0.682 –0.210 + 0.588 –0.165
The data show the intact osteocalcin contents measured by sandwich immunoassay as described in Materials and Methods. The osteocalcin contents in the isografts (isogenic marrow/HA composites) were measured by conventional RIA as described in ref [3]. The composites were crushed immediately after harvesting and followed by osteocalcin extraction in formic acid.
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contents of the non-frozen composites were detennined by conventional radio-immunoassay (RIA). The data indicate that there was few degradation of osteocalcin molecule during the steps of the measurement and therefore the molecule is quite stable under the low temperature of -80 C and 4 C in the presence of forniic acid. Furthennore, the data indicate the reliability of our previous reports of osteocalcin measurement (RIA) in marrow/HA composite. The bone formation occurred in allogeneic m a r r o w / H A c o m p o s i t e under immunosuppression with FK506. However, it was not observed in allogeneic composite without FK506 and only a trace of the osteocalcin was detected in the composite (Table 1). In this experiment, we measured intact osteocalcin molecule, because it is known that degradation of bone tissue accompanies the degradation of osteocalcin molecule. Therefore, measuring the intact form is crucial in identifying nonnal bone tissue. As shown in Table 1, the amount of osteocalcin in the allogeneic marrow/HA composite with FK 506 was comparable to that of the isografts (with and without FK506). The data of this experiment i n d i c a t e that the bone formed in allogeneic m a r r o w / H A c o m p o s i t e s u n d e r immunosuppression with FK506 did not show rapid degradation which might be initiated by the immunological reaction. Therefore, the surface of HA can support natural process of osteoblastic differentiation of allogenic marrow cells under immunosuppression with FK506.
REFERENCES 1. 2. 3. 4. 5. 6. 7.
Ohgushi, H., Goldberg, V. M. and Caplan, A. I J.Ortop.Res.,7:568-578,1989. Yoshikawa, T., Ohgushi, H., Okumura, M., Tamai, S., Dohi, Y. and Moriyama, T. Calcif Tissue Int., 50:184-188,1992. Inoue, K., Ohgushi, H., Yoshikawa, T., Tamai,.S., Dohi, Y., Hosoda, K. and Ohta, T. Bioceramics Volume. 8:99-102,1995. Sempuku, T., Ohgushi, H., Okumura, M. and Tamai, S. J.Orthop.Res.,14:907-913,1996. Sempuku, T., Ohgushi, H., Okumura, M. and Tamai, S. Bioceramics Volume. 8:397401,1995. Ohta, T., Azuma, Y., Kiyoki, M., Eguchi, H., Hosoda, K., Tsukamoto, Y. and Nakamura, T. Calcif Tissue Int., 59:283-290,1996. Fujiwara, K. Yasuno, M, and Kitagawa, T. Cancer Res., 41:4121-4126,1981.
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
SI-CA- P XEROGEL S AND BON E MORPHOGENETI C PROTEI N AC T SYNERGISTICALL Y ON RA T STROMA L MARRO W CEL L DIFFERENTIATIO N IN VITRO E. M. Santos^, P. Ducheyne^’^, S. Radin^. B. Shenker^, I. and Shapiro^ Departments of ^Bioengineering, ^Pathology, ^Biochemistry and ^Orthq)aedic Surgery, University of Pennsylvania, Kiiladelphia, PA 19104.
ABSTRAC T The effect of a novel bioactive xerogel glass carrier with and without bone morphogenetic protein (BMP) on the osteogenic activity of rat stromal marrow cells was studied in vitro. Cell differentiation was more pronounced on xerogel glass without BMP than that of cells grown on plastic with BMP. Stromal cell differentiation, as measured by alkaline phosphatase activity and osteocalcin synthesis was most increased when the BMP was incorporated or adsorbed onto the xerogel glass. The data suggest that the xerogel glass concentrates osteoinductive proteins at its surface and potentiates their function. KEYWORDS : bioactivity, growth factor, cell culture, cell differentiation INTRODUCTIO N Fracture non-unions and large bone defects represent major clinical pDblems in the practice of reconstructive orthopaedic surgery. ^ Since current treatments for these conditions, such as autogenous bone grafting, have limitations inherent in their use, new approaches for bone tissue repair are valuable.^’*’ One novel approach is the use of osteoinductive bone growth factors, such as bone mOTphogenetic proteins (BMP).^ Bioactive glass has been shown in numerous studies to bond to bone in ydvo. In our group we have shown that porous bioactive glass can serve as an effective template for the growth of bone like tissue in vitro.^These studies also revealed the importance of pre-treating the glass surface. The treatment led to the formation of a calcium phosphate surface layer with proteins adsorbed and incorporated into it. With this treatment neonatal rat calvaria osteoblasts expressed the markers of the osteoblast phenotype extensively within 4-7 days of culture. In contrast, without the treatment, the osteoblast phenotype was not yet expressed within the same culture duration. Using sol gel synthesis exclusively at room temperature, a glass has been made that releases functional bone growth factors in a sustained manner over a period of several weeks. ^ In this paper we document the effect of this material with and without BMP-2 on the proliferation, differentiation and function of rat stromal marrow cells. 101
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MATERIA L AND METHOD S Synthesis: Xerogel discs with a composition of 70% Si02 - 25% CaO - 5% P2O5 (S70) were synthesized using a room temperature sol-gel procedure. Tetramethylorthosilane, calcium metbov^’^thoxide and triethylphosphate were mixed under an argon atmosphere. After casting, soluuu.. ryith or without BMP in 0.1 N acetic acid mixed with a-MEM containing 15% fetal bovine serum was added to sol samples. Solutions with BMP (recombinant human BMP-2, Genetics Institute, Cambridge MA) contained 25 ^lg of it. Gels were aged and dried to 50% of their original weight. The resulting discs (10 mm in diameter and 4 mm in height) were sterilized by exposure to ultraviolet light. Material pre-treatment to form a calcium phosphate surface laver was conducted before cell culturing by immersion in sterile Dulbecco’s phosphate-buffered saline (GibcoBRL/Life Technologies, Grand Island, NY) for 3 hours. Treamient parameters were selected such that the treatment would not cause a significant loss of incorpwated protein from the sol-gel prior to the cell culture experiments. After the treatment, Fourier transform infrared spectroscopy (FTIR) (Nicolet 5DXC) was performed to establish that P-0 bend peaks were present, thereby revealing the formation of calcium phosphate layer. As a second pre-treatment step, for which we developed the rationale in our lab before"*, S70 discs were immersed in 3 ml of tissue culture medium (TCM) containing a-MEM + 15% fetal bovine serum for 1 hour prior to cell culture. Rat Stromal mam?W ggUs were harvested from 4-5 week female Wistar rats using the methods described by Maniatopoulos et al.^ Isolated cells were plated on tissue culture plates in medium containing a-MEM + 15% FBS with 50 U/ml penicillin, 50 ^g/ml streptomycin, and 10" 8 M dexamethasone in a 37 ""C, 5 % CO2 - 95% air incubator. Non-adherent cells were removed by washing after 24 hours. Thereafter, the medium was supplemented with 50 M^g/ml of ascorbate and exchanged every two days. Once the cells were confluent (after 1 week of primary culture) the adherent cells were detached using 0.25% trypsin in Ca- and Mg- free Hank’s Balanced Salt Solution and resuspended in culture medium. 1 x 10^ stromal cells in a 100 |xl TCM solution were seeded on the surface of S70 discs or tissue culture plates (35 mm in diameter) and allowed to attach for one hour. Medium was then added to the culture dish and incubated for either 6 or 10 days. The medium was exchanged every other day. Control groups included cells cultured in tissue culture dishes without BMP (C) otwith 10 ng BMP added to the initial medium and with every medium exchange (C-BT). Experimental groups included sol-gel discs without BMP (SG), or with BMP added as follows: 25 ^g of BMP incorporated into the S70 discs (SG-BI) , 100 ng of BMP added to TCM during the second pre-treatment step (SG-BP), 10 ng of BMP added to the initial medium as well as with every medium exchange (SG-BT) . Aliquots of the medium were collected before every exchange and before cell harvesting. Cell lysate was obtained by aspirating the TCM from the plates, washing the plates with PBS, and then extracting the sample with 1 ml of 3% Triton X-100 in PBS. Cell extracts were analvzed for total protein content, total DNA content, alkaline phosphatase activity (AP) and osteocalcin production using techniques described elsewhere.^ AP activity and osteocalcin synthesis results were normalized to cell number (DNA content) and surface area available for cell growth. Collagen typing was performed by SDS-polyacrylamide gel electrophoresis (SDS-PAGE) run at 100 mV.
Si-Ca-P Xerogels and BMP on Rat Stromal Marrow Cell DifferentiationIn Vitro: E.M. Santos et al.
103
RESULT S The average DNA content of groups and alkaline phosphatase (AP) activity, normalized by DN A content and surface area available for cell culture, are displayed in Figures 1 a, b.
mc 0 0 ^
m 3 days
10 days
1
SG
SG-BI SG-BP SG-BT C-BT
10 days
Figure 1 a,b. Average DNA content (a) and normalized AP activity (b) of control and experimenta l groups. Samples without BM P containe d considerably more DNA than controls (p<0.001) , indicating that cells in samples without BM P proliferate d extensively . This can be explained by considering that the expmmenta l groups with BM P showed an elevate d normalized AP activity over the correspondin g groups without BMP . With BMP , cells move on their differentiatio n pathway at the expense of proliferatio n in the the undifferentiate d state. Cells on S70 (SG) also showed an increased normalized AP activity when compared with stromal cells on tissue culture plastic (C). Similarly, a higher normalized AP activity was noted when cells were grown on S70 discs plus BM P (SG-BI , SG-B P and SG-BT) , when compared to the group of cells grown on tissue culture plastic with BM P added to the medium (C-BT). SG-B P and SG-B I groups were not statisticall y different (p>0.2) at 10 days, but both SG-B P and SG-B I were significantly higher in normalized AP activity than the groups C, SG and C-BT (p<:0.001) . It is also worth noting that the AP activity of the control group of cells on S70 (SG) was higher at both 6 and 10 days than for stromal cells grown on tissue culture plastic with BM P added to the culture medium (C-BT). In general, samples with BM P had much greater protein content than controls (p<0.001) . The samples cultured on S70 discs with incorporate d BM P (SG-BI ) containe d the greatest amount of protein (data not shown). In geno^, the osteocalci n levels of the groups with BM P were greater than those of the groups without BM P (data not shown). At 10 days, medium from cells cultured on S70 pre-treate d with BM P (SG-BP ) containe d the highest concentratio n of osteocalcin . SDS-PAG E analysis, used to vCTify the synthesis of type I collagen, indicated that the most intense bands were present in SGBI samples, followed by the oth^ samples that receive d BM P treatment . DISCUSSIO N AN D CONCLUSIO N The data shows that the effect of the material substrate on cellular differentiatio n was as important as that of BM P added to the tissue culture medium. The material itself without BM P produced a greater response in AP activity of cells than tissue culture medium to which BM P was
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added. BMP is a powerful stimulator of diff^^ntiation in stromal marrow cells in the presence of the surface-treated sol-gel material. Actually, the effect of BNfP on cell diff^entiation is potentiated as follows from the low amount of DNA, and the high AP activity and osteocalcin production. At shorts time periods (six days) BNfP added to solution combined with the xerogel (SG-BT) had the greatest effect on stimulating cell differentiation, while at the longer time periods (ten days) BMP adsorbed to the surface of the xerogel (SG-BP) ac incorporated within (SG-BI) had the greatest effect on differentiation. In view of the data on the effect of BMP adsorbed or incorporated into the glass on differentiated cell function, our results suggest that the sol-gel material concentrates osteoinductive and growth regulatory proteins, naturally present in serum, at its surface in a biologically active fwm. Furthermore, the results indicate that the calcium phosphate surface layers formed on Si-containing, biologically reactive materials play an important role in stimulating osteoblastic differentiation of stromal marrow cells. BMP added to the materials stimulates extracellular matrix formation as measured by total protein present in culture. The cells exposed to xerogel with incorporated BMP produced substantially more protein (40-50% more) than cells on xerogels with adsorbed BMP (p<0.01) and even more than control cell populations without BMP. These results suggest that BMP is a powerful stimulator of extracellular matrix formation in stromal marrow cells and that a sustained release carrier is the most effective way to stimulate matrix production by cells. In summary, this study together with other studies from our laboratory reveals numerous benefits in using a porous xerogel for delivery of a growth factor. The porosity and chemical structure of the xerogel allow sustained release of the factor over a defined time period. In addition, the surface of the xerogel can be modified to generate a local environment that promotes both cell adhesion and extracellular matrix formation. In this study, we optimized synthesis of the xerogel with respect to effecting cellular activity and controlled release. And, as the data suggest, BMP is taken up and then continuously made available from the Si-Ca-P xerogel. Importantly, the growth factor is biologically active in that it influences cell differentiation and function.
REFERENCES
1. Praemer, A., Fumer, S. and Rice, S.D., Park Ridge, IL: Am.Ac. ofOrthop.Surgery,1992. 2. J. M. Lane and H. S. Sandhu, Orthop.Clin.NAm., 1987,18 (2), 213-225 . 3. C. J. Damien and J. R. Parsons, JAppliedBiomater.,1991, 2, 187-208 . 4. A. El-Ghannam, P. Ducheyne, and I. M. Shapiro, JMiomed.MatJies,1995, 29, 359-370 . 5. S. B. NicoU, E. M. Santos, S. Radin, R. S. Tuan and P. Ducheyne, Biomaterials,in press. 6. C. Maniatopolous, J. Sodek, and A. H. Melcher, Cell Tissue Res.,1988, 254, 317-330 .
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EFFEC T OF SURFAC E INSTABILIT Y OF CALCIU M PHOSPHAT E CERAMIC S ON GROWT H AN D ADHESIO N OF OSTEOBLAST-LIK E CELL S DERIVE D FRO M NEONATA L RA T CALVARI A T. Suzuki, M. Hukkanen,! L. D. K. Buttery,i J. M. Polak,i Y. Yokogawa,^ K. Nishizawa,^ F. Nagata,2 Y. Kawamoto,^ T. Kameyama,^ and M. Toriyama,^ Department of Biological Science and Technology, Science University of Tokyo, 2641 Yamazaki, Noda, Chiba, 278 Japan; i Department of Histochemistry, Royal Postgraduate Medical School, Du Cane Road, London W12 ONN, United Kingdom; ^National Industrial Research Institute of Nagoya, Kita-ku, Nagoya 462, Japan; ABSTRAC T Using composite sinters of hydroxyapatite (HAP) and 6-tricalcium phosphate (TCP) as culture carriers, whose Ca/P molar ratios were adjusted in stepwise fashion to values of L50, L55, L60, L64 and 1.67, osteoblast-like cells derived from neonatal rat calvaria and a mouse fibroblast L-929 cell line were cultured and cell growth rates, adhesion and metabolism were investigated. Growth rates of both cells were accelerated on the TCP-HAP ceramics compared to zirconia ceramics. Adhesiveness of osteoblast-like ceUs on TCP-HAP ceramics was also enhanced more than on zirconia alone. Measurement of alkaline phosphatase activity showed that the cellular activity was enhanced by culturing osteoblast-like cells with high concentrations of calcium. These results suggest that the surface of TCP-HAP ceramics, especially 100% HAP ceramics, is effective to accelerate growlh and differentiation of osteoblast-like cells. KE Y WORDS : bioceramics, osteoblast, hydroxyapatite, tricalcium phosphate, tissue culture INTRODUCTIO N In our previous studies the chemical and physical structure of the sintered mixtures of HAP and TCP carriers were shown to influence the adhesion and growth of mouse fibroblast L-929 cells [1-3]. Observations derived from in vivo studies of implanted ceramics have suggested that the slow solubility of calcium phosphate ceramics seems to be necessary for osteoconductivity and remodeling of the fibrous connective tissues, and for natural bonding of the tissue with the implant [4]. Furthermore, hydroxyapatite ceramics have been reported to have good biocompatibility with osteoblastic cells with no biotoxic reactions [5]. These results suggest that the surface of calcium-phosphate ceramics have the potential to stimulate remodellmg of bone in vivo by acceleratmg the growth of osteoblastic cells. However, the TCP-HAP ceramic hybrid surface can be relatively unstable and may be causative factor for reduced adhesion of cells to the ceramic material, cellular damage and for subsequent inflammatory reactions in osteoarticular tissues. In view of this we have investigated the effects of the surface structure of biocompatible ceramics on the growth and adhesion of osteoblast-enriched neonatal rat calvaria cells cultured on the TCP-HAP composite ceramics. 105
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MATERIAL S AN D METHOD S Ceramic disk plates, 30 mm in diameter and 2 mm thick, consisting of TCP and HAP in various ratios, were used in this study. The Ca/P ratio was adjusted to values of 1.50, 1.55, 1.60, 1.64 and 1.67. Zirconia (ZrOz) disk plate and a polystyrene culture dish (LUX) were used for comparison. Primary osteoblast-like cells were obtained by sequential collagenase digestion of rat calvaria from neonatal rats according to the method of Wong and Cohn [6]. The mouse-fibroblast cell line, L-929 (NCTC Clone 929), was used for comparison. The effect on cell adhesion and growth patterns of Ca-enriched medium was investigated by culturing cells in the medium supplemented with CaCh. The concentration of Ca was adjusted to 2.5, 5, 10, 17, 33, and 50 mM by dissolving CaCh into MEM containing fetal bovme serum (10%). The initial cell density was adjusted at 2 x 10^ cells/cm^ of osteoblast (OB)-like cells and at 5 x 10^ cells/cm^ of L-929 cells. Cultured cells were sonicated and aliquots of the supematants were analyzed for total alkaline phosphatase activity using an Alkalme phosphatase B-test Wako assay kit (Wako Chem. Co., Osaka, Jpn). RESULT S Figure 1 shows the time-dependent variations of concentrations of Ca and PO4 ions after immersion of TCP-HAP ceramics in distilled water and culture medium. The solubility in distilled water was higher for TCP than for HAP. Tlie dissolved Ca and PO4 in the medium were likely to be removed from the solution by deposition on immersed TCP-HAP ceramics. The variation of zeta potential of the TCP-HAP ceramics was unstable with significant changes in the charges as shown in Figure 2. These results suggest that the stability of the surface is closely
(a) 2.0
1.50
1.55 1.60 1.65 1.70 Ca/P molar ratio Figure 1. Time-dependent profiles of calcium and phosphate in distilled water (a), and in culture medium (b) by immersing TCP-HAP ceramics for 1, 3, and 6 days.
1.50 1.55 1.60 1.65 1.70 Ca/P molar ratio Figure 2. Time-dependent changes of surface zeta potential of TCP-HAP ceramics by immersing in distilled water (a), and in culture medium (b).
Effect of Surface Instabilityof TCP-HAP Ceramics on Growth and Adhesion: T. Suzuki et al.
S 40
LJU^LA^
J 30
< I 201U
107
J
jri’"^^
6dj--|
10
o
0 1.50
1.55 1.60 1.65 Ca/P molar ratio Figure 3. Population of OB-like cells on TCP-HAP ceramics and on a ZrOz; Z.
’
II
’ ,
’
,
,’
’
Ca/P molar ratio Figure 4. ALP activity of OB-like cells on TCP-HAP ceramics and on a Zr02; Z.
related to both reactions of association and dissociation of calcium and phosphate in culture medium. Figure 3 shows time-dependent variations in populations of anchored OB-like cells on TCP-HAP plates as a function of Ca/P molar ratio of the carrier material. The cell numbers on Zr02 disk plates at 2 d were higher than on TCP-HAP disk plates except those on TO-HIOO plate. However, the cell numbers on TIOO-HO, T70-H30, and TO-HlOO plates reached higher values than those on Zr02 plates at 9 d. Variations in total cellular ALP activity is shown in Figure 4. Total ALP activities of OB-like cells on TCP-HAP plates were expressed at the higher level than those on ZrOz throughout the time-course studied. Figure 4 also shows that ALP activity of OB-like cells on TCP-HAP plates was elevated according to increase in the HAP content. The adhesiveness of OB-like cells on TCP-HAP plates except for TO-HIOO was weaker than on ZrOz during the first four days of culture as shown in Figure 5. The adhesion strength measured by trypsination method [1], F^^^^, at 4 d culture on ZrOz was 0.53, while it was from 0.39 to 0.53 on TCP-HAP plates. However, the adhesion of OB-like cells on TCP-HAP plates increased during the later growth phase. The highest value of F^ ^ for OB-like cells was obtained on TO-HIOO plate at 9 d of culture.
Figure 6 shows the effects of Ca concentration on adhesion of
1.55 1.60 1.65 Ca/P molar ratio Figure 5. Adhesiveness of OB-like cells on TCP-HAP ceramics and ZrOa.
1
10 100 Initial Ca concentration (mM) Figure 6. Effects of Ca concentration on adhesiveness of OB-like cells.
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OB-like cells on LUX carrier. High concentrations of CaCh damaged the cells and decreased their growth rate, while total cell number and ALP activity increased gradually when Ca concentration was lower than 17 mM. The adhesion of OB-like cells at 2 d and 4 d were almost constant when Ca concentration was lower than 10 mM. However, cells cultured in medium with 5 mM and 10 mM of Ca demonstrated increased adhesion at 6 d and further at 9 d of culture. These results support the view that the adhesion of OB-like cells on TCP-HAP plates is increased compared with inert materials and dependent on Ca/P ratio of the ceramic carrier material. DISCUSSION Implanted porous calcium phosphate ceramics have been shown to have strong osteogenic potential with new bone forming in the pore regions. The instability of TCP-HAP ceramics is thought to be a factor in accelerating osteoconductivity and improving the chemical affinity and connectivity with bone tissue in vivo. However, the high solubility and reactivity of bioceramic surfaces may cause damage to adherent cells, which in turn may stimulate inflammatory responses in surrounding tissues. Our previous studies showed that the high solubility of the surface structure reduced the adhesion strength between L-929 cells and TCP-HAP ceramics [1]. Nevertheless, morphological observations and a series of other analyses showed that the weak solubility of the TCP-HAP ceramics were preferable not only for the grov^h of cells, but the surface was shown to be stable enough also to enhance cell adhesion. The accelerated increase in the ALP activity of OB-like cells on TCP-HAP ceramics suggests that the calcium phosphate surface supports both the growth and differentiation of osteogenic cells. In conclusion, this study has further clarified the effects of TCP-HAP composite ceramics, and especially 100% HAP, on the growth and adhesion of OB-like cells during the early phase of the mineralization process. Time-dependent variation of the cellular ALP activity suggests that the TCP-HAP ceramics are able to stimulate differentiation of osteoblasts and therefore to promote osteogenesis in vivo, ACKNOWLEDGMEN T This work was supported by the Bilateral International Joint Researdi Grant by the Japanese Government, the British Council (promotion of the joint research project between United Kingdom and Japan) and the Medical Research Council U.K. The authors are grateful to Dr. Francis J. Hughes of Department of Periodontology, London Hospital Medical College for his advise on osteoblast culture.
REFERENCES 1.
2. 3. 4. 5. 6.
Suzuki, T., Yamamoto, T., Toriyama, M., Nishizawa, K., Yokogawa, Y., Mucalo, R.M., Nagata, F., Y. Kawamoto, Y., and Kameyama, T., /. Biomed. Mater. Res., 1997, 34, 507-517. Nishizawa, K., Toriyama, M., Suzuki, T., Kawamoto, Y., Yokogawa, Y., and Nagata, F., Nihon Kagaku Kaishi, 1995, 63-67. Toriyama, M., Kawamoto, Y., Suzuki, T., . Yokogawa, Y., Nishizawa, K., and Nagata, F., J. Ceram.Soc. Jpn, 1995, 103, 46-49. Niki, M., J. Jpn. Soc. Biomater.,1994, 12, 5-21. Bagambisa, F.B., and Joos, U., Biomaterials,1990, 11, 50-56. Wong, G.L., and Cohn, D.V., Nature, 1974, 252, 713-715 (1974).
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
HISTOLOGICA L EVALUATIO N OF CRYOPRESERVE D MARRO W CELL S
CULTURE D
BON E
GRAF T
USIN G
H. Nakajima^’^, T. Yoshikawa^’^, H. Ohgushi’, M. Akahane\ S. Tamai\ K. Mishima^ and K. Ichijima^ Departments of Orthopaedic Surgery^ and Pathology^, Nara Medical University, Kashihara City, Nara 634, JAPAN.
ABSTRAC T Rat bone marrow cells obtained from femora were cultured in a standard medium for ten days. The cultured cells were released by trypsin treatment and stored at -196 C (liquid nitrogen) in a preservative culture medium containing 10% dimethylsulfoxide. After 3 months, the cryopreserved cells were thawed and subcultured in porous hydroxyapatite(HA; Interpore 500). The medium consisted of Eagle-MEM containing 15% fetal bovine serum, antibiotics, ascorbic acid, P-glycerophosphate and dexamethasone. After 2 weeks of the subculture, the composites of the cultured bone and porous HA were subcutaneously implanted into syngeneic rats. These implants were harvested 2 and 4 weeks postimplantation and prepared for histological analysis, which showed bone formation together with active osteoblasts in many pore regions. These results indicate the osteogenic capacity of cryopreserved marrow cells in porous HA. KE Y WORDS : cryopreservation, hydroxyapatite, bone marrow, dexamethasone, bone graft
INTRODUCTIO N Bone marrow stromal cells are fibroblast-like cells that are thought to include precursors of different connective tissue phenotypes, including adipocytes and osteoblasts[l]. Previously, Maniatopolous et al. demonstrated that in the presence of dexamethasone, cultures of rat marrow stromal fibroblasts formed mineralized nodules that resembled bone[2]. We also have reported the appearance of osteogenic cells in porous biomaterials by transplantation of marrow/materials composites to heterotopic (extraosseous) sites[3, 4], and more importantly we have found that not only fresh marrow but also cultured human fibroblastic cells can show the osteogenic response resulting in bone formation at heterotopic sites[5]. Cryopresevation is clinically available technique for the prolonged storage of many mammalian tissues such as pancreatic islets, blood, bone marrow cells and spermatozoa[6]. At present experiment, we analyzed the osteogenic capacity of cultured bone graft derived from cryopreserved marrow cell culture. 109
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MATERIAL S AND METHOD S Materials; Coralline hydroxyapatite ceramic (HA) was used in this experiment. A hydrothermal chemical exchange converts the original calcium carbonate exoskeleton of coral into a completely inorganic replica of HA. The solid and porous components of the microstructure are completely interconnected. The pores average is 600 micrometers in diameter and their interconnections average is 260 micrometers in diameter. Interpore International (Irvme, CA) manufactured the HA and labeled it as Interpore 500. We used the cubic structure (5x5x5mm) of the HA. Preparation of Cryopreserved Marrow Cells; Marrow cells were obtained from the femoral shafts of a seven-week-old male Fischer rats. The femora were taken and cut at both ends of epiphysis under aseptic conditions. The marrow cells were flushed out with the 10ml of a standard culture medium using a 10-ml syringe with a 21 gauge needle. A standard culture medium consisted of an Eagle’s minhnal essential medium (MEM) containing 15% fetal bovine serum (ICN Biomedicals Japan Co. Ltd.) and antibiotics (lOOU/ml penicillin, lOO^ig/ml streptomycin and 0.25|Lig/ml amphotericine B, Sigma). The cells were collected in a T-75 flask (COSTAR) containing 15ml of the culture medium. The medium was changed after 24 hours to remove nonadherent cells. Subsequently, the medium was renewed 3 times a week. Cultures were maintained in a humidified atmosphere of 95% air and 5% CO2 at 37 C. After 10 days in primary culture, the cells were released by 0.1% trypsin treatment and concentrated to 10^ cells/ml in a preservative culture medium by centrifiigation at 900rpm for 5min at room temperature, then frozen and stored at -196 C in liquid nitrogen for 3 months. The preservative culture medium consisted of Eagle’s MEM containing 10% dimethylsulfoxide and 20% fetal bovine serum. At the time of culture, the frozen cells were thawed and reseeded on T-75 flask in 15ml of standard medium for removal of non-adherent cells which were probably dead. After 24 hours, the cells were trypsinized and concentrated to 10^ cells/ml in a standard medium. The trypsinized cells were used for subcultures on well or in porous hydroxyapatite(HA). Culture of Cryopreserved Cells on a Culture well; To ascertain the osteogenic capacity of cryopreseved marrow cells, we cultured these cells for alkaline phosphatase(ALP)-stain. The cells were seeded at 100 X 10^ cells/35mm Falcon tissue culture plate and cultured for 2 weeks. The subculture was done with 2 ml of the standard medium supplemented with lOmM Na (3 glycerophosphate, 82|ig/ml vitamin C phosphate(L-ascorbic acid phosphate magnesium salt nhydrate) and with or without lO’^M dexamethasone(Dex), which is known to induce osteoblastic differentiation of marrow stromal cells in culture conditions. The culture medium was renewed 3 times a week. After 2 weeks of subculture, ALP was stained as described previously[7]. Culture of Cryopreseved Cells in Porous HA ; The cells were subcultured in HA as described previously[8]. Briefly, HA block were soaked in 4ml of the cell suspension in a CO2 incubator. After 2 hours of incubation, each composite of HA and cells was transferred into one well of a 24-well plate and subcultured in 1ml of the medium containing P-glycerophosphate, vitamin C phosphate with or without Dex, as described above. The medium was renewed 3 times a week. After 2 weeks of subcultures, the composites were prepared for following implantation. Six composites were implanted subcutaneously at 6 sites of the back of each syngeneic rat: three composites subcultured in the presence of Dex were separately implanted in the right side, and the other three composites subcultured in the absence of Dex were separately implanted in the left side. The grafts were harvested at 2 and 4 weeks after implantation and prepared for decalcified sections. The specimens were fixed in 10% buffered formaline, decalcified (K-CX solution, Falma Co., Tokyo) and stained with hematoxylin and eosin. These specimens were examined under light microscopy.
Evaluation of CulturedBone Graft Using CryopreservedMarrow Cells: H. Nakajima et al.
Dex (+)
Ill
Dex (-)
Fig. 1 ALP stain of cryopreserved cells cultured with or without Dex. Many ALP positive nodules were seen in the cultured tissue with Dex.
2W
4W
Fig.2 Dex-treated composite 2 weeks (left) and 4 weeks (right) postimplantation. Both show active bone formation on the HA pores. Right shows regenerated marrow tissue associated with new bone formation at 4 weeks. (HE stain, X 40)
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RESULTS Cryopreserved marrow cells which were subcultured for 2 weeks with Dex on culture dishes showed many calcified nodules. The calcified nodules were stainable for ALP (Fig.l). The same ALP activity was seen on the nodules when subcultures were done using fresh marrow cells. These findings indicate that cryopreserved marrow cells have osteogenic potency same as fresh marrow cells, because alkaline phosphatase activity is known to reflect osteoblastic activity. The Dex-treated composites using cryopreserved cells showed obvious bone formation sevral weeks after implantation. As shown in Fig. 2, many active osteoblasts in contact with the newly formed bone matrix were clearly seen in the pore areas. Number of osteoblasts and area of bone increased as time passed. In contrast, it was hard to detect bone formation in the Dex-untreated composites. DISCUSSION We demonstrated that cryopreserved marrow cells, which were stored for 3 months in liquid nitrogen, show histological active bone formation both in vitro and in vivo conditions. This indicates that cryopreserved marrow cells have comparable osteogenic capacity to that of fresh marrow cells. HA ceramics are useful as bone graft substitutes and can show osteogenesis when combined with freshly isolated marrow cells[3, 4] or with their cultured bone[8-12]. Present resuhs also showed the osteogenic potential of cryopreserved marrow cells. If new methods are applied clinically, it may eliminate the limitation of bone graft and furthermore be available for elderly patients with low osteogenic ability using marrow cells frozen and stored at young age of the patients. REFERENCES 1. Owen M. In: Bone and Mineral Research3 Elsevier, Amsterdam 1985, 42-53. 2. Maniatopolous C, Sodek J. and Melcher A.H., Cell TissueRes., 1988, 254, 317-330. 3. Ohgushi H., Goldberg V.M. and Caplan A.I., J. Orthop.Res., 1989, 7, 568-578. 4. Yoshikawa T., Ohgushi H., Okumura M., Tamai S., Dohi Y. and Moriyama T., Calcif Tissue Int.,1992, 50, 184-188. 5. Ohgushi H. and Okumura M., Acta Orthop.Scand, 1990, 61, 431-434. 6. Farrant J. In: Low Temperature Preservationin Medicineand Biology Pitman Medical, 1980, 1-18. 7. Ohgushi H., Dohi Y., Katsuda T., Tamai S., Tabata S. and Suwa Y., J. Biomed Mat. Res., 1996,32,333-340. 8. Yoshikawa T,, Ohgushi H. and Tamai S., 1 Biomed Mat. Res., 1996, 32, 481-492. 9. Ohgushi H., Dohi Y., Yoshikawa T., Tamai S., Tabata S., Okunaga K, and Shibuya T., J. Biomed Mat. Res., 1996, 32, 341-348. 10. Yoshikawa T., Ohgushi H. and Tamai S. In: Bioceramics Volume 9, Pergamon, Oxford 1995, 421-426. 11. Yoshikawa T., Ohgushi H., Akahane M., Sempuku T., Tamai S. and Ichijima K. In: BioceramicsVolume 9, Pergamon, Oxford 1996, 65-68. 12. Yoshikawa T., Ohgushi H., Dohi Y. and Davis J.E., Bio-Medical Mater. Eng., in press,
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
INTERACTION S OF BIOCERAMIC S ON HUMA N OSTEOARTHRITI S (OA) TYPE B SYNOVIOCYTES . EFFECT S ON INTERLEUKI N LEVEL S AND LIPOXYGENAS E PATHWAY S Bertrand Liagre^, Jean-L. Charissoux^, Marie-J. Leboutet^, Didier Bemache-AssoUant^ and Jean-L. Beneytout^ ^ Facultes de Medecine et de Pharmacie, E.R.S. CNRS 6101, 2 rue du Docteur Marcland, 87025 Limoges Cedex, France ^ Service dOrthopedie-Traumatologie et ^ Service d’Anatomie Pathologique, CHRU de Limoges, Hopital Dupuytren, 2 avenue Martin Luther King, 87042 Limoges Cedex, France ^ LMCTS, U.R. A. CNRS 320,123 avenue Albert-Thomas, 87060 Limoges Cedex, France ABSTRAC T We investigated the effects of the biomaterials hydroxy2q)atite (HAP) and fluorapatite (FAP) on cultured human OA type B synoviocytes by analysing interleukin-la (IL-la) production and arachidonic acid metabolism via lipoxygenase pathways. A portion of opsonized particles were endocytosed and found in numerous phagolysosomes in human synoviocyte cytoplasms. The present stucfy demonstrates that HAP and FAP calcined at 700 C induced a decrease in IL-la production but markedly decreased the synthesis of lipoxygenase products after one month incubation with the particles. This model will allow us to stucfy the possible inflammatory response (arachidonic acid metabolism, proinflammatory cytokines) that can be induced by any biomaterials used in orthopaedics. INTRODUCTIO N The nature of the response of human tissue to particulate debris from total joint prostheses has not been clearly defined. Some early investigators thought that hip prostheses were inert [1], but others postulated a hypersensitivity rejection phenomenon in some patients [2]. Most histological studies have confirmed that the predominant cell types in the tissue surrounding loosened total hip prostheses are macrophages, giant cells andfibroblasts[3]. IL-1 is a pleiotropic cytokine and a potent mediator of the inflammatory response in connective tissues with the c^)acity to induce the synthesis and the secretion of coUagenases such as matrix metalloprotease-1, stromelysin-1, gelatinase A and prostaglandin Ej (PGEj) by synoviocytes and chondrocytes [4]. Lipoxygenase metabolites have been demonstrated to modulate a host of biological activities and appear to be of particular importance in inflammation, fibrosis and the immune response [5]. The aim of our study was to understand \^ether release of particulate debris from total joint prostheses in organism could induce an inflammatory phenomenon over time. The majority of previous studies were done either in vivo on animal models by injections of particulate prosthesis material or in vitroon macrophages [6] and on OA synovial lining cells [7]. The behaviour of human osteoblasts cultivated on HAP-coated disks [8] was studied in vitroand comparison of the effects of HAP and FAP were studied on the release of growth hormone on human osteosarcoma cells [9]. In our work and contrary to previous studies on macrophages, we investigated the effects over time of biomaterials such as HAP and FAP on human OA type B synoviocytes in culture by analyzing IL-la production and arachidonic acid metabolism via lipoxygenase pathways. 113
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MATERIAL S AND METHOD S HAP, Caio(P04)5(OH)2, is a chemical component ^ o s e composition is very close to that of natural bone. HAP powder was prepared in an aqueous medium by slow addition of a (NH4)2(HP04) solution containing NH4OH to boiling Ca(^0^)2,4 HjO solution at pH 9. The powder was calcined at 700^C in order to improve its crystallization state, and to eliminate the volatile impurities (NH4NO3). For characterization, a small fraction of the powder was heated at lOOO^’C and then analysed t^ X r ^ difGraction and a basic test to reveal the presence of CaO. Xray difGraction patterns (Fig. 1) show no additional peak of C2^(?0^ and CaO, and the basic test with phenolphtaleine (AFNOR 894-066) was negative. This is a very sensitive test vAiich confirmed that the Ca/P ratio of this powder was 1.67 – 0.004, the theoretical stoichiometry of pure HAP. Particles were sterilized and opsonized was done according to Estensen et al. [10]. OA synoviocytes were isolated from fresh synovial biopsies obtained from OA patients undergoing hip synovectomy. Synovia were minced and digested by enzymatic reaction. Adherent cells were cultured in complete medium and, at confluence, cells were trypsinized and passed. Cells were used between passages 4 and 8 when they morphologically ressembled "fibroblast-like** synoviocytes. HAP and FAP powders were opsonized and incubated at 10^ particles/10^ cells for one week and for one month. Particles and cells were seeded in culture flask at the same time. Culture medium was changed every two days. No additional particles were added during the entire study. Particles were not removed during medium change because they bound to synoviocyte cell membranes. Synoviocyte culture supematants, henceforth referred to as synoviocyte conditioned medium, were collected, centrifiiged, aliquoted and stored at -80^C. All synoviocyte conditioned media were evaluated for IL-la and the collection period was two days because of the short IL-la half life. The amounts of IL-la in synoviocyte conditioned media were measured by commercially available EIA kits (Cayman Chemical Company, Ann Arbor, NO, U.S.A.). Furthermore, synoviocyte cultures were collected for analysis of exogenous arachidonic acid metabolites t^ reverse-phase-high pressure liquid chromatogr^hy (RP-HPLC) according to Bonnet et al. [11]. Acidified samples and synthetic standards were injected on a 5 ^m Radial-Pack C18 cartridge (Waters-Millipore, Saint-Quentin en Yvelines, France) protected by a Waters CI8 Guard Pack precolumn. Lipoxygenase metabolites were eluted at a flow rate of 2 mL.min’^ using a tertiary methanol-acetonitrile-H20 gradient as previously described. Radioactivity was detected at the exit of the column by a Flow One/Beta A500-Packard system. Metabolites of [1-^^C] arachidonic acid were identified by comparing their retention times with those of synthetic standards and quantitated by measuring the areas under the respective peaks. Each peak count was adjusted relative to the total radioactivity of the corresponding sample. For each incubation, the amount of every metabolite was compared with the control assay. RESULTS AND DISCUSSIO N In the presence of HAP or FAP, synovial proliferation slows over time (Fig. 2). However, under our experimental conditions, no cytotoxic effects were observed when particles were present. Transmission electron microscopy confirmed the nontoxicity of the two materials, as there was no spontaneous detachment of cells, and human synoviocytes presented normal cytoplasm and nucleus. Human OA type B synoviocytes spontaneously synthesized IL-la yMch increased over time (Table I). The presence of HAP and FAP inhibited II-la synthesis and II-la levels after one week were decreased compared to control cultures: 75% and 67% for HAP and FAP respectively (Table I). Moreover, by one month incubation, II-la synthesis dramatically declined compared to control cultures: 86% and 83% for HAP and FAP respectively (Table I)- In agreement with the results of Boimet et al. [11], we observed the formation of metabolites issued from the lipoxygenase pathways after incubation of 10^ cells with 1 \iC\[1-^^C] arachidonic acid in the presence of calcium ionophore A23187 (Fig. 3 A).
Interactionsof Bioceramieson Human Osteoarthritis(OA) Type B Synoviocytes:B. Liagre et al.
4S
A«gl«
SO
{19}
SS
115
to 0»vs
Figure 1: Xray difGraction pattern of HA P powder.
Figure 2: Proliferation of human OA type B synoviocyte s exposed to HA P ( -M- ) and FAP ( - x- ) particles; control culture ( - o ).
After one month, HA P and FAP particles did not seem to influence the formation of LTC 4 and LTE4. In fact, a marked decrease in HETE s (15-, 12- and 5-HETE) , LTB4, its isomer A6-/ran5^LTB4, and LTC 4 ^^^ ^^^’LTE 4 was not detecte d (Fig. 3B and 3C). In conclusion , using our experimenta l conditions, HA P and FAP decrease the secretio n of IL-l a and leukotrien e and HET E synthesis in human OA type B synoviocyte s in culture in a similar manner. Macrophages, which are representativ e of an inflammatory cell type, can only be used in short term invitrostudies. Our model, because cell culture is possible, has the advantage of prolonged exposure of human OA type B synoviocyte s to powders. This model will allow us to study the possible inflammatory response (arachidonic acid metabolism , proinflammatory cytokines) that can be induced by any biomaterials used in orthopaedics . One must be prudent in making assumptions about the invivopropertie s of a biomaterial based on a simple extrapolatio n of/>iv//ro results. Table 1. IL-l a release d by human OA synoviocyte s exposed to HA P and FAP particles by one week or one month incubation. Stimulus^
D - l a (pg/mL)b
Unstim. (1 week)*
6.33–1.5 8
Unstim. (1 month)*
13.73–3.2 9
106 HA P (1 week)*
1.58–0.0 6
106 HA P (1 month)*
1.92–0.1 3
106 FA P (1 week)*
2,08–0.5 2
106 FA P (1 month>
2.33–0.5 7 ^Stimulus abbreviations : Unstim., spontaneou s release by unstimulate d human OA type B synoviocytes ; 10<5 HA P or 10^ FAP, 10^ HA P or FAP particles/10 ^ ceUs. ^EL-l a concentratio n is expresse d as pg/mL/10^ cells. ** P < .05 relative to unstimulate d group, paired Student / test).
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8
s
\Q
u
iQ , u
k^ ia
u
ifwtr’-^ 4a
Figure 3: Extended scales of radioactive profiles of human OA type B synoviocyte s incubated with [1-^^C] arachidonic acid. A: referenc e profile; B: with HA P particles for one month; C: with FAP particles for one month. Metabolites identifie d were A6-/>-anj^LTB4 (peak 1), LTB 4 (peak 2). 15-HETE (peak 3), H-HET E (peak 4), 5-HETE (peak 5). LTC 4 (peak 6), LTE 4 (peak 7). Peak 8 is the peak of [1-^^C] arachidonic acid. REFERENCE S Chamley, J., London, Churchill Livingstone, 1972, p 9. 1. 2. Evans, E.M., Freeman, MAR. , Miller, AJ., and Vernon-Roberts, B.J., J, Bone JointSurg,,1974,56B , 626-642 . Goldring, S.R, Schiller, AL. , Roelke, M., Rourke, CM. , ONeill, D.A., and Harris. W.H., J. BoneJointSurg.,1983, 65A, 575-584 . M c Guire-Goldring, M.B., Meats, J.E., Wood, D.D., Ihrie, E.J., Ebsworth, N.M., and Russel, K.G.G.,Arthritis Rheum.,1984,27 , 654-662 . 5. Henderson, W.R.,yim. Rev.Respir. Dis., 1987,135,1176-1185 . 6. Charissoux, J.L., Najid, A , Cook-Moreau, J., Setton, D., and Rigaud, M., Clin. Orthop. Rel Res.,1996.326 , 259-269 . 7. Tawara, T., Shingu, M., Nobunaga, M., and Naono, T., Inflammation, 1991,15 , 145-157 . Labat. B.. Chamson, A . and Frey, J., J. Biomed. Mater.Res., 1995,29,1397 1401. 9. Downes. S., Clifford. C.J.. Scotchford. C , and Klein. C.P.AT.,7. Biomed. Mater.Res,,1995.29 , 1053-1060 . 10. Estensen. R.D.. White. J.G.. and Holmes. B., Nature,1974,248 , 347-348 . 11. Bonnet, C . Bertin, P.. Cook-Moreau. J.. Chable-Rabinovitch . H., Treves, R , and Rigaud, M.. Prostaglandins, 1995, 50, 127-135 .
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
A LON G TER M IMPLANTATIO N HYDROXYAPATIT E
OF CULTURE D
BON E
IN
POROU S
T. Yoshikawa^’^, H. Ohgushi^, H. Nakajima^ ’^, M. Akahane^, S. Tamai^ and K. Ichijima^ Departments of Pathology^ and Orthopaedic Surgery^, Nara Medical University, Kashihara City Nara 634, JAPAN.
ABSTRACT : In vitro bone tissue in porous areas of hydroxyapatite (HA) can be constructed by marrow cell culture treated with dexamethasone (Dex). To investigate in vivo osteogenic potential of the cultured bone for a long tenn period, the composites of HA and Fischer rat marrow cells were cultured in a Eagle’s Minimum Essential Medium containing 15%fetal bovine serum and Dex for 2 weeks, and were subcutaneously implanted into 7-week-old male syngeneic rats. The implants were harvested 52 weeks (one year) postimplantation and examined histologically. Almost all pores were filled with lamellar bone and some porous areas showed the appearance of regenerated bone marrow. Fluorochrome labeling (calcein) administered 50 weeks postimplantation was seen near the surface of bone in porous areas of the HA. These resuks indicate the persistent in vivo osteoblastic activity of the in vitroprefabricated culture bone in HA. KE Y WORDS : Hydroxyapatite, Bone marrow cell, Dexamethasone, Bone graft INTRODUCTIO N Bone like tissue can be produced by culturing bone marrow cells in a medium containing dexamethasone (Dex) and P-glycerophosphate, as previously reported [1-6]. Hydroxyapatite ceramics (HA) has been used as bone graft substitutes which can support osteoblastic differentiation [7-9] and we recently reported that HA combined with this cultured bone tissue possesses a high osteogenic ability when grafted in v/vo[10-12]. Therefore, the composite (HA and cultured bone) can be expected to be a bone grafting substitute having high osteogenic ability. The high osteogenic ability was confirmed by biochemical analysis. High alkaline phosphatase activity, which indicates osteoblastic activity, in the harvested composites began to appear at 1 week postimplantation and was maintained until 8 weeks. Bone Gla protein, an specific product of osteoblasts, began to increase at 1 week, followed by a steady increase until 8 weeks [8]. These findings indicate that this composite in in vivo situations can be categorized as the hybrid artificial organ. To confirm the ftinction of the composite as organ, the composite should maintain its activity for a long term period. In present study, we examined the osteogenic capacity of Dex-treated composites after one year postimplantation. 117
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Materials and Methods Materials; Coralline hydroxyapatite ceramic (HA) was used in this experiment. A hydrothermal chemical exchange converts the original calcium carbonate exoskeleton of coral into a completely inorganic replica of HA. The solid and porous components of the microstructure are completely interconnected. The pores average is 600 micrometers in diameter and their interconnections average is 260 micrometers in diameter. Interpore International (Irvine, CA) manufactured the HA and labeled it as Interpore 500. We used the cubic structure (5x5x5mm) of the HA. Marro w cells preparation and Culture; Marrow cells were obtained from the bone shaft of femora of Fischer 344, male, 7-week-old rats. Both ends of the femora of a rat were cut away from the epiphysis, and the marrow was flushed out by using 10 ml of culture medium expelled from a syringe through a 21 gauge needle. The released cells were collected in two T-75 flasks (COSTAR) containing 15 ml of the below-mentioned standard medium. Subsequently, the medium was renewed three times a week. Cultures were maintained in humidified atmosphere of 95% air with 5 % CO2 at 37 C. A standard medium consists of an Eagle-Minimal Essential Medium (MEM) containing 15% fetal bovine serum ( ICN BIOMEDICALS JAPAN Co. Ltd. ) and antibiotics ( 100 units/ml Penicillin, lOOj^g/ml Streptomycin, and 0.25|ig/ml amphotericine B, Sigma ). After 10 days in primary culture, marrow stromal cells were released from their culture substratum using 0.1% trypsin. The cells were concentrated by centrifiigation at 900 rpm for 5 minutes at room temperature, resulting in 10^ cells/ml. The HA blocks were soaked in 4 ml of the marrow cell suspension (10^ cells/ml). After 2 hours of incubation, each block was transferred into a 24 well-plate ( FALCON ) for subcultures. One HA block was subcultured in one well with 1 ml of the standard medium supplemented with lOmM Na (i-glycerophosphate (MERCK), 82 jUg/ml vitamin C phosphate ( L-Ascorbic Acid Phosphate Magnesium Salt n-Hydrate, C6H609PMg2/3.nH20, WAKO PURE CHEMICAL INDUSTRIALS ) and lO’^M dexamethasone (Dex, Sigma) [1,8]. The medium was renewed three times a week and the subcultures were maintained for 2 weeks. Implantation; Six composites after 2 weeks of subculture were implanted subcutaneously at 6 sites of the back of each Fischer 7-week-old male rat. These implants were harvested at 52 weeks (one year) postimplantation and prepared for histological analysis. To determine the bone dynamics in the composite, some rats were intramuscularly given one dose (15 mg/kg) of calcein at 50 weeks postimplantation and prepared for undecalcified histology at 52 weeks. Histological evaluation; The implants were harvested 52 weeks postimplantation, and were fixed in 10% buffered formaline, decalcified (K-CX solution, Fahna Co., Tokyo) and stained with hematoxylin and eosin. For undecalcified sections, the implants were immediately fixed with 70% ETOH and stained with Villanueva bone stain. These were dehydrated in an alcohol series, defatted, embedded in methylmetacrylate, and observed under light microscopy or fluoromicroscopy.
A Long Term Implantationof CulturedBone in Porous Hydroxy apatite:T. Yoshikawa et al.
119
Figure 1. The composite graft 52 weeks after implantation. White area shows the ghost of hydroxyapatite ceramic produced by decalcification procedures. Black area indicates the new bone appeared in porous area and M indicates regenerated marrow tissue. (Decacified sections stained with hematoxylin and eosin. x80)
Figure 2. The composite 52 weeks postimplantation. Undecalcified section under fluoroscopy. Arrowheads indicate the interface between bone and hydroxyapatite. B indicates bone and C indicates hydroxyapatite ceramic. Arrows indicate the calcein labeling, administered 50 week after implantation. (Villanueva bone stain, xlOO)
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RESULT S AND DISCUSSIO N Fifty-two weeks after implantation of the composites, almost all pores were filled with bone and some porous areas showed the appearance of regenerated bone marrow in association with the new bone formation [Fig.l]. Not only resting osteoblasts but also active osteoblasts could be detected on the surface of bone. All of Dex-treated composites (12/12) showed bone formation. By fluorochrome labeling, the line of calcein administered 50 weeks postimplantation could be detected near the surface of the bone in many porous areas [Fig.2]. These results indicate that the composite of cultured bone and porous hydroxyapatite maintains the osteogenic ability for along time period and, therefore, the composites can fimction as hybrid artificial organ. The findings clearly showed that the tissue culture technology applied in present experiments can stimulate the inherent osteogenic ability of marrow stromal cells is pore regions of HA and importantly the ability persists for a long term period. Based on the fact that human cultured marrow cells possess osteogenic potential[13-15], the present graft procedures show the possibility in clinical application of reconstructive surgery.
REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14. 15.
Maniatopoulos C , Sodek J. and Melcher A.H. Cell TissueRes., 1988, 254, 317-330. Yoshikawa T., Ohgushi H., Dohi Y., and Davies J.E. Bio-Med.Mater. Eng., in press. Ohgushi H., Dohi Y., Katsuda T., Tamai S., Tabata S. and Suwa Y., J. Biomed. Mat. Res., 1996, 32, 333-340. Ohgushi H., Dohi Y., Yoshikawa T., Tamai S., Tabata S., Okunaga K. and Shibuya T., Biomed Mat. Res., 1996, 32, 341-348. Yao K-L., Todescan R.J. and Sodek J., J. Bone and Mineral Res.. 1994, 9,231-240. Davies J.E., Chemecky R., Lowenberg B. and Shiga A., Cells Mater. 1991, 1, 3-15. Ohgushi H., Goldberg V.M., and Caplan A.I. J.Orthop.Res.,1989, 7 568-578 Yoshikawa T., Ohgushi H., Okumura M., Tamai S., Dohi Y. and Moriyama T. Calcif. TissueInt., 1992,50, 184-188. Okumura M., Ohgushi H., Yoshikawa T. and Tamai S., Ceramics in Substitutiveand ReconstructiveSurgery, 1991, 353-361. Yoshikawa T., Ohgushi H., and Tamai S., J. Biomed Mat. Res. 1996, 32, 481-492. Yoshikawa T., Ohgushi H., and Tamai S., BioceramicsVol. 8, 1995, 421-426. Yoshikawa T., Ohgushi H., Akahane M., Sempuku T., Tamai S. and Ichijima K., Bioceramics Vol. 9, 1996, 65-68 de Bruijin J. D., vd Brink I. and Bovell Y. P., BioceramicsVol.9, 1996, 45-48. Haynesworth S.E., Goshima J., Goldberg V.M. and Caplan A.I., ^o«^, 1992,13, 8188. Ohgushi H. and Okumura M.,Acta. Orthop.1990, 61,431-434.
SURFACE BEARING CERAMICS
This Page Intentionally Left Blank
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
OXID E CERAMIC S FOR ARTICULATIN G COMPONENT S OF TOTA L HI P REPLACEMENT S G. Willmann CeramTec Medical Product Division, P.O.B. 1149, D-73201 Plochingen, Germany
ABSTRAC T The state of the art of ceramics for total hip repalcement (THR) is reviewed, reporting some tribological facts and future trends. Modem hip systems are designed in a modular way. By using the concept of cup inserts the surgeon has the option to use either a cup insert of PE-UHMW or an alumina cup insert formetal backed socket. The wear couple alumina-on-alumia offers extremely low wear rates when approved components are used. The combinations alumina-on-zirconia and zirconia-on-zirconia are not approved. WEA R COUPLE S USED IN THR Alumina ceramics were introduced more than 20 years ago as a material for bearing surfaces in total hip replacements (THR) [1]. The objective was to minimise the wear debris to overcome the problem of osteolysis caused by the polyethylene wear particles. Osteolysis is still a problem in orthopedics. To minimise wear of the articulating surfaces in THR there are two approaches classified as "hard / soft" and "hard / hard" (table 1), both offering the option to minimize wear. Approved wear couples in total hip replacement [1-7]
Table 1. Femoral head
socket
remark
typical wear
cobalt chromium
PE-UHMW
standard
0.2 to 0.5 mm per year
almnina
PE-UHMW
standard
< 0.1 mm per year
zirconia
PE-UHMW
new approach
alumina
alumina
used since 25 years
alumina
CFRP
5 |im per year (clinical results) < 1 ^un per year (HIPed alumina) a few ^m per year
123
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One option is to use ceramic articulating against polyethylen e (PE-UHMW) , the better one is the wear couple alumina-on-alumin a [6,7]. The standard combination for THR s is a metal femoral head articulating against an acetabula r cup of polyethylen e (PE-UHMW ) The wear rate for the mostly used combination metal / PE is > 0,2 mm, sometime s up to 0,5 mm / year [9-11], Biolox forte alumina heads articulating against polyethylen e (hard / soft) have a lower wear rate. It is below 0,1 mm / year [9-11]. Dueto that the loosening rate of stems is lower, too. Since 1974 more than 2 million alumina component s have been successfiill y used. The mechanica l propertie s of Y-TZ P zirconia are very atractive, e. g.fracturetoughness and bending strength. This offers the option to design heads with diameters smaller than 28 mm. Femoral ball heads of Y-TZ P were introduced . Alumina and zirconia femoral heads in combination with polyethylen e cups (hard / soft) have similar tribological properties . In respect to wear Y-TZ P offers no adavantages . Heads of Y-TZ P are more expensiv e than alumina. The combination alumina femoral head articulating against an alumina cup or cup insert (hard / hard) has the lowest wear rate known. Clinical results are < 5 pm / year [2,6]. There are lots of investigation s on Biolox forte femoral heads articulating against Biolox forte acetabular cups or cup inserts. Histological investigation s of retrieve d Biolox implants prove that the wear particles have a diamete r in the order of 1 ^m [6 ]. Alumina particles don’t cause any problems, they are biocompatible . The investigatio n of retrieve d Biolox implants (in vivo up to 20 years) prove that the wear rate is extremel y low, e.g. lessthan 5 fim per year [2,6]. Ring-on-disc tests and simulator test prove that the combination Y-TZ P ceramic / Y-TZ P ceramic and Y-TZ P zirconia / alumina are disastrous [4,12] . Table 2.
Approved and not approved wear couples in THR
Femoral head
socket
apporved
remarks
alumina
PE-UHM W
yes
standard
Biolox forte (HlPed alumina)
Biolox forte
yes
Biolox forte (HIPed alumina)
any alumina
no
any alumina
Biolox forte (alumina)
no
Y-TZ P zirconia
PE-UHM W
yes
Y-TZ P zirconia
alumina
no
desastrous wear
Y-TZ P zirconia
Y-TZ P zirconia
no
desastrous wear
} } }
see comment on next page
Oxide Ceramicsfor ArticulatingComponentsof Total Hip Replacements:G. Willmann 125
Table 2 shows a list of wear couples that are apporoved or not approved. It is important to draw the surgeons’s attention to the fact that only approved wear couple shall be used. Approval is a legal problem. A system or a wear couple will only be apporved if tests have been performed and if reports about good results are available. In the USA there are very strict regulations for that, in Europe the European Community passed directives for evaluationg the risk of a medical divice. Table 2 shows a list of wear couples that are apporoved or not approved. The good clinical results for the wear couple alimiina-on-alumina are only achieved if there is a gap between the head and the socket. Due to that lubrication is possible. There is some kind of standardization for the diameters of the ceramic femoral heads, but no standardization for the tolerances of the diameters. Therefore components offered by different companies shall not be combined. It may well be that due the bad tolerancesfrictionof the wear couple may be too high. If a surgeon combines a not approved wear couple he will be responsible for quality assurance. That is law in most countries. The advice has to be: NEVER MIX AND MATCH. RELIABILIT Y The regulations of the American Food and Drug Adminisatration and the directives of the European Community state that a medical device shall be safe, i. e. the risk due tofracturehas to be minimized. Ceramics are brittle and the probabilty of failure may be higher than for metal components. Today’s ceramic heads and cup inserts have a very low risk of failure due to fracture. Statistics prove that thefrtacturerate is in the oerder of 0.01 % [1,13]. The revision rate due to failiu-e of a ceramic compomnent is much lower than the one for septic or aseptic loosening, see table 3. This result is based on more than 1.5 million Biolox heads. To improve reliabilty in THR modem alumina is hot isostatic pressed (HIP) to maximize density and to minimize the grain size of the mircostructure to below 2 ^un. Both materials properties are correlated to mechanical strength. Heads and cups are laser marked to eliminate critical flaws. Using a 100 % proof - test only heads and cups that withstand fracture loads far above maximum physilogical loads which is in the order of 6 to 8 times body weight [14]. Table 3:
Revision rate in THR
Revision rate
due to
to 10 %
aseptic loosening
about 1 %
septic loosening
0.02 %
fracture
of a ceramic component
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Bioceramics Volume10
Figure 1.
Modular designed hip system with the option to use metal-on-polyethylene , alumina-on-polyethyne , zirconia-on-polyethylene , and alumina-on-alumin a
FUTUR E TREND S For systems in total hip replacemen t the future trend is then wear couple hard / hard [6,7]. Modem hip systems are designed ikn a modular way, see figure 1. By using the concept of cup inserts the surgeon has the option to use either a cup insert of PE-UHM W or an alumina cup insert for metal-backe d sockets. REFERENCE S 1. Clarke, I., G. Willmann in: Cameron, H .U .(ed.) BoneImplant Interface (1994) 203 - 252 2. Henssge, E .J., I. Bos, ans G. Willmann J. Mat. Sci.Mat. Medicine 1994 5. 57 - 661 3 Willmann, G.,H. J. Friih, and H. G. Pfaff Biomaterials 1996 17. 2157 - 2162 4 Friih, H. J., G. Willmann, H. G. Pfaff Biomaterials 1997 18. (in press) 5 Friih, H.-J., G. Willmann Biomaterials (in press) 6. Puhl. W. (ed.) Die Gleitpaarung BIOLOXinderHuftendoprothetik Enke Verlag (1996) 7. Puhl, W. (ed) Performace of theWearCoupleBIOLOXforte in TotalHip Arthroplasty Enke Verlag (1997) 8. Willmann, G. Biomed. Technik 1997 42. (in press) 9. Weber, B. G., Th. Fiechter Orthopadie 1989 18. 370 - 376 10. Zichner, L. P., H.-G. Willert Clin: Orthop: Rel:Res 1992 282. 86 - 94 11 Th. Lindenfeld Der Orthopade 1997 (in press) 12. Oonishi, H., M . Ueno, H. Okimatsu, and H. in: Kokubo, T., T. Nakamura, F. Miyaji (eds.) Bioceramics 9 (1996) 503 - 506 13. Willmann, G. Mat. wiss. Werkstofflechni k 1996 27. 280 -286 14. Richter, H. G. and G. EWillmann in [6] 77 - 83
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the10thInternational Symposium on Ceramics in Medicine, Paris,France,October1997) '1997 Elsevier Science Ltd
EX VIV O AND IN VITR O ANALYSI S OF THE ALUMINA/ALUMIN SYSTE M FOR HEP JOIN T PROSTHESE S
A BEARIN G
HJ. Refior*, W. Plitz**, A. Walter** *Orthopadische Klinik, Klinikum GroBhadern, MarchioninistraBe 15, D-81377 Munchen **Labor fur Biomechanik und Experimentelle Orthopadie, Klinikum GroBhadern, MarchioninistraBe 23, D-81377 Munchen
ABSTRAC T The increased awareness of polyethylene debris as one of the main reasons causing osteolysis around implants raised the expectation of a renaissance of the ceramic/ceramic bearing system. Simulator tests result in linear wear rates below 1 ^m per million cycles and per component corresponding to a gravimetrical wear estimation of less than 1 mg per million cycles for a current alumina quality. These improved alumina materials show an advantage for linear wear of factor 5 at minimum compared with quality levels ten years ago. Contrary to simulations, collections of up to date more than 50 retrieved alumina/alumina implants have shown that linear wear rate frequency maxima exceed by 2 orders of magnitude those couples which were not exposed to loosening artifacts. The wear rate probability is not normally distributed and arithmetic means do not seem adequate to describe or compare invitro and ex-vitro results. KEYWORD S Tribology, hip joint replacement, ceramics INTRODUCTIO N Endoprosthetic loosening is increasingly considered to be the consequence of wear dependent osteolysis in the anchorage area of the implant, even though there are less secured findings about the biologic effects of the characteristics of wear particles. This is obviously valid for ceramic/ceramic couplings, too. Considering the increasing knowledge \vith regard to wear problems of polyethylene components a renaissance of the hard selfpaired s}’Stem was announced at the beginning of the nineties. In this respect there may be pointed out that chemical reactions are not to be expected in case of pure alumina contrary to the reaction products of metals. MATERIA L Subject of the investigations were 54 revised as well as 11 brand-new alumina/alumina implants (Table 1). The alumina quality of the presented revised ceramic couplings varies according to the state of the art in the deliver}^ year as well as the different manufacturers. The minimal quality requirements are to be found in the 1981 edition of the International Standard 6474. In the course of a revison of this International Standard the upper limit for the avarage grain size was reduced from 7 ^m to 4.5 ^im (ISO 6474 2nd edition). 127
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Table 1:
Material, density and average grain size Number [n] 54
Dcnsitiy fg/cm^l 3.86-3.97
Revision samples 1997-1991 (different manufacturers) Rosenthal, Feldmiihle/Cerasiv, Ceraver 1 Simulator Specimens 1986 3.94 - 3.96 9 (Biolox I, Feldmiihle) 1 Simulator Specimens 1996 3.98* 2 1 (Biolox forte, CeramTec) * typical value according to the manufacturer (CeramTec)
Mean Grain Size 1 fum] 2.2 - 16.4 1 4.1-4.9 1.8*
The material and design quality of 9 brand-new ceramic components were tested in the hip joint simulator corresponded to the state of the art of the year 1986. The grain sizes were at the maximum limit of the actual standard. For the sake of comparison two further tests were performed under identical test conditions using ceramic couplings recently manufactured (Biolox forte, CeramTec). By means of hot isostatic pressing after sintering it is possible to achieve a lower avarage grain size while improving density and strength at the same time. Presenting a mean grain size of 1.8 fim and a densitiy of 3.98 g/cm^ the new material greatly exceeds the standard requirements, METHOD S The tribologic tests (Table 2) were performed in a hip joint simulator type Miinchen I. The three axial movements are controled electronic-hydraulically within mechanically defined limits so that certain contact positions of the gliding components are not permanently exact reprocduced. The stress pattern with its PAUL type double peak simulating normal walking deserves rather critical consideration according to recent results of telemetering total hip prostheses. This fact does also refer to the use of distilled water as lubricant, since compared to solutions of bovine serum and physiological salt it may represent a more severe condition for degradation of polluted ceramics due to the corresponding metals. A turntable device (Taylor-Hobson TR-30 PC) was used for profilometric wear quantification of simulator specimens as well as retrieved samples. These linear wear data of the simulator specimens were completed with planimetry (MOP AM02) of abrasion areas that seemed to be macroscopically roughened and with gravimetric measurements (Sartorius Balance 1712 MP8). In the case of retrieved objects mean grain size (SEM) and chemical structure (AAS) was analyzed where the X-ray documentation refering to positioning or migration of the prosthetic component respectively made it possible. Table 2: Test conditions in the hip joint simulator type Miinchen I Movement: Loading:
Frequency: Lubrication Temperature:
3-axial (frontal 12 / sagittal 45 / transversal 14 ) double peak 1. Maximum 2500 N 2. Maximum 3000 N Preload 300 N 0.9 Hz aqua dest. 37 – 2 C
Analysis of the Alumina/AluminaBearing Systemfor Hip Joint Prostheses:HJ. Refior et al.
129
SIMULATO R RESULT S OF EARLIE R QUALIT Y CERAMIC S The abrasive character of the dominant wear mechanism is \vithout doubt at high wear rates III. Peak loads induce initial grain excavations getting between the gliding surfaces and may cause avalanche like increasing wear due to 3-body wear. From the sphericity deviations and the planimetric values of the abrasive areas before and after each test interval it is possible to estimate mathematically a wear volume. In case of small abrasive quantities this volume calculated from dimensional data is up to 90 % below the volume that can be estimated by means of weight loss measurements. One cause coming into question is the general insuflFiciency of the profilometric assessment method to register concentric form changes if wear mechanisms result in a regular material loss as shown for relief polishing documented in SEM investigations. Such large area wear with maximum depth of fractions of micrometers was neither profilometrically reliably to register nor macroscopically detectable. Nevertheless, gravimetric control measurements showed at these relief polishing phenomena a wear rate up to 1.5 mg per million cycles. Contrary to a simulator test under normal test conditions wear may increase more than hundredfold in case of a singular harsh repositioning even without simulation of subluxation. On the other hand, a high angle positioning of the cup from 45 to 55 causes a wear increase rate of only 25 to 65 per cent. Overcritical sphericity deviations 111 do obviously have a much greater influence on wear rates increase than such a suboptimal inclination angle. The effects of the above mentioned conditions for peak loads may superimpose additively or multiplicatively and should be used for the interpretation of clinical wear phenomena. At present, due to the lack of a consistent testing concept, it is impossible to present a quantitative estimation of loosening dependent wear for explants that were revised because of loosening mainly. Nevertheless, laboratory tests may serve as guiding data and for 9 specimens they provide an avarage wear rate for alumina selfpaired couplings of 3.8 and 1.9 micrometers per million cycle for the ball and cup component respectively. But for the explanation of sometimes contradictory wear values of revision objects the conclusion seems much more improtant that with an increasing number of cycles wear rates are not constant but they descent down to the limits of the available measuring device. SIMULATO R RESULT S FOR OPTIMIZE D ALUMIN A Compared to Biolox I the selfpaired Biolox forte presents a determined measurable improvement of linear wear by factor 5 and if taking considering different error possibilities by a factor over 10 /3/. The maximum measurable linear wear of a single component was at 0.9 [im per million cycle in the run-in phase, whereas after the second and third million of cycles no change could be registered any more by profilometry. Weight loss estimations resulted in a maximum wear volume of 0.2 mm^ for both components. Abrasion is the predominant wear mechanism again but contrary to the grain excavations of the earlier materials it is characterized by primarily transcristalline abrasion. RESULT S OF THE RETRIEVE D OBJECT S The collection of retrieved objects comprises 54 couplings of different design and an implantation time between 3 months and 13 years. The components were manufactured between the years 1975 and 1991. The wear rates that were established profilometrically range from 30 nm to 3.7 mm per year with a maximum frequency most appropiately described with the geometric mean value of 3.1 micrometer per year. The adequate arithmetic mean value in case of normal distribution would result here in a value of more than 120 micrometer per year and is influenced drastically by few high maximum values of the left inclined distribution. The largest group of the same design were the ceramic/ceramic couplings with a design after Mittelmeier. It comprised the component with 3.7 mm maximum wear in the cup rim noticed already in laboratory tests. It presented sharp-edged unpolished reliefs by means of which the potential contact areas of the bearing components are essentially reduced at small deviations of inclination angles of the cup from an ideal
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Table 3:
Range of estimated linear wear rates per component based on the changes of the sphericity deviations number
1 Revision samples 1997 -1991 (different manufacturers) Rosenthal, Feldmuhle/Cerasiv, Ceraver 1 Simulator Specimens 1986 (Biolox I, Feldmuhle) 1 Simulator Specimens 1996 1 (Biolox forte, CeramTec)
54
minimum
maximum
fjim]
0.03
3.7 X 10^
9
0.1
6.3
2
0.2
0.9
1
position. Wear rates of retrieved samples with cup positions below 55 slightly decrease from about 2 to 1.5 micrometer per year, the values for higher positions show the danger of excessive wear and shorter implantation time. Of course the final wear performance is a result of multifactorial processes. But it seems remarkable that all components with wear rates below 0.5 micrometer per year present mean grain sizes of less than 4 micrometer. DISCUSSIO N An overview of the range of all estimated linear wear rates is given in Table 3 demonstrating a large scattering of the data of the mainly because of loosening retrieved explants and the comparable narrow ranges of the simulator wear data which present the estimated wear rates of a run-in phase during the first million of simulated walking cycles. One of the most frequently mentioned cause for the reserve of the users of ceramic/ceramic couplings was besides the fracture possibility the supposed material specific sensitiveness of the system to critical inclination angles of the cup. Simulator results show that a steep cup position of 10 from the ideal positioning results in a wear increase of at maximum 65 per cent, but other peak load conditions are able to generate far more wear. Steep positioned cups of ceramic/ceramic revision objects seem to show clearly higher wear rates, but the same time these wear rates are usually overlayed with loosening artifacts difficult to be estimated. An explanation might be provided by comparing the frequency distribution of wear rates from simulator tests, autoptical explants and loosened revision samples. Based on the data provided by Henssge et al. 74/ in his investigations of autoptic ceramic components there may be derived a maximum frequency below 0.1 micrometer per year and a wear rate distribution coinciding with the corresponding simulator results the maximum frequency of which corresponds to that of the revision objects. Taking into account the design-dependent wear phenomena of explants it can be concluded that the improvements of the tribological behavior found in laboratory tests for components of hot isostatic pressed alumina will present a 5 or 10 times less wear rate in vivo only in case of an adequate cup design. LITERATUR E 1. Walter, A.: On the Material and the Tribology of Alumina-Alumina Couplings for Hip » Joint Prostheses. Clin. Orthop. 282:31, 1992 2. Walter, A., Bdhmer, F.: On the geometrical shape of artificial hip joint bearings with articulation of identical materials. Implant Materials in Biofunction, Advances in Biomaterials, Elsevier, 8:143, 1988 3. Walter, A.: Investigations on the Wear Couple Biolox forte/ Biolox forte and earlier Alumina Materials, Proceedings of the 2""* CeramTec Symposion 7/8.3.1997 in press 4. Henssge, E.J., Bos, I., Willmann, G.: AI2O3 against AI2O3 combination in hip endoprostheses. Histologic investigation with semiquantitative grading of revision and autopsy cases and abrasion measures. J. Mater. Med. Sci. 5:657, 1994
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
HYBRI D ALUMINA-ALUMIN A HI P REPLACEMENT : A SURVIVORSHI P ANALYSI S AN D RESULT S AT A MINIMA L FIV E YEA R FOLLOW-U P M. HAMADOUCHE, P. BIZOT, R.S. NIZARD, and L. SEDEL Department of Traumatology and Orthopaedic Surgery, Lariboisiere Hospital , 2 rue Ambroise Pare 75010 Paris, France Laboratoire de Recherches Orthopediques. URA CNRS 1432, Faculte de Medecine Lariboisiere-St Louis, 10 Ave de Verdun, 75010 Paris, France
ABSTRAC T The survivorship of an hybrid alumina-alumina total hip arthroplasty has been evaluated in a series of 55 patients (62 hips). A press-fit bulk alumina socket and a cemented titanium alloy stem were implanted. Bearing surfaces were a 32 mm alumina head articulating with the alumina socket. Four failures occured: 3 were reoperated for aseptic loosening of the socket and 1 for femoral head fracture. The overall survival rate at six years was 90.8% when revision was considered as failure. At a mean follow-up of 82.5 months, 95.2% hips were graded very good or good while 71.4% showed radiolucent lines around the socket. Socket fixation has to be improved. INTRODUCTIO N Since Charnley introduced metal on polyethylene (PE) total hip arthroplasty (THA) in the 1960s (1), an increasing number of papers reported on PE wear debris generating aseptic loosening and osteolysis mediated through biological reactions (2). Alumina on alumina bearing surfaces was proposed as an alternative to metal on plastic friction couple by P.Boutin in 1970 (3). Since this period, an improvement in the manufacturing process made alumina a reliable material. Until 1982 the acetabular component was always cemented using a conventional cementing technique, loosening of the acetabular cemented components (4,5,6). Loosening of the acetabular cemented components (4,5,6) was the main reason for failure. For this reason a short series of patients had a bulk alumina socket implanted with a press-fit technique (Cerapress). The aim of this study was to evaluate prospectively the survivorship of an hybrid alumina-alumina prosthesis using a press-fit cementless fixation in a limited number of patient. PATIENT S AND METHOD S From March 1982 to December 1990, 62 hybrid THAs were performed in 55 patients. The 32 mm head and socket component were made of surgical dense alumina ceramic. The femoral stem was made of anodised titanium alloy (TiA16V4). The socket was a bulk alumina cup with Inun grooves allowing bone ingrowth. Of the 55 patients, 36 were male (65.5%) and 19 female (34.5%). The average age at the time of hip replacement was 49.6 years (range 22.5 to 75.4). The average weight was 75.4 Kg (50 to 108) and height 172.5 cm (154 to 193). The right hip was operated on 25 patients and the left hip in 23, while 7 had bilateral replacement. Initial diagnosis is presented in Table 1. THA was the first surgical procedure for 43 patients while 12 had previous hip surgery procedures. All arthroplasties were performed by the senior author (L.S.) through a posterolateral 131
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approach in all cases except for one lateral approach with trochanteric osteotomy. Allograft in one patient and autograft in seven were used to reconstruct an acetabular deficiency. Table 1. Initial diagnosis Underlying disease Primary osteoarthrosis Avascular necrosis Hip dysplasia and CDH Posttraumatic 1 osteoarthrosis Ankylosing spondilitis 1 Other coxopathy
1 TOTA L
Figure 5. Bilateral THA at 3 and 11 years
Number of hips 23 21 7 5 3 3 62
% 1 37.1 33.9 11.3 8.1 4.8 4.8
100 1
Systematically, the stem was cemented using a second generation technique.The socket component was impacted in a 2 nmi under-reamed acetabular cavity (Cerapress). Clinical ratings (Merle d’Aubigne-Postel grading system PMA) were determined pre- and postoperatively (7). On the immediate post-operative and at the latest follow-up anteroposterior radiographs, several parameters were investigated on the socket side: inclination angle of the cup, evolution of radiolucent lines in six zones of the socket and vertical or horizontal migration (8). RESULT S 13 patients (14 hips) were lost to follow-up. 25 patients (28 hips) were examined and 12 patients (14 hips) were interviewed by telephone at least five years after THA (mean 82.5 months, range 59.9 to 144.7). One patient with bilateral hip replacement died two years after THA from septicaemia. At the latest follow-up, both sides were graded excellent. Four revisions were documented in this series: one for femoral head fracture and three for aseptic loosening of the socket. The femoral head fracture occured 34.8 months after THA and remained unexplained: seven days before fracture the hip was graded excellent and no radiographic changes were noticed. Revisions for aseptic loosening occurred respectively 36.3, 53.1 and 68.4 months postoperatively. The mean PMA was 14.5 (range 12 to 17). Two patients had a patent socket migration and one was operated for permanent pain. The remaining 37 patients (42 hips) had a mean preoperative PMA of 11.2 – 2.2 (3.6 for pain, 3.5 for walking ability and 4.1 for motion) Vs 17.1 – 1.8 (5.5 for pain, 5.5 fro walking ability and 5.8 for motion) posoperatively (p< .0001). No radiological modification was noticed on the femoral side. On the socket side 4 hips (9.5%) showed migration of the acetabular component but had good clinical results except for one. 25 hips out of 37 (71.4%) showed a two-zone radiolucent line or more, 6 of them being complete (5). There was no correlation between the presence of radiolucent lines and clinical results. Neither osteolysis nor bone loss was observed even in case of cup migration or aseptic loosening. The survivorship analysis was performed using revision and revision for aseptic loosening of the socket as failure . The survival rates were respectively 90.8% (82.1% - 99.5%) and 92.7% (84.7% 100.7%) at six years as shown on fig. 1 and 2. An univariate analysis was then performed on sex, age (fig.3), initial diagnosis (fig.4), previous hip surgery, position of the socket. None of these parameters did significantly influence the results.
Hybrid Alumina-AluminaHip Replacement:A SurvivorshipAnalysis: M. Hamadoucheet al. Figure 1. THA survivorship. Revision considered as failure ^99.5 %
100 p
-o
, 90.8 %
80 4
82.1 %
I
"f
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Mean Surv. 95% Inf. (Surv.) 95% Sup. (Surv.)
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30 40 50 60 Postoperative months
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70
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80
Figure 2. THA survivorship. Revision for aseptic loosening considered as failure. 1.007% 100
1 = ^ ^
80 i
84.7 %
T3
> 60 i Mean Surv.
^
40 i
95% Inf. (Surv.) 95% Sup. (Surv.)
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30 40 50 60 Postoperative months
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Figure 3. THA survivorship. Revision for aseptic loosening considered as failure for patients older than 50 years of age and younger than 50 years of age. 100-f 80
"" ^^^
> 60 4
^
40
Mean Surv. Age< 50 years old * Mean Surv Age
50 years old
20 4 0 4 1
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40 50 60 70 Postoperative months
80
133
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Figure 4. THA Survivorship. Revision for aseptic loosening considered as failure for patients with primary ostearthrosis and for patients with other initial diagnosis 95.1 % 89.8 %
100 4 801 0) C/3
> 60
* Mean Surv. Primary osteoarthrosis
0^
o c
40 4 Mean Surv. Other initial diagnosis 20 4 0 4 -| ’ I
0
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^"i ’ I "-n ’ I
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20 30 40 50 60 Postoperative months
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-
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DISCUSSIO N AND CONCLUSIO N Alumina-alumina combination has proved to generate 4,000 times less particles than metal-PE, wear debris being considered as bioinert (3). However, cemented fixation of the socket was not reliable enough. In order to improve this situation, a cementless bulk alumina socket was used. The fracture of an alumina implant remains a matter of concern, but it must be noticed that in our experience this one was the fourth in a 17 years period on more than 1,500 alumina-alumina hip replacement. In this case of cementless cup fixation, a micromotion is probably tolerable as demonstrated by the absence of correlation between radiolucent lines and clinical data. Despite the relatively large number of failures, the overall revision rate was comparable to metal-PE combination. The absence of osteolysis could be attributed to the lack of wear debris. Nevertheless, the initial anchorage of the bulk alumina acetabular component must still be improved. REFERENCES 1. Charnley J., Springier Verlag 1979. 2. Savio J.A. Ill, Overcamp L.M. and Black J., Clin.Mater., 1994, 15, 101-147. 3. Boutin P., Christel P., Dorlot J.M., Meunier A, de Roquancourt A., Blanquaert D., Herman S. and Sedel L., J.Biomed.Med.Res., 1988, 22, 1203-1232. 4. Sedel L., Kerboull L., Christel P., Meunier A. and Witvoet J., J.Bone Joint Surg., 1990, 72-B, 658-663. 5. Nizard R.S., Sedel L., Christel P. Meunier A., Soudry M. and Witvoet J., Clin.Orthop.Rel.Res. 1992, 282, 53-63. 6. Sedel L., Nizard R.S., Kerboull L. and Witvoet J., 1994, 298, 175-183. 7. D’aubigne R.M. and Postel M., J.Bone Joint Surg., 1954, 36-A, 451. 8. DeLee J. and Charnley J., 1976, Clin.Orthop.Rel.Res., 1976, 121, 20.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
LO W TEMPERATUR E AGEIN G BEHAVIO R OF ZIRCONI A HIP JOIN T HEADS . J. Chevalier, J.M. Drouin, B. Cales Norton Desmarquest Fine Ceramics, ZI n l, 27025 Evreux Cedex, France
ABSTRAC T Low temperature ageing of zirconia ceramic hip joint heads (PROZYRfi, HIP) was analyzed. Transformation kinetics from 134 C down to 60 C allow an accurate prediction of ageing at body temperature : an incubation period of about 10 years followed by a slow surface transformation is expected. 25 years are required to obtain 20% monoclinic content. Wear simulations show 40% monoclinic content does not increase Polyethylene wear significantly. KEYWORD S zirconia, ageing, phase transformation, wear INTRODUCTIO N Surgical grade yttria-stabilized zirconia (Y-TZP) ceramics have been successfully introduced in orthopaedy for 12 years as hip joint heads. They present reduced polyethylene (PE) wear [1], which is one of the major concerns today in hip arthroplasty. Their mechanical properties are much better than recorded with alumina ceramics in terms of toughness and strength. This improvement in mechanical properties is related to the stress induced transformation from the tetragonal to the monoclinic phase around the cracks, conunonly known as transformation toughening [2]. Up today, about 300.000 zirconia hip joint heads have been successfully implanted, mainly in Europe and in the United States. In recent years, a surface tetragonal to monoclinic (t-m) transformation of yttria-stabilized zirconia has been recorded for so called low temperature ageing in humid atmosphere [3]. The origin of this mechanism is believed to be attributed to a reaction between water and Zr-O-Zr bonds of the ceramic [4]. This reaction has been clearly identified from 100 C to 500 C and occurs with a maximum rate at 250 C. Following these results, the long term stability of surgical grade zirconia ceramics has sometimes been questioned in the literature [5]. However, extrapolation of simple first order transformation kinetics from high temperature studies down to 37 C must be regarded with suspicion. Moreover, results of flexural strength on bending bars cannot be interpreted for hip joint heads because of the strong effect of machining on transformation kinetics [6]. It is the aim of the present paper to accurately predict the long term ageing behavior at 37 C of a surgical grade zirconia ceramic (PROZYRfi) from ageing experiments on hip joint heads between 134 C and 60 C. A first attempt is made to correlate surface transformation to wear behavior of polyethylene. 135
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MATERIAL S AND METHOD S Experiments were conducted on commercially available zirconia hip joint heads (PROZYRfi), processed from 3Y-TZP powders, sintered then hot isostatic pressed (HIPed) in order to increase their strength. The mean grain size calculated from the linear intercept method is 0.5 |im. Kinetics experiments were conducted using both steam sterilization at 134 C and 120 C and ageing in Ringer’s solution from 100 C down to 60 C, with femoral heads of different batches in order to obtain a good insight into manufacturing process. At least 3 heads of each batch were tested for each ageing temperature. Near surface monoclinic content was followed by X-ray diffraction (CuKa, penetration depth « 5 \xm).Surface transformation was also observed by means of an optical interferometer with height resolution of – 0.1 nm (X-Y resolution of 0.5 |j,m). Wear experiments were conducted on the same zirconia heads versus conunercial PE acetabular cups, coupling on a MTS hip simulator. Hip joints were submitted to physiological loading (maximum load 2500N) in a synthetic serum (Plasmionfi, Roger Bellon) containing 25g/l proteins. RESULT S AND DISCUSSIO N Figure 1 shows the monoclinic content measured by X-ray diffraction on the surface of the zirconia heads for one given batch of heads as a function of ageing in steam and in Ringer’s solution. It appears that the variation of monoclinic content versus time is perfectly described with sigmoidal curves showing that a simple first order kinetic, which was sometimes used in literature is unsuitable to fit Low Temperature Transformation (LTT) data. This also suggests that the Mehl-Avrami-Jonhson (MAJ) law [7] already used in the literature for higher temperatures by Tsubakino et al.[8], is relevant. MAJ law gives an expression of monoclinic content/versus time t on the form :
/ = l-exp-(6/) "
(1)
where n is a constant that must be determined for each zirconia ceramic and b gives the activation energy Q of the transformation mechanism by :
* = ^exp(^--^J
(2)
where T is the absolute temperature, A is a constant and R is the gas constant, n and b parameters can be deduced from a logarithmic plot of Equation 1 which defines a straight line of slope n and ordinate origin ln(b). It has been observed that n is independent of temperature and equal to 3+0.4. We must remember the MAJ equation, which was often used to describe time-transformation isotherms in alloy systems, assumes an incubation period, then nucleation and growth for n values between 3 and 4. Optical interferometer effectively revealed tm transformation proceeds by nucleation of monoclinic sites from one grain and growth to the neighboring grains after an incubation time dependent on temperature. Nucleation was detected by small isolated spots of about 1 nm height at the surface of the heads corresponding to the volume expansion caused by transformation. Growth was observed by an increase of these spots in size.
Low TemperatureAgeing Behaviour of Zirconia Hip Joint Heads: J. Chevalieret al.
137
^ CI
o
'
o o
s
1
10
100
1000
10000
time (hours) Figure (1): Kinetics of phase transformation between 60 C and 134 C The plot of ln(b) as a function of 1/RT gives a straight hne of slope Q which was found to be about 109 KJ/mol. for this zirconia. The activation energy allows the simulation of ageing at 37 C as presented on Figure (2) which takes into account the scattering from one batch to another. Slow surface transformation occurs after an incubation period of 10 years. Approximately 25 years are then required to obtain a monoclinic content of 20%. I. Thompson et al. calculated a monoclinic content of 20% after only 5 to 11 month at 37 C in another zirconia ceramic. This stresses the need to produce a well controlled microstructural/chemical composition like in the presently studied zirconia ceramic to obtain a high degree of reliability.
80 : 70 60 o 50 40 - h 30 o 20 o o 10 0-
^
^
v i ^ ^ ^ ^ ^n > ^ ^ ^ ^ ^ J^^^^^
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^^J^^^p^ 1
\ ^ \
10
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15 20
\
\
25
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\
\
35 40
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’
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time (years) Figure (2): Predicted ageing behavior at 37 C for PROZYRfi heads.
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Wear of UHMWPE is one of the main issues in hip arthroplasty. The question that can be risen is the effect of surface transformation on wear behavior of zirconia femoral heads against polyethylene cups. To answer the question a first series of tests was conducted on heads artificially aged or not against commercial PE cups. Results are summarized in table I for 4 different monoclinic contents, in terms of PE weight loss after 5 millions cycles and monoclinic content after testing. Table I shows no significant effect of transformation on PE weight loss, even for 40% monoclinic content. On the other hand, wear has no effect on surface transformation because monoclinic content is the same before and after wear tests. Wear rates measured in this study (0.6 to 1.10 mg / million cycle) are low when compared to the limited number of publications dealing with zirconia - UHMWPE pairing [1,9]. Indeed, McKellop [1] and Saikko [9] found a wear rate of 10 mg / million cycle in bovine serum and 0.4 to 7.1 mg / million cycle in Ringer’s solution respectively. This stresses the effect of lubricant conditions (in particular protein content) and of the quality of initial surface finish (2 or 3 nm for these zirconia heads) on wear. This preliminary study should be continued up to 10 millions cycles and completed by a microstructural analysis of both heads and cups. Also the influence of the media on wear results must be understood in order to propose relevant lubricant conditions. 1
1
monoclinic content (%) before wear experiment
monoclinic content after 5 millions cycles
< 2% < 2% 30 %
< 2% < 2% 30% 40%
40%
Polyethylene weight loss (+ I mg)
5mg 3mg 4mg 3mg
1 1 1 1
Table I: Wear results after 5 millions cycles. CONCLUSIO N Low Temperature Transformation of commercially available zirconia hip joint heads is investigated in a large range of temperatures. The activation energy is calculated with a good accuracy, leading to a prediction of ageing at 37 C. A slow surface transformation can be recorded after an incubation period of 10 years in body situation. A first series of wear tests on a hip joint simulator suggests surface tetragonal to monoclinic transformation doe not increase wear significantly up to a level of 30 to 40 % monoclinic content which corresponds to 25 to 35 years at 37 C. This kinetic study cannot be extended to other zirconia ceramics and specific characterization of one given zirconia is highly recommended. REFERENCES Mc Kellop H., Lu B., Benya P., Park S.H., OrthopaedicResearchSociety,1992, paper n 402. Gupta T.K., Lange F.F., Bechtold J.H, J. materialsScience,1978, 13, 1464-1470. Kobayashi K., Kuwajima H., Masaki T., SolidStateIonics,1981, 3/4, 489. Sato T., Shimada M, J. Am. Ceram. Soc, 1985, 68, 356. Thompson I., Rawlings R.D., Biomaterial,1990, 11, 505-508. Lilley E., ceramictransactions,1990, 10, 387-406. Christian J.W., The theory of transformations in metals and alloys, Pergamon Press, 471 p. Tsubakino H., Sonoda K., Nozato R., J. MaterialsScienceLetters,1993, 12, 196.
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CHARACTERIZATIO N OF ZIRCONI A COATE D BY BIOACTIV E GLASS : PRELIMINAR Y OBSERVATION S M. Bosetti^, M. Santin^, M. Mazzocchi*, A. Krajewski*, M. Rastellino’^, A. Ravaglioli* and M. Cannas’^ ’^Department of Medical Sciences, Human Anatomy, University of Torino, 28100 Novara, Italy Institute for Technological Research on Ceramics, (IRTEC), CNR Faenza, Italy ABSTRACT The cell/biomaterial interface is one of thefimdamentalaspects that should be studied to evaluate the biocompatibility of a new implantable material. Cell attachment and orientation in vitrodepends on mechanical and physico-chemical characteristic of the biomaterial, influenced particularly by protein adsorption to the material itself This abstract reports observations on serum protein adsorption, growth and adhesion of fibroblasts on Zirconia covered surfaces by two different phosphate vitreous substances (differing for Tantalium and Lantanium presence). The pattern of adsorption of serum proteins onto Zirconia surfaces was characterized by an increased presence of proteins in the molecular weight range of 20-90 kDa thus permitting to conferm that bioactive glass coating led to a drastic increase of adsorbed proteins. Cell cultures has defined an accettable biocompatibility with a significantly greater cell growth on AP40 and RKKP coated materials than on uncoated Zirconia. In conclusion, these coatings permit the combination of the mechanical qualities of Zirconia (elsewhere demonstrated), which is not very biocompatible, with the properties of new ceramics which are not mechanically resistant, but biocompatible. KEYWORD S Zirconia; in-vitro biocompatibility; glass coating INTRODUCTIO N Biomaterials for bone implants can be subdivided into four categories according to their relative surface reactivity with surrounding tissue: (1) toxic when tissue dies, (2) non› toxic, biologically inactive when tissue forms a non-adherent fibrous capsule around the implant, (3) non-toxic, bioactive when tissue forms an interfacial bond with the implant, (4) non-toxic, dissolution of the implant when tissue replaces implant. To combine the mechanical properties of a high-strength inert ceramic with the specific properties of bioactive materials a composite was suggested based on Zirconia [1] as a support and biological glasses as coating. An optimized thermal treatment process was carefully developed to coat Zirconia substrates by amorphous or glass-ceramic AP40 and RKKP. The preparation of the coatings consisted in covering the Zirconia substrate by a 139
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slip (binders and water), then heating them at temperatures slightly above the melting, obtaining 100-200 mm thick layers [2, 3]. We studied two coatings, differing for Tantalium and Lantanium presence, and evaluated their in vitro protein adsorption and cytocompatibility, taking Zirconia as reference material. MATERIAL S AND METHOD S Zirconia substrates (d « 6.00–0.05 g cm"^ , corresponding to a relative density of 98.5%– 0.05 in respect to the theoretical one) sintered at IRTEC-CNR Faenza, Italy, and obtained by 3% Y203-stabilized Zirconium oxide powders, with an average grain size of about 1 mm diameter, was used As coating materials two bioactive glasses were used: AP 40: 44.30 Si02 / 24.50 h-C^^^^^h’ ^^-^^ CaO / 4.60 Na20 / 0.19 K2O / 2.82 MgO/4.99CaF2. RKKP: 43.82 Si02 / 24.23 b-Ca3(P04)2 /18.40 CaO / 4.55 Na20 / 0.19 K2O / 2.79 MgO / 4.94 CaF2 / 0.99 Ta205 / 0.09 La203. We underline that the few % of Ta205 and La203 was added to enhance the osteoproductivity. The bioactive glass-ceramic were prepared by melting of the starting products in a platinum crucible at 1450 C in a laboratory Kiln for 2 hours. The melted glasses were normally quenched into cold water or poured on preheated graphite moulds to obtain bars. The coarse granulated glasses were powdered by an Alumina vibratory ball mill and sieved up to a granulometry less than 100 mm; they were utilized at average granulometries of 40 mm (RKKP) and 100 mm (AP40). Each coating was characterized by optical and scanning electron microscopy, compositional analysis and characterized by X ray dififractometry in order to stucfy the microstructure. The adhesion of the coatings on Zirconia was tested by Vickers Indentation Method at the coating/Zirconia interface [4]. The bioactivity of glasses and glass-ceramics, both as coating and as bulk, was evaluated and compared by examining serum protein adsorption and growth and adhesion of fibroblasts on the materials. Squares of biomaterials were incubated with serum for Ihoiu* at room temperature. After removal of serum, squares were washed four times in phosphate buffer saline. The squares were then cut in pieces, trasfered into Eppendorf tubes and incubated at 90 C for 10 min in a 2% (w/v) SDS. Samples were analysed by SDS-PAGE according to the method of Laemmli [5] and the gel stained with a Silver Stain kit (BioRad). For adhesion and proliferation tests, humanfibroblasts(MRC5) were added to the luminal surfece at 5x10^ cell/cm^. After 2 hours of cell adhesion the culture media has been added in each 1.5 cm diameter well in a 24 microplate well. Adhesion test was performed after 6h culture time while proliferation test after 5 days culture. The materials, were rinsed in PBS,fixed20 min at 60 C and stained 5 minutes in a 0.025% Acridine Orange solution, a nucleic acid staining [6]. Cell mmiber on each material, was counted using a fluorescent microscope. Statistical analysis of the data was carried out using a Etell computer equipped with SPSS for Windows software. Duncan test was performed to compare, adhesion and proliferation
Characterizationof Zirconia Coated Bioactive Glass: M. Bosettiet al.
141
results for the four materials tested ; p value was obtained from the ANOV A table. The conventiona l 0.05 level was considere d to reflect statistica l significance . RESULT S AND DISCUSSIO N Studies of serum protein adsorption allowed to give clues about the formation of a conditioning layer on the material surfaces. The electrophoreti c patterns of the serum protein desorbed from the surfaces clearly showed an increase d protein binding after coating of the Zirconia with AP40 and with RKK P (Figure 1). No significant difierenc e was observed concernin g fibroblastattachmen t on to the three substrates studied: Zirconia, AP40 coated Zirconia and RKK P coated Zirconia (Figure 2). Fibroblast proliferatio n (Figure 2) showed a significantly greater cell growth on AP40 and RKK P coated Zirconia than on uncoate d control; no statistica l difference s were seen betwee n the two coated materials. The appearance of MRC 5 fibroblasts stained with acridine orange, after 6 days growth on bioactive glass coated Zirconia, evidence d a higher cell spread with a cell isles organization , characteristi c of new tissue formation. The results obtained in this paper seems to indicate that the coating of Zirconia either with AP40 or with RKK P may enhance the interaction s of the surfaces with biological component s thus favouring the integration of the prosthesis in the damaged tissue. In conclusion, these coatings permit the combination of the mechanica l qualities of Zirconia, which is not very biocompatible , with the propertie s of new ceramics which are not . mechanicall y resistant, but biocompatible
KDa
51.6
Figure 1: SDS-PAG E of proteins desorbed from material surfaces conditione d by serum. Lane 1: molecular weight standards; Lane 2: 1:25 diluted human serum; Lane 3: Zirconia; Lane 4: AP40; Lane 5: RKK P (for details see Materials and Methods).
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25
iSproliferation
MMM
ladheston
20
Zirconia Materials Zirconia AP40 RKKP
AP40 Proliferation 15,16*3,40 20,9318,01* 23,4117,50*
RKKP Adhesion 12,11*4,88 11,1215,64 10,00 s 4,88
*p<0.05 Figure 2: Human fibroblasts (MRC5) adhesion and proliferation results. Cell number referred to a 0.17 mm^ surface respectively after 6 hours and 5 days culture.
REFERENCES 1. Hulbert S.F., In: "An introductionto Bioceramics" Advances series in Ceramicsyolumc1, World Sc. Publ. Co. Ed. by L.L.Hench and J.Wilson, 1993, 25-29. 2. Berger G., Gildenhaar R., In: "Ten years of clinical experience"4thWorld BiomaterialsCongress,Apr., 24-28, 1992, Berlin, RFD, 33-35. 3. Ferraris M., Verne E., Moisescu C, Ravaglioli A., Krajewski A, In: Proceedingsof the 3rdMeetingandSeminaron Ceramics, Cells and Tissues,May, 2-4, 1996, Faenza Ed Italy. 4. Ferraris M., Verne E., Appennino P., Moisescu C, Krajewski A, Ravaglioli A, Proposed for prints in J. ^w. Cer. Soc. Bull. (1997). 5. Laemmli U. K., Nature1970, 227, 680. 6. Kirkpatrick C I , Dekker A, Biomat.TissueInterface1992,10, 31-41.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CALCIU M PHOSPHAT E PARTICLE S AR E FOUN D AT THE POLYETHYLEN E INSER T SURFAC E WHETHE R IMPLANTE D WIT H HA-COATE D DEVICE S OR NOT . A SEM-EPM A STUDY . Patrick Frayssinet*^, Laurent Gineste^, G. Bonel*, N. Rouquet* * Bioland, 132 Rte d’Espagne, 31100 Toulouse, France. See de Biologie Buccale, Ecole Dentaire, Universite Paul Sabatier, Toulouse ^ Laboratoire du Tissu Osseux et des Pathologies Osteo-articulaires, Universite Paul Sabatier, 31100 Toulouse, France. ABSTRAC T Calcium phosphate debris have been searched using SEM-EPMA at the surface of several polyethylene cups whether implanted with HA-coated devices or not. It appeared that calcium phosphate particles which had characteristics suggesting that they were not from artifact origin were found at the surface of the most of cups. Several hypothesis can explain the presence of these particles at the polyethylene surface. INTRODUCTIO N Polyethylene debris have proved to be responsible for dramatic osteolytic processes leading to aseptic prosthesis loosening. Third body wear, due to the presence of metal or ceramic particles at the polyethylene surface, has been suspected in many cases. Bloebaum et al. (1) reported the presence of calcium phosphate particles at the surface of polyethylene hip inserts which had been implanted with HA-coated metal backs or HA-coated hip stems. He suggested that HA-particles from the coating had migrated to the articulating surface causing the release of UHMWPE particles, hence a foreign body reaction. We examined the surface of several polyethylene hip implants with scanning electron microscopy (SEM) coupled to energy dispersive spectroscopy (EDS) microanalysis in order to detect any calcium phosphate particle surface contamination. The UHMWPE implants were implanted with either HA-coated or non-coated devices and for various periods of time. MATERIAL S AND METHOD S Twelve different inserts were obtained during revision or autopsy. None of these inserts were retrieved for an abnormal wear but for a reason concerning the stem. The inserts examined were not selected to appear in this series which explains its heterogeneity. The characteristics of the implants are given in table 1. Different methods were used to prepare the UHMWP surface for SEM observation in order to avoid contamination or calcium phosphate precipitation on the surface after retrieval. The surface was not handled by the surgeon during retrieval and was first gently washed in an isotonic solution of NaCl. Upon arrival in the laboratory, each insert was immersed in a 1.5% solution of sodium hypochlorite for 24 hours to remove all organic remnants on the polyethylene surface. It was then washed with demineralized water. Once processed, the inserts were then divided into three groups based on the implant depth. Each was treated differently. Group 1 : The surface of the implant was gold palladium coated for direct observation by SEM at 15kV; Group 2 : An epoxy polymer was molded in the cup and the polymer 143
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detache d from the cup surface once polymerizatio n had taken place to make a print of the cup surface. This was then gold-palladium coated and observed by SEM ; Group 3 : An epoxy polymer was cast in the cup as described . The cup was then cut in half and the epoxy polymer detache d from the surface. The cupfragmentswere then gold-palladium coated and observed by SEM . Abrasion wear was characterize d by the presence of scratches , adherenc e wear by the presence of wrenche d polymer sheet and fatigue wear by the presence of small cracks. Table 1: characteristic s of the samples examine d implant number "^’ 1
granuloma
HA-coated device ""Ye r
femoral head
group 2
implantation time 0.5 year
group 3
1 year
Stainless steel
0
Yes
group 3
2 years
Stainless steel
0
Yes
group 3
3 years
Stainless steel
0
Yes
yes
group 1
2 years
0
No
yes
group 1
3 years
0
No
no
group 1
11 years
0
No
unknown unknown
no no
group 1 group 1
2 months 5 years
yes yes
No No
10
unknown
no
group 1
15 years
yes
No
11 12
unknown unknown
no no
group 1 group 1
12 years 6 years
Stainless steel Stainless steel Cobaltchromium alumina stainless steel Cobaltchromium Alumina stainless steel
yes yes
No no
metal back
7
type of acetabula r wmponent biarticulate d (Zimmer, Swindon, UK ) biarticulate d (Zmimer, Swindon, UK ) biarticulate d (Zimmer, Swindon, UK ) biarticulate d (Zimmer, Swindon, UK ) Harris (Zimmer, France) Harris (Zimmer, France) unknown
8 9
2 3 4 5 6
process
Stainless steel
’
Q
RESULT S Implant n l: Scratches were found at the surface of the insert base with much debris inside these scratches and their vicinity. ED S reveale d that these particles were made of iron-chromium-nickel , titanium-aluminu m or calcium phosphates . The debris were less than SOfim in size. Implant n^2: debris were rare at the implant surface. Calcium phosphate particles were found on one polymer fragment only. However, there were numerous particles of stainless steel (ironchromium-nickel ) and titanium alloy. Scratches were found on the base of the insert. There was a pollution by silicon products. Implant n 3: iron-chromium-nicke l particles were found at the surface of every polyethylen e part. Some titanium alloy particles were also found. Pollution caused by silicium products disseminate d at the polymer surface of one part was observed . Small cracks were observed at the polymer surface. Implant n 4: No scratch was found on the implant surface, however , some poljmier sheets detache d from the material were found. A cluster of calcium phosphate s and iron-chromium-nicke l debris was apparent.
Calcium PhosphateParticles are Found at the PolyethyleneInserts Surface: P. Frayssinetet al.
145
Implants VL’S and 6: No scratch was found on the implant surface. Sodium chloride and calcium phosphates were found on the implant surface. Implant n 7: A lot of particles consisting of Ca, P, O, Na, CI, Si and Mg were found at the implant surface. These particles were of fluffy appearance and less than 60 ^im in size. Implant n 8: A few particles composed of Ca, P and O, of porous aspect were found at the polymer surface located close to the edge of the acetabular cup where scratches were visible. Other much more numerous particles consisted of aluminum oxide compound containing traces of titanium. These particles were smaller than 20 \xm. Implant n 9: no particles were found in the observed region. Implant n 10: A lot of particles composed of Ca, P and O were found encrusted in the polymer surface. Their surface was flattened and they were fragmented. Implant n ll: small fragments of a compound made of Ca, P and O formed encrusted clusters of particles in the polyethylene surface. Other particles consisted of Ca and O. Na and CI were sometimes found associated with the Ca-P-0 particles. Implant n 12: Particles composed of Ca, P and O were found encrusted within the polyethylene surface, particularly m the scratched region. Calcium oxide and Fe particles were also found. All such particles had a flattened surface and were fragmented (fig. 1). Figure 1 : SEM photograph (a) and EDS (b) of a flattened particle encrusted in the surface of the implant n 12 showing that the particle was made of Ca, P and O.
DISCUSSIO N AN D CONCLUSIO N This study showed that the presence of many particles having different origins can be found at the UHMWP inserts surface. We have found calcium phosphate particles at the surface of every kind of polyethylene inserts whether they,were implanted with HA-coatings or not. We were not able to show if these particles were constituted of one or several calcium phosphate phases and what is the
146 Bioceramies Volume 10
Ca/P ratio of these materials. The reason for this is that the use of a standardles s EPM A (electro n probe microanalysis ) technique without some independen t confirmatio n of the compositio n in test specimen s can lead to very large errors, which may go unrecognize d (2). Several origins may be attributed to the calcium phosphate particles found at the polyethylen e surface : a particle contaminatio n before the implantation , particularly during the manufacturin g processes , or after the implantation at the retrieval or during the specime n preparation. a debris migrationfrombone tissue. a secondary nucleatio n at a foreign body surface in a supersature d liquid in calcium and phosphorus. d by the ED S analysis, and the absence of collagen The absence of carbon in these particles assesse traces (3) in material make the bone debris origin very unprobable. The presence of NaCl crystals or silicated compounds at the surface of some implants in our series shows that the control of surface contaminatio n is a critical step in the qualitative electro n microscopy analysis. NaCl crystals are undoubtedl y formed during the surface drying step and some Si compounds may be brought in on the surgeon gloves or during sectionin g of an implant even with a protecte d surface. However, the presence of calcium phosphate particles encruste d in the polymer surface is in favor of the presence of these particles before the retrieval. Calcification of syntheti c or natural polymers once implanted in human or animal bodies is very well documented . Vascular devices , breast implants, collagen based materials very often show calcificatio n that may be visible on radiographs. Serum and extracellula r fluids are solutions supersaturate d in calcium and phosphorus. Furthermore , Radder et al (4) demonstrate d that copolymers rich in polyethylen e oxide were associate d with rapid and abundant calcification . It is therefore not surprising that heterogeneou s nucleatio n can take place at the polyethylen e surface once it is in contact with such a medium. We can suppose that the polyethylen e surface would provide a good seed material for heterogeneou s crystallizatio n of calcium phosphate , specially in the zone where the polyethylen e surface presents molecular periodicit y or in the cracks and crevices . Furthermore , the fragmentation of calcium phosphate s at the surface of the cups, creates nuclei for secondary nucleatio n (5, 6). REFERENCE S 1. Bloebaum, R.D., Beeks, D., Dorr, L.D., Savory, C.G., DuPont, J.A., Hoffmann, A.A., Complications with hydroxyapatit e particulate separation in total hip arthroplasty. Clin Orthop, 1994, 298: 19-26. 2. Goldstein, J.I., Newbury, D.E., Echlm, P., Joy, D.C., Romig, Jr, A.D., Lyman, C.E., Fiori, C , Lifshin, E., Quantitative X-ray analysis: theory and practice. In: Scanning Electron Microscopy and X-Ra y Microanalysis. A Text for Biologists, Materials Scientists , and Geologists. Second Edition. Plenum Press. New York and London 1992. 3. Reid, S.A., Micromorphological characterizatio n of normal human bone surfaces as afimctionof age. Scanning Microscopy 1, 1987: 579-59 7 4. Radder, A.M., Davies, J.E., Sodhi, R.N.S., van der Meer, S.A.T., Wolke, J.G.C., van Blitterswijk, C.A., Post-operativ e carbonate-apatit e formation in PEO/PBT copolymers (Polyactivefi) . Cells and Materials 1995, 5: 55-62. 5. Mersmann, A., Kind, M., Chemical engineerin g aspects of precipitatio n from solution. Chem Eng Technol 1988, 11:264-276 . 6. Mullin, J.W., Heterogeneou s nucleation . Crystallization, 3. ed, Butterworth, London 1985: pp 182-185 ,
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BON E REMODELLIN G AROUN D LOAD-BEARIN G CONDITION S
IMPLANTE D
MATERIAL S
UNDE R
M. Oka, Y.S.Chang, S. Yura, K. Ushio *, J. Toguchida, T. Nakamura * Department of Artificial skeletal systems, Research Center for Biomedical Engineering, * Dept. of Orthop. Surg., Kyoto University, 606 Kyoto, Sakyoku, Shogoin ABSTRAC T Bone formation and remodeling around implanted materials are influenced by the load-bearing conditions. In this study, three types of material were implanted into dog femoral condyles and bone formation and remodelling were observed for 24 weeks thereafter. Even thickening of lamellar bone was observed around bead-coated alumina implants, whereas thick fibrous tissue surrounded by corticalized bone formed around those made of smooth alumina. With an implant made of an artificial osteochondral composite material, thickening of ingrown trabeculae could be observed as early as 4 weeks. Bone ingrowth into the titanium fiber mesh was abundant and increased with time after implantation. This interstitial bone ingrowth resulted in the complete integration of this implant and the viable host bone. These results suggest that artificial osteochondral material is a promising material for joint replacements. KEYWORD S Bone remodelling. Artificial articular cartilage. Load-bearing condition. Bone ingrowth. Porous structure. INTRODUCTIO N Bone formation and remodeling around implanted material is influenced by the type of material, its surface properties and the anatomical site of implantation [1]. Bone formation and remodeling also depend on the loading conditions. As mechanical stress influences the remodelling process of bone and subsequently its structure and strength, in accordance with Wolffs law, materials implanted into bone should be under load-bearing conditions to elucidate interstitial bone ingrowth into the implants [2]. In this study, we compared bone formation and remodelling around solid and porous materials under load-bearing conditions. Another purpose of this study was to investigate the rate and extent of interstitial bone ingrowth into the pores of the titanium fiber mesh of an artificial osteochondral composite device, which we have developed as an artificial articular cartilage [4, 5, 6]. MATERIAL S AND METHOD S Three types of rectangular parallelepiped test pieces were used. One type of alumina test piece was manufactured into an implant with a mirror-finished articular surface and bead-coated sides; the other type was of the same size and shape with smooth side surfaces. The third type of test piece was the same shape, but manufactured ft-om pure titanium fiber with a porosity of 70% and mean pore size of 170 M ni, and the articular surfaces were 147
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shaped using the mirror-finished polyvinyl alcohol (PVA) hydrogel we have developed as an artificial articular cartilage. The composite material was made by infiltrating PVA (molecular weight, 220,000) solution into the pores of the distal half of the titanium fiber implant and then binding it to the fiber by gelling the PVA. The shear strength between the PVA hydrogel and titanium fiber was about 2.2MPa. In the control dogs, osteo-chondral defects were left open in the femoral condyles. Thirty-two mature adult mongrel dogs, weighing approximately 15-20 kg, were divided randomly into four groups of eight, which bore the implants for 4, 8, 12 and 24 weeks. All the specimens retrieved from the animals were examined using radiography, and all segments of the femur containing implants were then excised and fixed. Undecalcified sections of femur including the implants were cut and then ground into thicknesses ranging from 100 to 150 // m. Giemsa surface staining and contact microradiographic examination of the same cut section were performed. Histological changes of tibial joint surfaces, including menisci, were examined using Thionin staining. RESULT S Four weeks after the operation, thin and irregular ingrowth of newly formed immature bone into the pores of the bead-coated alumina implants was found, some of which reached the deepest layer. From 12-24 weeks after implantation, trabecular bone remodelling advanced further, producing a regular lamellar pattern. Contact microradiography (CMR) showed good fixation and uniform trabecular thickening around the bead-coated implants. (Fig. 1-A) Examination of the sections containing smooth-sided alumina showed that immature bone had also formed and surrounded the implants 4 and 8 weeks after implantation. However, 12 weeks postoperatively, a radiolucent zone surrounded by radiopaque changes around the smooth surfaced alumina at the bone-implant interface was apparent. After 24 weeks, the thickness of the fibrous layer had increased and the radiolucent zone surrounded by sclerotic bone observed on the radiographs was even more obvious. (Fig. 1-B) Abundant new bone ingrowth into the pores of the titanium fiber implant was present 4 and 8 weeks postoperatively (Fig. 2-A,B), and thereafter lamellar bone remodelling advanced further. By 24 weeks, neither demarcates nor radiolucent zones around this material were observed and, like the bead-coated alumina implant, no interposing fibrous tissue surrounding the titanium fiber mesh was present. Mature bone formed in the pores of the titanium fiber, which enabled this composite material to become attached firmly to the adjacent bone. DISCUSSIO N In order to determine the actual bone ingrowth and remodelling around the materials under conditions closer to the clinical situation, we believe the materials should be implanted under load-bearing conditions [1, 2]. Takagi et al [3]. implanted bead-coated and uncoated alumina test pieces into the load-bearing portions of the medial and lateral femoral condyles of dogs. They observed that the uncoated implants showed signs of loosening similar to those observed clinically. In contrast, bonding of the bead-coated implants was significantly stronger than that of the uncoated implants and increased with time after implantation. We used Takagi’s method and implanted three types of test piece. The results of our experiments revealed the differences between the bead-coated and smooth-surfaced alumina
Bone RemodelingAround ImplantedMaterials Under Load-Bearing Conditions:M. Oka et al.
149
implants particularly clearly. At the bone-implant interface, uniform lamellar bone thickening was observed around bead-coated alumina, whereas thick fibrous tissue surrounded by corticalized bone formed around smooth alumina. These prominent bone remodelling pattern differences might not be manifest unless the materials are implanted under load-bearing conditions. During the early postoperative stages micromotion of the implant will occur more often with the smooth-surfaced than bead-coated alumina, which may produce favorable conditions for fibrous tissue production. In the case of bead-coated alumina, however, micromotion will be prevented by the new bone formation around the beads [1]. Load-bearing conditions can be considered to be a more severe test for materials implanted in the bones than non-load-bearing conditions. Therefore, the considerable bone ingrowth into the pores of the titanium fiber mesh implant observed as early as 4 weeks after implantation, encourages us to use this material for joint replacement [2]. The reasons why we developed an artificial articular cartilage have been described in detail in our previous publications [4, 5, 6]. In short, it is necessary to preserve, as far as possible, the cancellous bones, which play an important role in the shock-absorbing effects of the joints. With respect to some of the mechanical properties of this material that are desirable for an artificial articular cartilage, such as lubricating and shock-absorbing functions, we obtained very encouraging results, which are described in detail in other papers [4, 5, 6]. The most important problem in developing an artificial articular cartilage is to fix the material quickly and firmly to the bone. With this aim, we manufactured an artificial osteo-chondral composite material composed of porous titanium fiber with PVA infiltrated into its pores. We implanted this composite material under the same loading conditions in this study as alumina implants. Excellent bone ingrowth into the pores of the titanium fiber mesh was observed 8 weeks after implantation and neither demarcated nor radiolucent zones around the implant were observed [5, 6]. As the wear-resistant properties of PVA hydrogel have been improved strikingly by gamma-ray irradiation, this composite material appears to be a very promising material for artificial joints. Furthermore, as the tibial joint surface remained intact against the artificial articular cartilage, in contrast to marked pathological changes with both types of alumina implant, PVA hydrogel is a very promising joint prosthetic material [6]. REFERENCE S 1. Chang, Y.S., Oka, M., et al., Biomat.,1996, 17, 1141-1148. 2. Oka, M., et al., Clin. Biomech.,ed. by Hirasawa, et al. Springer, Berlin, 1994, 124-137. 3. Takagi, H., Yamamuro, T., Hyakuna, K., Nakamura, T., Kotoura, Y. & Oka, M., Appl. Biomater.,1989, 23(A2), 161-181. 4. Oka, M., Noguchi, T., Kumar, P., Ikeuchi, K., Yamamuro, T., Hyon, S. H. & Ikada, Y., Clin. Mater.,1990, 6, 361-381. 5. Oka, M., Ikeuchi, K., Tsutsumi, S., Yamamuro, T., Noguchi, T. & Nakamura, T., Bomechanics in Orthopedics,ed. S. Niwa, S.M. Perren & T. Hattori, Springer, Berlin, 1992, 282-298. 6. Oka, M., Cahiers d’enseignement de la SOFCOT 57, Expansion scient, Francaise, 1996, 62-74
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Fig. 1-A: CMR of bead-coated alumina, 12 weeks after implantation. (X 10)
Fig. 2-A: New bone formation into the pores of titanium fiber, 8 weeks after operation. CMR X 10
Fig. 1-B: Bone remodeling around smooth sided alumina at 24 weeks after implantation. CMR X 10
Fig. 2-B: The artificial osteochondral composite material was integrated to the host bone, 8 weeks after operation. Giemsa surface stain X 10
CLINICAL USE OF CERAMICS
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CLINICAL COMPARATIVE STUDY BETWEEN POROUS COATED AND HYDROXYAPATITE POROUS COATED FEMORAL IMPLANTS - Retrospective matched pair study, average 5 years results Young Ho Kim,
Jeong Hwa Shon,
D Yong Choi
Department of Orthopedic Surgery, Hanyang University, Kuri Hospital, 249~1, Kyomoon-Dong, Kuri, Kyunnggi-Do, 471-020, Korea ABSTRACT Authors performed a retrospective matched-pair study for 64 uncemented total hip arthroplasty which include same number of hips with porous coated and HA porous coated femoral components using same kind of implant (Zimmer Anatomic^) to identify whether HA coat was efficient to get more favorable clinical results. Excellent function was achieved at 6 months in HA group and 1 year in porous group and mean Harris hip score maintained higher in HA group compare to porous group for whole period of follow up. Endosteal bone formation was noted mostly in zone 2, 6 in both groups and appeared earlier with larger extent and higher incidence in HA group compare to porous group. Radiolucencies were seen in zone 4 most commonly in both groups and the extent and the incidence of radiolucency were smaller and lower in HA group compare to porous group. Endosteal osteolysis were seen in 2 cases in HA group and 3 cases in porous group, and that in HA group was noted first earlier and smaller in mean size compare to that in porous group. Conclusively, HA coated implants improved and stabilized clinical and radiologic results earlier, and maintained superior results compare to porous coated implants until average 5 years follow up. INTRODUCTIO N Porous implant which was developed to use an alternative to cemented implant for young patients revealed the less than expected bone ingrowth and the more than expected fibrous ingrowth on retrieval specimens[2]. So thin HA coating was applied on the porous surface to solve the problem. But there is still debate regarding whether the use of a thin HA coating on a porous implant surface results in increase rate and amount of bone ingrowth emd such increase would result in improved clinical and radiographic performance. Dorr mentioned that there was no difference in the clinical results between HA porous coated and porous coated implants with same design[4]. But Whitecloud et al reported that early clinical and radiologic results with HA systems appeared superior to non HA coated systems with the support of animal study[6]. So we performed a retrospective matched pair clinical comparative study for 64 uncemented total hip arthroplasties which included same number of porous coated and HA porous coated femoral components to identify the efficacy of HA coating clinically. 153
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MATERIAL S AND METHOD S The uncemented anatomic hip system(Zimmer, Warsaw, U.S.A.) with porous pad of wire mesh were used as porous coated implant and the same design femoral prosthesis with additional CALCICOT(70% HA + 30% TCP) coating implant. Our study included 32 cases respectively in both groups. Harris-Galante II acetabular components with porous coating and modular Cr-Co femoral heads of 28mm diameter were used in all patients. All the variables such as age, sex and quality of bone etc. were similarly matched and average follow up period was 61.2(48-72)months in porous coated group and 59.2(48-68)months in HA coated group. All data were assessed and analized clinically and radiographically. RESULT S Excellent function was achieved at 6 months in HA group and 1 year in porous group and mean Harris hip score maintained higher in HA group compare to porous group for whole period of follow up. Mean time of weight bearing without support since operation was 6.4 weeks in HA group and 10.1 weeks in porous group, which was statistically significant. The incidence of thigh pain was 9.4 % in HA group, which was clinically lower than that in porous group, 18.7%. Endosteal bone formation appeared mostly in zone 2, 6 in both groups and extent and incidence of endosteal bone formation were lager and higher in HA group compare to porous group. Endostead bone formation was noted first at 3 months in HA group and at 6 months in porous group. In overall view, endosteal bone formation appeared earlier with larger extent and higher incidence in HA group compare to porous group. Radiolucencies were seen in zone 4 most commonly in both groups and extent and incidence of radiolucency were smaller and lower in HA group compare to porousp group. Radiolucencies were noted first at 6 months in HA group and at 3 months in porous group. In overall view, radiolucencies were seen later with smaller extent and lower incidence in HA group compare to porous group. Cortical bone remodelling reactions were commonly seen in zone 2, 6 in HA group and in zone 3, 5 in porous group and overall incidence of them were similar in both groups. Calcar remodelling was seen earlier with higher incidence in HA group compare to porous group. Incidence of pedestal formation including stable or unstable type was lower in HA group compare to porous group. Endosteal osteolysis around stem were seen in 2 cases in HA group, which was noted first at 32.7 month on averge and 3.5cm in mean size, and in 3 cases in porous group, which ws noted first at 25.5 month on average and 6cm in mean size. Mean linear wear rate was 0.30 – 0.13mm in HA group and 0.28 – 0.15mm in porous group, which were similar in both groups. Engh’s radiologic score for fixation and stability was increasing in fixation scale, and decreasing in stability scale with time in both groups, and maintain higher in HA group compare to porous group. Definite radiographic sign of bone ingrowth was achieved at 6 months in HA group, and 1 year in porous group.
Clinical ComparativeStudy BetweenPorous Coated and HA Porous Femoral Implant: Y.H. Kim et al.
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CAS E ILLUSTRATIO N This 51 year old female patient had total hip arthroplasty with HA porous coated anatomic stem for avascular necrosis of left femoral head. The patient could walk without difficulty in the absence of crutch since 4 weeks after operation. She had proximal femoral fracture with long spiral oblique fashion assoacted with the subsidence of stem after slip down at postoperative 6 week(Fig. 1). In retrieved stem, remained HA coating about porous surface could not be observed and more than 90% of porous surface was replaced by bone and porous pore were filled with bone mostly(Fig.2). DISCUSSIO N AN D CONCLUSION S This study shows that HA coated implants improved and stabilized clinical and radiologic results earlier compare to porous coated implants. HA group improved Harris hip score earlier and had lower incidence of thigh pain and could walk without support earlier than porous group. When we assessed the patients clinically, we felt that patient who had HA coated implants usually answer more comportable than those who had porous coated implants. These results might be due to achievement of the early initial stability of HA coated implant resulting from its early osteoconductive property. The radiographic results such as earlier appearance with larger extent and higher incidence of endosteal bone formations and later appearance with smaller extent and lower incidence of radiolucencies in HA group might also support the clinical results.
Fig. 1 Six weeks postoperative roentgenogram reveals long spiral oblique fracture of the femur from proximalportion to distal to the tip of stem.
Fig. 2. Microscopic finding of retrieved porous pad area of implant. Notes bony femoral ingrowth between fiber-mesh of titanium.
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Bone growth on to HA or porous coated stems showed endosteal bone condensation over the HA or porous coating on femoral stem followed by distal radiolucency around non-coated stem area. Distal radiolucency are probably due to micromotion between the stiff distal stem and the more elastic bone[3]. Higher incidence of distal radioluency in porous coated stems could be explained by less stability of porous coated stems resulting from less endosteal bone formation in addition to the difference of elasticity between bone and stems. Cortical remodelling probably reflected adaptive change of cortex to local stress pattern, which was regarded as evidence of good implant fixation[5]. Symmetrical cortical remodellings which was associated with extensive endosteal bone formation were appeared frequently in HA group, and asymmetrical type which was associated with increased localized intramedullary stress were appeared frequently in porous group. Mean linear wear rate of polyethylene was similar in HA and porous coated stem. This results were supported by the report in which HA coating does not increase polyethylene wear[l]. Endosteal osteolysis was appeared later and smaller size in HA group despite of similarity of mean linear wear rate in both groups. This results might be due to profuse impacted endosteal bone formation between proximal portion of femoral component and proximal portion of femoral canal functioned as banier to movement of polyethylene wear debri. Conclusively, HA coated implants improved and stabilized clinical and radiologic results earlier, and maintained superior results compare to porous coated implants until average 5 years follow up. REFERENC E 1. Bauer T.W., Taylor S.K., Jiang M. and Medendrop S.V., Clin Orthop, 1994, 298, 11. 2. Bobyn J.D., Engh C.A. and Classman A.H., Clin Orthop,1987, 224 : 303. 3. D’Antonio J.A., Apello W.N., Crothers CD., Jaffe W.L. and Manley M.T., / Bone Joint Surg, 1992, 74-A, 995. 4. Dorr L.D., / Bone Joint Surg, 1993, 75-A, 1728. 5. Geesink R.G.T. and Hoefnagels N.H.M., / Bone Joint Surg, 1995, 77-B, 534. 6. Whitecloud T.S., Cook S.D., Enis J.E., Amstrong D. and Lisecki E.J., / Bone Joint Surg, 1992, 74-B supp. HI , 252.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
REVISIO N RATE S AND RADIOGRAPHI C CHANGE S ASSOCIATE D WIT H DIFFEREN T SOCKE T INTERFAC E TECHNOLOGIES : CLINICA L RESULT S FRO M 456 PATIENT S AT FIV E TO EIGH T YEAR S FOLLOW-UP . M.T.Manley**, A.Edidin**, J.A.Epinette+, R.G.Geesinkt, J.A.D’Antonio*, W.N.Capello-. ** Osteonics Corp., Allendale, NJ, USA; +CRDA, Bruay la Buissiere, France; tState University of Limburg, Maastricht, Netherlands; * M, H & D Orthopaedics, Swickley, PA, USA; -Indiana University, Indianapolis, IN, USA. ABSTRAC T Over the last seven years, we have followed the clinical course of 456 total hip replacement patients all of whom received the same hydroxylapatite coated cementless femoral stem and either a control porous coated acetabular component, a press-fit hydroxylapatite coated acetabular component or a threaded, hydroxylapatite coated acetabular component. We found acetabular bone remodeling changes and implant revision rates were different for the different cup types; the hydroxylapatite coated threaded components needmg the smallest number of cup revisions for aseptic loosening (0.44%) while the press-fit hydroxylapatite coated components needing most revisions (10.2%). These rates compared with 1.0% for the porous controls. Radiographic data suggests that interface failure with the press-fit cups tended to initiate at the implant/bone interface inferior to the component where the shear and tensile stresses caused by physiologic loads are greater than elsewhere around the implant periphery. We concluded that the bond between bone and smooth hydroxylapatite coated acetabular component interfaces may be compromised by physiologic tensile stresses unless interlock between unplant and bone is achieved by threads or other features incorporated in the implant design. KEYWORD S Hydroxylapatite coatings, implant fixation, clinical resuhs. INTRODUCTIO N Clinical studies of uncemented cups at five years or more follow-up have shown failure rates from zero to about eight percent for porous implants [1-4], but as high as thirty to forty percent for other cementless designs [5-7]. Attempts to improve fixation by adding threads to the implant design have shown mixed results with early clinical failures reported [7-10] as well as good results with threaded cup designs which incorporated also a biological fixation surface in the design of the metal shell [7,11,12]. Our study was designed to compare the clinical performance at a minimum of five years follow up of acetabular components with two different hydroxylapatite interfaces and to compare the results of both to a control implant with a porous fixation coating. The intent was to determine whether a bonded or an interlocked interface produces the best prospects for long term stable fixation in the cementless acetabulum. 157
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MATERIAL S AND METHOD S In this clinical study conducted in the U.S. and Europe, we have followed 456 patients with 497 total hips to a minimum follow-up of five years. All patients received the same type cf femoral component (Omnifit-HA, Osteonics Corp., Allendale, NJ) and one of three acetabular cups each with a different fixation surface (porous, hydroxylapatite coated press-fit and hydroxylapatite coated threaded). Ninety-six patients received 103 cups in the porous cup group, 152 patients had 166 cups implanted in the hydroxylapatite press-fit cup group and 208 patients received 228 cups in the hydroxylapatite threaded cup group. After surgery, clinical and radiographic evaluations of each patient were conducted at five to seven weeks post-operation and at six months, at one year, and at each year thereafter. Analysis of radiographs included evaluation for radiolucencies, cancellous condensation, component migration, and osteolysis. Selected radiographic series from patients with both stable and unstable components were digitized and computer enhanced using commercially available image enhancement software (Imagika, Clinical Measurement Corp., Upper Saddle River, NJ).
RESULT S At the 5 year evaluation, 395 patients (86.6 per cent of the population) reported no pain associated with their joint or joints. Clinical results broken down by cup type showed that at 5 year follow-up the porous press-fit cup group had a mean Harris score of 94.1, the hydroxylapatite press-fit cup group had a mean score of 96.0 and the hydroxylapatite threaded cup group had a mean score of 97.6. These five year scores were not statistically significantly different one from another. Radiographic observations at a minimum of five years follow-up showed that a stable bone interface was maintained by 92.7% of the porous implants, by 88.2% of the hydroxylapatite coated press-fit implant group and by 100 % of the hydroxylapatite coated threaded implants (Figure 1). Analysis of revision rates showed one implant revised for aseptic loosening in the porous cup group (1%) and one m the hydroxylapatite threaded group (0.44%). By comparison, in the hydroxylapatite coated press-fit implant group, seventeen of 166 unplants were revised (10.2%) and a further four of these implants were found to be loose by radiographic criteria at five years follow-up. Analysis of the radiographs of these press-fit implants prior to failure showed a region of radiolucency inferior to the implant which increased in width as follow-up progressed (Figure 2). DISCUSSIO N Comparison of the failure rates between the cups in our study shows that the hydroxylapatite coated press-fit components performed significantly worse than the other two implant types (p<0.001). In assessing the difference between implants, the hydroxylapatite coatings were manufactured to the same specifications, the geometry of all cup types was similar and all implants were fabricated from commercially pure titanium. The difference between the successful groups and the unsuccessful group is in the morphology of the fixation surfaces. Both the hydroxylapatite threaded surface and the porous coating could achieve interlock with bone, while the hydroxylapatite coated press-fit surface relied for fixation on a bond between bone and implant coating.
Revision Rates and Radiographic Changes Associated With Different Cup Types: M.T. Manley et al.
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Finite element analyses [13,14] have suggested that if a metal backed implant is bonded to the acetabular preparation, very high local tensile stresses, caused by physiologic activity, will be found at the fixation interface close to the component rim. A potential for failure cf bonded interfaces is suggested by these models, and once bond failure occurs the models suggest that component loosening, distraction between inferior cup and bone and finally cup migration will result. Radiographic evidence of this failure of this type is shown in our study, with the hydroxylapatite press-fit interfaces tending to show initial signs of failure at the inferior interface. The absence of such failures with the hydroxylapatite coated threaded and the porous designs suggests that these interfaces are able to withstand the stresses imposed by patient activity, presumably because of the interlock achieved between the implants and the acetabular bone. SUMMAR Y This clinical comparison of three acetabular components of similar geometry but different bone fixation interfaces has shown acetabular revision rates that are dependant on the fixation surface used. The hydroxylapatite coated press-fit cups had an unacceptable loosening rate. Failure of the fixation interface with these implants began inferiorly to the implant and then progressed toward the pole. Similar interface failures with the hydroxylapatite coated threaded cups and the porous cups appear to have been prevented by the mechanically mterlocked interfece which these implants achieved with bone by the design of their fixation surfaces. Our study suggests strongly that interlock between an acetabular component and bone is a prerequisite for successful long term cementless fixation in the acetabulum.
REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14.
Heekin R.D., Callaghan J.J.,Hopkinson W.J., Savor>’ C.H., and Xenos J.S. J. Bone and JointSurg.1993, 75-A:,77-91. Kim Y-H., Kim V.E.M. J. Bone and JointSurg.1993, 75-B, 6-13. Schmalzried, T.P., and Harris, W.H. J Bone and JointSurg., 1992 74-A, 1130-1139. Schmalzried, T.P., Wessinger, S., Harris, W.H. J Arthroplasty.1994, 9, 235-242. Engh, C.A., Glassman, A.H., Griffin, W.L., and Mayer J.G. Clin Orthop.1988, 235, 91-110. More R.C., Amstutz H.C., Kabo, J.M.,Dorey, F.J.,and Moreland J.R, Clin. Orthop., 1992 282, 114-122. Pupparo, F.; and Engh, C.A, Clin. Orthop.,1991 271, 201-206. Apel, D.M., Smith, D., Schwartz, CM., Paprosky, W.G., Clin. Orthop.,1989, 241, 183-189. Capello,W.N. Clin.Orthop.,1990, 261, 102-106. Fox, G.M., McBeath, A.A., and Heiner, J.P., J Bone and JointSurg.,1994, 76-A: 195-201. Cartillier, J.C, In: HydroxyapatiteCoatedHip and Knee Arthroplasty,Expansion Scientifique Francaise, Paris, 1994, 165-168. Duthoit, E., Epinette, J.A., and Carlier, Y., In: HydroxyapatiteCoatedHip and Knee Arthroplasty,Expansion Scientifique Francaise, Paris, 1994, 176-186. Mann K.E., Osteonics Corp. Internal Report, 1996. Rapperport, D.J., Carter, D.R., Schurman, D.J., J. Orthop. Res.,1987, 5, 548-561.
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Figure 1 Hydroxylapatite Coated Threaded Cup at 7 Years Follow-up.
Figure 2 Enhanced Image of Hydroxylapatite Coated Press-fitCup at Four Years Follow-up.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
COMPARATIV E STUDY OF THE RESULT S BETWEE N CUSTO M NON-COATE D CEMENTLES S fflP IMPLANT S AND MIRRORE D CEMENTLES S HA - COATE D HI P IMPLANT S ON THE CONTRA-LATERA L SIDE . Dr M Mulier, G. Deloge University Hospitals U.Z. Leuven Departement of Orthopeadic Surgery Weligerveld,! 3212 Lubbeek Pellenberg (Belgium) ABSTRAC T Introduction Hydroxyapatite coatings to prostheses are reported to give a very early and reliable stability and show a better osseointegration compared with identical non-coated implants. Materials SL Methods: Since 1990 we have implanted 43 HAP- coated femoral stems based on the computer mirrorimage of the contra-lateral side custom hip stem. The patients received this custom prosthesis earlier by surgery using the I.M.P. ( = Intra-operative Manufactured Prosthesis) system of Stewal Implants. This study was setup in order to determine the influence of HA coatings on the performance of the I.M.P. implant. For each patient a HA coated stem and a custom-made rasp was made, based on the geometry data of the custom-made contralateral side. Of the 43 implanted prosthesis we have 30 patients with a minimum follow-up of 3 years on both sides.(3 - 7 years). We evaluated both sides clinically and radiographically. 161
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Results; All the 43 rasps and respective stems could be introduced without much diflSculty. This proves the high grade of similarity of the femoral cavities of both femurs Radiographically: On the uncoated side bony pedestal formation below the distal tip was observed in 50 % of the cases. On the contra-lateral HA coated side this pedestal formation was observed in 1 case. This pedestal formation may be triggered by excess force transmissionfromthe prosthesis to the bone possibly by late subsidence, (see figure 1 and 2) No other significant differences were observed radiographically. The HA coated implants didn’t show any form of osteolysis caused by defragmentation of the HA coating. Clinically: Most of the patients have a more secure feeling on the HA coated side during monopodal stance; this appeared insignificant. To compare Harris scores between sides in the same patient was difficult. Conclusions; The HA coated implants show a much better radiographical result then the non -coated implants. It is a challenge to coat the implants during surgery. This way the I.M.P. system could become available to all the patients with a difficult and unpredictable femoral geometry (i.e.: revisions, C.D.H.).
Comparative Studyof Non-Coated Implants andHA-CoatedImplants: M. MulierandG. Deloge 163
CEMENTLESS IMP
PEDESTAL 43 %
Figure 1: Radiological evaluatio n ( Pedestal) : Cementies s IM P stem.
CEMENTLESS IMP WITH HA COATING
PEDESTAL 5%
Figure 2: Radiological evaluation (Pedestal) : Cementies s IM P stem with HA coating proximal.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
IMPROVEMEN T OF THR WIT H SPONGIOS A META L SURFAC E USIN G TH E WEA R COUPL E CERAMIC-ON-CERAMI C Gerd Quack, MD, PhD; Gerd Willmann, PhD.*; Hans-Gerd Pieper, MD, PhD; Hartmut Krahl, MD, PhD., Prof. Department of Orthopaedic Surgery of the Alfried Krupp Hospital, D-45117 Essen (Germany) * CeramTec Medizintechnik, D-73207 Plochingen (Germany) Keywords: all-ceramic wear couple, total hip endoprosthesis with spongiosa-type metal surface Abstract Based on positive 6-years results of the SMS cup system (*Spongiosa A/etal 5’urface), we improve this THR to an all-modular system with the wear couple ceramic-on-ceramic. As corresponding results are available for ceramic/polyethylene couples of the SMS system, such improvement of the survival curve of this system may reasonably be ascribed to the use of all-ceramic couples. It is thus possible to clearly assign any failures which may additional occur to the new components. The follow-up of the first 50 patients shows good short-term results. Introduction In the meantime, uncemented hip endoprosthetics can look back on highly successful results. Despite this, the fact that THR tend to be subject to loosening after a dwelling period of 10 years cannot be ignored. Such process of aseptic loosening is essentially caused by increased abrasion of the polyethylene (PE) used in material-mix couples in which metal or alumina ceramic balls are used together with PE-cups. The abraded PE particles which are not carried off by the organism, cause high formation of granuloma which again cause osteolysis alongside the prosthetic components - which in the end causes the loosening of the implant [1]. Alumina ceramics Replacing the metal femoral head by a Biolox femoral head allows for the rate of abrasion to be reduced by > 50%. Further reductions of such rate can only be achieved if the PE insert is replaced by a ceramic insert. The use of all-ceramic couples, i.e. THR femoral head and acetabular cup made of Biolox forte, allows for the rate of abrasion a reduction by factor 100 (tbl. 1) [6, 8]. The properties of Biolox (AI2O3) ceramics, which are described below, make the material particularly suited for hip endoprostheses: Aluminaceramicsare bioinert The high-purity alumina available today are extremely resistant to corrosion and will not emit any ions to the body. Hence, they can be classified as both, biocompatible and bioinert materials. This fact also means that they will not induce any connective osteogenesis. Wettability ofAluminaceramics The wettability of almnina in respect of polar liquids such as synovial fluid is better than the one offered by PE or CoCrMo. Hence, the material offers the best prerequisites for a lubricating film to establish between the prosthetic counterparts. Formation of the lubricating film is supported by the gap provided between the ceramic femoral head and the ceramic insert, which is generated by the difference in diameter between the head and the insert. The type of friction involved in vivo is a mixed friction which among other things is due to the oscillating movement of the hip joint. 165
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Configuration fern, head/cup
Abrasion / year in mm
Metal/PE diam.
heads of 22 mm up to 0.5 mm rest: > 0.2 nun Biolox/PE <0.1 mm Biolox/Biolox up to 0.005 nun Biol. forte/Biol. forte < 0.001 mm Basic prerequisite: no loosening process involved of the stem and /or the cup Tbl. 1: Rate of abrasion
Fig. 1: SM S cup, Biolox forte insert
Wear ofAluminaceramics Owing to its extreme hardness (pyramid hardness of 2000 HV), the material is subject to extremely low wear. The material’s median grain size of < 0.002 nun and its high density allow for highly polished functional surfaces, i.e. of surfaces featiuing a low peak-to-valley height to be achieved. In combination with the good wettability offered by AI2O3 ceramics, the material enables excellent tribological properties. According to the studies conducted by Saikko [6] and by Walter [8], the rate of abrasion obtained for Biolox forte/Biolox forte couples ranges below 0.001 nun/year. DimensionalstabilityofAluminaceramics Owing to their high elasticity modulus of 380 GPa, alumina ceramica are extremely rigid and are not subject to neither plastic nor elastic deformation. This means, that they offer long-term dimensional stability even when exposed to high strain. Aluminaceramicsare safe The material offers an extremely high compressive strength of > 500 MPa. This means that the femoral heads and cup inserts made of it will notfracturewhen exposed to the different types of stress incurred in vivo. The resistance tofractureoffered by Biolox forte femoral heads of 28 nun in diameter exceeds 50 kN (> 5 tons), while the one offered by the smallest Biolox forte insert is 86kN(>8tons)[l,9]. However, the current experience in the use of ceramic cups has shown that the bone will not integrate the ceramic surface in the sense of a connective osteogenesis (ace. to Osbom), owing to the ceramic material’s absolute bioinertness. As a result, the implant is subject to early loosening. This basic property of ceramics is the cause of the failures experienced for monobloc cups. The autopsy examinations performed by Plenck [4], Hensge [3] and Fritsch [2] represent the connecting link between the theoretical and the in-vitro properties of all-ceramic couples, and the clinical results obtained for them. Such results can be summarized as follows: - Abrasion rates of < 0.005 nun/year are obtained if neither the stem nor the cup have loosened. - The diameters of the abrasion particles of AI2O3 ceramics ranges between 0.001 and 0.002 mm which is definitely lower than the ones observed for PE particles which were up to 0.02 mm. - The amount of abrasion particles is extremely low (< 0.005 mm/year) as long as there is no loosening of the implant and the cup and the stem. - Abrasion particles of AI2O3 ceramics are classified as biocompatible material. They will not cause any undesirable tissue reactions, which is due to their bioinertness and the small size and small amount of abrasion particles.
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Osseointegratio n of the uncemente d SMS-TH R system The degree of osseous integration of uncemented metal cup components essentially depends on the design of the prosthetic surfaces. As far as the modular SMS cup system is concerned, the biological idea is to the fore: - Preservation of as much bone as possible. Resulting from this a spherical shape of tiie cup was devised. A cone is provided on tiie inside of the cup in order to allow for optimumfixationof the PE or of the ceramic insert. - Osseous integration of the prosthesis into the osseous bed through spongiosa-type wide-meshed enlargement of surfaces, which is achieved by coating the implant with tripodes: the cancellous bone represents the load-carrying and load-transmitting structure for positioning the cup and the stem. It reacts to the implant by integrating it. For this reason, the implant surfaces must feature a load-carrying structure and must provide suflBcient gaps into which the trabeculae of the spongiosa and the supplying vessels may vascularize [5]. Until December 1996, a total of 709 of these SMS cup systems were implanted. The follow-up examination of the first 95 uncemented hip endoprostheses (implanted in 1987 to 1991) with wide-meshed SMS cups yielded good medium-term results, with only one cup loosening and protruding into the minor pelvis (for data refer to table 2). The system used consisted of a ceramic femoral head (Biolox) and a PE cup insert. Yielded from this was a loosening rate of < 1.05% after a dwelling time of 6.03 years, and the good osseous integration predicted was confirmed. The consequences which can be drawnfromthefindingsobtained for the use of alumina ceramics [7] and from the positive medium-term results (*tained for the SMS cups must be a division of fiinctions [5]: - Ceramics should be used as gliding components offering minimum abrasion for the articular fimction as such. - Metal (CoCrMo, Ti) should be used for thefimctionof osseous integration. A cup system featuring a 3-dimensional wide-meshed surface and using Biolox forte ceramic inserts was developed, which as the central element offers a modular structure of the gliding couple. Resulting from this development was im imcemented and completely modular THRsystem offering optimum prerequisites to enable connective osteogenesis and low wear of the aJumina articular surfaces (fig. 1). On the basis of the experience gained from the use of the SMS cup system over the years, the improvement achieved in respect of the life of the prostheses is clearly due to the use of ceramic/ceramic couples. Also, it is possible to assign the causes of failures. Such findings shall be specified fiirther with the help of a comparative study investigating the use of Biolox^E and Biolox forte/Biolox forte couples. SMS-CU P / aver. 6.03 y. - 95 SMS cups -47 females (49,47%) - 48 males (50,53%) Mean Score: Harris hip score 92.80 pts (+/- 7.11) Merle d’Aubigne sc. 16.74 pts (+/-1.36) - Aseptic loosening:! SMS cup = 1.05% TbL 2: Medium-term results
From November 1995 to March 1997, tiie first 50 patients below the age of 61 have received a gliding couple consisting of ceramic Biolox forte femoral head and insert. There have been no complications due to the use of ceramics. So far, the regular follow-up examinations have yielded good clinical (Harris hip score) and radiological results. In respect to the surgical technique used, the items specified below should be accounted for: - Use of an exact surgical technique and optimum positioning of the cup must be ensured.
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- Both, the cone available on the inside of the cup, and the cone of the stem must never be damaged during handling and implantation . - Necessit y of using special implantation instruments . - To enable optimum cone-fitting , the ceramic insert must not be allowed to rest on the metal bottom of the cup (fig. 1). Summary The SM S THR-system (manufacture d by ESKA ) offers optimum osseointegratio n owing to the spongiosa-typ e design of its surface. The abrasion rate is drastically reduced as a result of the improved gliding propertie s of the all-ceramic (Biolox forte) couple, as long as the use of an exact surgical technique is ensured. The residual abrasion particles produced are bioinert and will be carried off by the organism without any problems. The ceramic component s will prolong the survival of the THR caused by the decrease of any osteolysi s due to abrasion. As a resuh the prosthesis will not loosen, or at least will loosen at a much later point in time. This means a much better prognosis, especiall y in the case of younger patients. A comparing study with the wear couples Biolox/PE and Biolox forte/Biolo x forte will help specifying these statements . First experience s with 50 cases, 2 years afler the first implantation , show good short-term results. Reference s 1. Clarke, I.C , P. Campbell, N. Kossovsky: Debris-mediate d osteolysi s - a cascade phenomeno n involving motion, wear, particulates , macrophage induction and bone lysis. In: St. John, K.R. (ed.): Particulate Debris from Medical Implants. AST M STP 1144, Philadelphia 1992, pp. 7-26. 2. Fritsch, E., H. Mittelmeier, J. Heisel, K. Remberger, S. Pahl: Micro- and macroscopic findings on capsular tissues of the hip after alumina arthroplasty. Proc. 6th Biomaterials Symposium "Ceramic Implant Materials in Orthopaedic Surgery", Sept 21-23, 1994, Gottingen (Germany); H.G. Buchhom, H.-G. Willert, in press 1996. 3. Hensge, E.J., I. Bos, G. Willmann: AI2O3 against AI2O3 combination in hip endoprostheses . Histologic investigation s with semiquantitativ e grading of revision and autopsy cases and abrasion measures. J. Materials Science Materials in Medicine 5 (1994) pp. 657-661 . 4. Plenck, jun. H., M . Buhler, A. Walter, K. Knahr, M. Salter: Fifteen years experienc e with alumina-cerami c total hip-joint endoprostheses : a clinical, historical and tribological analysis. In: Ravaglioli, A., A. Krjewski (eds): Bioceramic and the Human Body. Elsevier Appl. Sci., London, New York 1992, pp. 17-25. 5. (Juack, G., G. Willmann, H. Krahl, H. Grundei: Konzeptionell e Uberlegunge n zur Verbesserung der Pfanne der ESKA-Hftendoprothese durch die Gleitpaarung Keramik/Keramik [Conceptiona l consideration s relating to the improvemen t of the acetabula r cup of the ESK A hip endoprosthesi s achieve d through the use of ceramic/cerami c couples] . Biomed. Technik 41 (1996) 9, pp. 253-259 . 6. Saikko, V.: Wear test of the couple BIOLO X forte/BIOLOX forte. In: W. Puhl (ed.): Performance of the wear couple BIOLO X forte in hip arthroplasty. Enke Verlag Stuttgart, 1997. 7. Sedel, L., RS . Nizard, I. Kerbouli, J. Witvoet: Alumina-alumina hip replacemen t in patients yonger than 50 yars old. Clin. Orthop. 198 (1994), pp. 175-183 . 8. Walter, A.: Investigation of the wear couple BIOLO X forte/BIOLOX forte. In: W. Puhl (ed): Performance of the wear couple BIOLO X forte in hip arthroplasty. Enke Verlag Stuttgart, 1997. 9. Willmann, G.: Hiiftgelenkersat z - eine tribologische und konstruktive Herausforderun g [Hip arthroplasty - a challenge in respect of tribology and design]. Mat. Wiss. u. Werkstofitechni k 27 (1996) pp. 199-205 .
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
ACETABULA R ARTHROPLAST
RECONSTRUCTIO N IN REVISIO N TOTA L Y U S I N G A BON E G R A F T S U B S T I T U T E
HI P
R.P. Pitto and D. Hohmann
Department of Orthopaedics, Friedrich-Alexander University, Waldkrankenhaus St. Marien, RathsbergerstraBe 57, 91054 Erlangen, Germany ABSTRAC T Twenty acetabular reconstructions in revision total hip arthroplasty with severe loss of bone stock were performed combining the use of autogenous bone grafts, synthetical hydroxyapatite ceramic and reinforcement rings. The rings were fixed with screws on the host bone. The grafts fused within 4 months after the operation in all the cases. No migration of the acetabular component or lysis of the mixed graft was seen in 19 cases after 2 years. One implant failed because of malposition and was revised 6 months after the operation. These preliminary findings give rise to cautious optimism that this is a reliable method for acetabular reconstruction. KEYWORD S Hip Prosthesis, Revision, Bone Stock, Bone Graft Substitute, Ceramic. INTRODUCTIO N Deficiency of bone stock is a major problem in revision arthroplasty. Filling of the cavities by cement or metal leads to ftirther bone defects, if renewed loosening occurs. The use of autogenous bone grafts is a biological way to solve the problem, but the quantity of the available harvested material is limited. For ethical, bacteriological and viral safety reasons, management of bone banks is becoming increasingly restrictive [2]. Synthetical bone substitutes offer an alternative to homologous grafts. The goal of this study was to evaluate prospectively the clinical and radiological results of acetabular reconstructions after revision of loose acetabular components with severe bone stock defects combining the autogenous grafts with synthetical hydroxyapatite ceramic. MATERIAL S AN D METHOD S Twenty revision arthroplasties were performed using the impaction grafting technique on the acetabular side [4]. The morselized autogenous grafts were mixed with synthetical hydroxyapatite ceramic (granulate or blocks, Synthacerfi, Scientific Development, Munich, D) (Fig.l and 2). Reinforcement rings of Muller (9 cases), Ganz (5 cases) and Burch-Schneider(6 cases) (Protek, Munsingen, CH) were used to anchor the new Prothesis, to impact the mixed autoheterografts and to protect them during the healing. The clinical assessment was performed according to the criteria of the Chamley 6-6-6 Hip-Score-System [1] and a similar method was used to classify the pre- and post-operative bone stock of the acetabulum [3]: normal acetabulum (grade 6); peripheral ectasis (grade 5); protrusion (grade 4); ventral defect (grade 3); ventro-cranial defect (grade 2); dorsal defect or discontinuity of the pelvis (grade 1). 169
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170
1 1
* ’’ ^ ^
*
^ 1 ^
ilk ^ii
% j’%^
Figure 1. The interconnecting porous framework of Synthacerfi, a synthetical hydroxyapatite ceramic [Ca5(Po4) 3OH].
Figure 2. The diameter of the pores of the synthetical hydroxyapatite ceramic is constant (600 jam). The porosity amount to 80%. Magnification: - = 100 |im.
Acetabular Reconstructionin Revision Total Hip Arthroplasty:R.P. Pitto and D. Hohmann 171
RESULTS The follow-up examination of the 20 patients at 24 months (min. 18, max. 28) showed the improvement of the pain score (av. pre-op, 2.6, av. Fw.-up grade 5.2), of the function score (av. pre-op. 2.5, av. Fw.-up grade 4.9) and of the motion score (av. pre-op. 4.3, av. Fw.-up grade 5.2). The roentgenological analysis of the grafts showed the fusion within 4 months after the operation in all the cases and a gradual condensation (Fig.3), but one of them had evidence of some degree of bony resorption. The bone stock had increased in all the cases (av. pre-op. 3.1, av. Fw.up grade 4.9). There were no signs of implant loosening. One case underwent a re-revision 6 months after the implantation because of malposition of the component and recurrent luxation. The graft showed in this case fusion at the host bone interface and osteointegration of the synthetical hydroxyapatite ceramic.
Figure 3. A) Aseptic loosening of a Wagner surface cup with severe osteolysis and protrusion. B) Revision and reconstruction of the acetabulum with impaction grafting and a reinforcement ring with hook of Ganz. C) Roentgenological signs of fusion of the graft, remodelling and stable implantation 2 years after surgery.
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DISCUSSIO N The revision of the loose acetabular component with severe bony defect filled with a mix of synthetical hydroxyapatite ceramic and autologous bone grafts have proved to be of value. The results shows the good tolerance of the heterografts and a roentgenological evolution similar to that observed with pure autografts [5]. Further study is necessary, but these preliminary findings give rise to cautious optimism that this is a reliable method for acetabular revision, reconstruction and reconstitution. Careful pre-operative evaluation and peri-operative assessment to match bone defects, grafting patterns and reinforcement ring are of paramount importance. AKNOWLEDGMEN T The authors would like to express their thanks to Prof. K. Draenert and Dr. Y. Draenert, Centre of Orthopaedic Research, Munich, Germany, for the support of this study.
REFERENCES 1. 2. 3. 4. 5.
Charnley, J. In: Low FrictionArthroplastyof theHip, Springer, Heidelberg 1979, 20-24. Levai, J.P., Boisgard, S. Clin, Orthop.Rel, Res. 1996, 330, 108-114. Pitto, R.P. J. Bone Joint Surg.(Br.)(in print). Sloff, T.J.J.H., Huiskes, R., Van Horn, J. Acta Orthop.Scand. 1984, 55, 593-596. Stringa, G., Pitto, R.P., Di Muria, G.V., Marcucci, M. Int. Orthop. 1995, 19, 72-76.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EFFEC T OF SOLUTIO N AGEIN G ON SOL-GE L HYDROXYAPATIT E COATING S B. Ben-Nissan, C.S. Chai and K.A. Gross Department of Materials Science, University of Technology, Sydney P.O. Box 123, Broadway, N.S.W., 2007, Australia
ABSTRAC T Sol gel technology offers an alternative technique for producing a bioactive surface for improved bone attachment. Hydroxyapatite was synthesized using the sol-gel technique with alkoxide precursors and the solution allowed to age up to 7 days. Coatings produced on MgO substrates were characterised by differential thermal analysis, thermal gravimetric analysis. X-ray diffraction and atomic force microscopy. It was found that, similar to the wet method of hydroxyapatite synthesis, an ageing time is required to produce a pure hydroxyapatite phase. KEYWORD S Hydroxyapatite, sol-gel, alkoxide, ageing, coating, characterisation INTRODUCTIO N Hydroxyapatite is an established material for applications such as maxillofacial reconstructive surgery and non load bearing applications [1]. One of the currently used methods to overcome low mechanical properties of bulk hydroxyapatite is to coat substrates such as titanium and its alloys. To date, many processes have been investigated. These include dip coating into a powder suspension [2], electrophoretic deposition [3], sputter coating [4] and plasma spraying [5]. Of these processes, plasma spraying is used commercially. Thermal spraying, however, requires good process control to avoid decomposition at high temperatures and is limited to coatings thicker than 30 jim. An alternative coating method is sol-gel deposition. While commonly being used for producing glasses and oxides, it has more recently been utilized to produce other more complex materials as well as non-oxide ceramics. The advantages of the sol-gel technique include (a) increased homogeneity due to mixing on the molecular scale, (b) reduced firing temperatures of ceramics due to small particles with high surface areas (c) ability to produce uniform fine-grained structures [6]. 175
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Sol-gel techniques have been previously used to synthesize hydroxyapatite powder [7] and coatings [8,9]. This work illustrates the necessity of ageing time on the phase composition of the coating. METHOD S Solution Preparation 1.5x10"^ moles of calcium diethoxide (Kojundo Ltd., Japan) was suspended in ethanol and then dissolved in ethanediol (BDH Chemicals, Australia) with the aid of vigorous stirring in a glove box under dry nitrogen atmosphere. A second solution consisting of a stoichiometric amount of triethyl phosphite (Aldrich, U.S.A.) diluted in ethanol was prepared and added to the calcium bearing solution. Stirring was maintained for a period of ten minutes. Solutions were allowed to mature for 0 and 7 days before being used to make coatings. Coating Procedure Magnesia single crystal substrates (Zirmat, U.S.A.), 10x10x0.5 mm in size were chosen to study the coating quality, without the influence of the interactions with a reactive substrate such as titanium [10]. Substrates were ultrasonically cleaned in acetone and ethanol and then coated using a Headway Research (U.S.A.) spin coater. A volume of 0.5mL of solution was applied to the substrate and spun at 2500 r.p.m. for 10 seconds. Coated substrates were hydrolysed in an air oven (Labec, Aust.) at 70 C for 10 minutes, followed by prefiring at 500 C in a mufile furnace (Ceramic Engineering, Aust.) for 15 minutes. The coating/hydrolysis/prefiring procedure was repeated until 5 layers were deposited. After the final layer had been prefired, the coated substrates were heated at 200 C/hr to 1000 C and soaked for 15 minutes followed byfiunacecooling. Characterisation Technique s X-ray diffraction (Siemens D5000, Germany) was conducted on coated substrates using CuKa radiation and a glancing angle geometry. This attachment was necessary due to the small coating thickness. Scan parameters included a scan range of 28 to 40 29, step size of 0.02 , step time of 5 seconds and X-ray incident angles between 0.5 and 5.0 . Thermal analysis techniques differential thermal and thermogravimetric analysis (DTA and TGA) were performed using a SDT 2960 simultaneous thermal analyser (TA Instruments, USA). Samples were heated at 10 C/min to 500 C, held for 15 minutes and then heated to 1200 C at 200 C/hr (3.33 C/min). This heating rate was chosen to replicate the heating schedule. The morphology of the coated substrates were examined using a Park Scientific Instrument (Autoprobe LS, U.S.A.) atomic force microscope (AFM). RESULT S AND DISCUSSIO N The use of a pre-firing stage at 500 C facilitates coating build-up. It also removes the volatile species allowing the rapid heating rate to sintering conditions. Thermal shock is minimized due to the small coating thickness and the relatively small amount of material deposited. Hence, the thin coatings have a low susceptibility to thermal shock cracking, and facilitates ease of gas (including alchohol) removal. In addition, the thermal gradient within the coating is very small and the sintering conditions in all locations of the coating are similar.
Effect of Solution Ageing on Sol-Gel Hydroxyapatite Coatings: B. Ben-Nissan et al.
177
Fired coatings appeared quite uniform except towards the edges where it was thinner (seen as interference fringes). This would be a thinning of the coating due to the edge effect. Complete coverage is thus dependent upon the wettability and geometry of the object. Ageing The X-ray diffraction patterns for coatings produced after 0 and 7 day ageing periods are shown in Figure 1. Hydroxyapatite is evident after 0 days ageing, however, the presence of CaO (JCPDS 4-777) and other peaks suggests that the reaction has not reached completion. The coating produced after an ageing period of 7 days appears to consist solely of hydroxyapatite. Thus, it is evident that an ageing period is necessary to allow the different species present in the coating solution to mix thoroughly. Given the complex kinetics of this system, it is possible that some chemical reactions may take place during this maturing period. This ageing phenomena is similar to the ripening procedure used in the "wet method" to produce a stoichiometric hydroxyapatite [11]. Thermal analysis of the hydrolysed gel produced after maturing time of 7 days exhibited an endothermic peak at 110 C and three exothermic peaks at 216, 430 and 550 C. The large endotherm corresponds to the evolution of residual solvent and adsorbed moisture. This is followed by two large exothermic reactions at 216 and 430 C respectively. These reactions correspond to the formation of chemical bonds through condensation and polymerisation as well as the evolution of residual water and/or alchohol. This has also been reported with zirconia gels [12]. A smaller exothermic reaction occurs at 550 C. It is believed that this reaction represents the crystallisation of hydroxyapatite [13]. The vertical translation observed on the DTA/TGA curve at SOC’C represents the 15 minute pre-firing heat treatment. Surface Morphology The surfaces of the coatings were examined using AFM. The coatings were crack free and consisted of 2 distinct regions. The surface was covered with small grains, approximately 200nm in size. These smaller grains exhibited a "cauliflower-like" surface which was broken up by larger grains, approximately 800nm in diameter. These were observed at random separations across the coating surface and can be identified as peaks in figure 3a and lighter regions in figure 3b. It is possible that the larger grains had formed as a result of exaggerated grain growth. 1
’
1
’
I
’
i
1 S
^ ’c
I e 1
28
30
1
32
34
, 36
0 da>{s ageing 38
40
Degrees (20)
Figure 1. X-ray diffraction pattern of coatings produced from solution matured for 0 and 7 days.
200
400
600
800
1000
1200
Temperature (^C)
Figure 2. DTA/TGA plots for hydrolysed gel matured for 7 days.
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A B Figure 3. Atomic Force Microscope scans of coatings, using solutions matured for 7 days. CONCLUSION S Hydroxyapatite coatings have been produced via the sol-gel route. It was found that to induce the formation of a coating that is predominantly hydroxyapatite, solutions should be aged prior to use. AFM examination revealed the presence of two distinct regimes consisting of grains 200nm and 800nm in size respectively after being sintered at 1000 C.
REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13.
de Groot, K., de Putter, C, Sillevis Smitt, P.A.E. and Driessen, A.A. In : Scienceof Ceramics,Brit. Ceram. Soc, Stoke on Trent, 1981, 433-437. Lacefield, W.R., Ann.NY. Acad. Sci.,1988, 523, 72-80. Ducheyne, P., van Raemdonck, W., Heughebaert, J.C. and Heughebaert, M., Biomater., 1990, 11, 244-54. Ong, J.L., Lucas, L.C., Lacefield, W.R. and Rigney, E.D., Biomater.,1992,13, 249-254. Gross, K.A. and Bemdt, C.C., J. Biomed.Mat. Res.,to be published in 1997. Johnson, D.W. and Gallagher, P.K. In : CeramicProcessingbeforeFiring, John Wiley and Sons, U.S.A., 1978. Masuda, Y., Matubaram, K. and Sakka, S., J.Ceram. Soc. Japan,1990, 98, 1266- 1277. Chai, C, Ben-Nissan, B., Pyke, S. and Evans, L., In.SurfaceModificationTechnologiesVII, T.S. Suddshan, K. Ishizaki, M. Takata and K. Kamata, Eds. Cambridge University Press, UK, pp. 509-525, 1994. Deptula, A., Lada, W., Olczac, T., LeGeros R.Z. and LeGeros J.P., In : BioceramicsVol. 9 University Press, Great Britain, 1996, 313-316. Chai, C. and Ben-Nissan, B., J. Aust. Ceram. Soc, 1993, 29 (1/2), 81-90. Osaka, A., Miura, Y., Takeuchi, K., Asada, M. And Takahashi, K., J. Mat. Sci.:Mat. in Med.,1991,2,51-55. Ben-Nissan, B., Anast, M., Bell, J., Johnston, G., West, B.O., Spiccia, L., de Villiers, D. and Watkins, I., Proc. 1stInt. Symp. Sci. ofEng.Cer.,S. Kimura and K. Niihara Eds., MikawaHaitsu,Koda, 1991,25-29. Gross, K.A., The amorphous phase in hydroxyapatite coatings, PhD dissertation, 1995, State University of New York at Stony Brook, USA.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
IONI C CEMENTS : INFLUENC E OF LIQUID/SOLI D RATI O ON POROSIT Y AND MECHANICA L PROPERTIE S F.Betchem*, P. Michaud*, F. Rodriguez*, Z. Hatim**. * Laboratoire de Pharmacie Galenique, Faculte des Sciences Pharmaceutiques, 35, chemin des Maraichers 31062 Toulouse Cedex ** Laboratoire de Chimie-Physique, Universite Chouaib Doukkali, Faculte des Sciences, B.P. 20 El Jadita-Maroc. ABSTRAC T Ionic cements are widely studied in orthopaedics and the biocompatibility, nontoxicity, and partial resorbability of hydroxyapatite are well known. In this study, we analyzed the effect of varying the liquid/solid ratio over a short range (0.40 - 0.50) with an ionic cement which has Ca/P atomic ratio 1.63. We examined the influence on axial and diametral tensile strengths, hardening and porosity. We observed that, when the amount of liquid is increased, the excess water is not used for the reaction, and occupies an interstitial position in the solid. So the porosity and the hardening time increase while the tensile strength decreases. KEYWORD S Ionic cements, hydroxyapatite, porosity, hardening, mechanical resistance. INTRODUCTIO N Ionic cements are being increasingly studied in orthopaedics. Their interest is due to their composition and structure, which are close to those hydroxyapatite (Caio(P04)6(OH)2, HAp), their easy utilization, and their non-exothermicity [1,2]. In order to reach total rehabitation, a cement has to present sufficient porosity to allow diffusion of body fluids but an increase in porosity is frequently correlated with a decrease in mechanical resistance. This paper shows how the variation of the liquid/solid ratio over a short range (0.40 - 0.50) leads to modifications of the mechanical and physical properties in vitroand in a moist atmosphere at 37 C. The value Ca/P of 1.63 was chosen in order to prepare a non stoechiometric apatite and to be sure to obtain a total reaction of all the initial reagents. MATERIAL S AND METHOD S Cement paste was prepared by addition of a liquid containing calcium and phosphate ions to a mixture of solid Calcium Phosphates (table 1) Ca4(P04)20, TTCP and a-Ca3(P04)2, a-TCP [3]. In the cement formula, the Sodium Glycerophosphate (NaGP) is useful for improving the paste homogeneity. The paste was placed in silicone molds (9 mm diameter and 5 mm thickness) for 10 minutes in a moist environment at 37 C. Then, the cylindrical samples were removed from the molds and kept in a moist environment at 37 C for 96 hours. Before analysis, each sample was dried in an infrared balance at 110 C for 10 minutes. These samples were used to determine the mechanical properties with a Diametral Tensile Strength (DTS) machine PHARMATEST PTB 311 and computerized single punch machine KORSH Ekod for Axial Tensile Strength (ATS). The specific area and total pore volume of the samples were measured by a nitrogen high speed surface area 179
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Powder
12000 7
8.68 g 1.32 g
TTC P and a-TCP NaGP
E E 10000 8000
Liquid
Ca(OH)2 H3PO4 H2O
Ca/P=1.635 liq/sol=0.4 3 "Ca/P=1.635 liq/sol=0.4 5 - Ca/P=1.635 Iiq/sol=0.5 0 B
6000 +
0.12g 0.29 ml to 100%
.5 4000 % 2000 24 30 36 Time (min)
Table 1: Composition of the cement
Graph 1 : Hardness versus time (Ca/P = 1.63, hquid/solid = 0.43) and pore size analyzer Quantachrome NOVA 1000 for specific area; mercury pore size analyzer Micromeretic autopore 11-9215 for total pore volume. To follow the hardening, we used a texture analyzer TAXT2. Cement paste was introduced into a mold of 12 mm diameter and 12 mm height which was placed in a steel block of great thermal inertia to keep the cement temperature at 37 C during measurements. One measurement was made every 3 minutes for one hour with a 1 mm punch in the following conditions : 2 mm/s for downward and rise speed; penetration of 5 mm. For each cycle we took the maximum strain to plot the hardening curve.
RESULT S AND DISCUSSIO N The graph 1 shows the setting and the hardening evolution of a cement with three different liquid/solid ratios in a moist atmosphere at 37 C. In the three curves, there is two main stages before reaching to hydroxyapatite. The part A corresponds to the formation of brushite (CaHP04, 2H2O, DCPD) by acidic and basic reactions between TTCP, a-TCP and phosphoric acid contained in the solution. 120 J
ir no f B
S 100 ^
^
^ g.
51 *Z 49 ^
70 f 60
0.35 0.43 045 Uquid/solid ratio
93.37 53’ 83.53 49
90 +
59 111.94 + 57 ’ 56 101.26 55 ^ 54
47 H-
-H
0.4
0.45
-+0.5
45 0.55
Liquid / solid ratio
Graph 2 : Tensile strength after 96 hours in Graph 3 : Porosity and specific area versus moist atmosphere at 37 C liquid/solid ratio Liquid/Solid ratio 0.40 0.43 0.45 0.50
Apparent density 1.56 1.50 1.48 1.38
Measured density 2.46 2.46 2.52 2.49
Theoretica l % of open porosity 37 40 42 45
Residual humidity rate (% ) 12.50 14 15 15.50
Table 2: Theoretical open porosity and residual humidity rate of sample differents from their liquid/solid ratio.
Ionic Cements:Influenceof theLiquid/SolidRatio:F. Betchemet al. 181 00
0.3
’JS^:
0.2 2 0.1
S 0.0-H^ ^ 00
o
H-3
rrr rr 10000
1000
100
Diameter (Angstroms)
s (view by Figure 1: Distribution of pore size by mercury Figure 2 Porosity and HA p needle intrusion in 0.43 liquid/solid ratio sample. SEM X 20000) in 0.43 liquid/solid ratio sample. The part B correspond s to the formation of octocalciu m phosphate (Ca8H2(P04)6, 5H2O, OCP ) [4] which in turn gives HA p after several hours. The curves also show that the setting and hardening times increase as the liquid/solid ratio increases . In graph 2, we observe the large modificatio n of mechanica l propertie s when the amount of liquid is increased . Graph 3 shows the rise of specific area and porosity as the liquid/solid ratio increases . The rise of porosity is due to excess water which occupies an interstitial position during hardening in a moist environmen t at 37 C. When the liquid is removed by drying in an infra red balance it contribute s to the porosity and specific area. Table 2 shows the results of density and humidity rate measurements . The differenc e betwee n apparent and measured density is explaine d by the open porosity. Moreover the rise of the residual humidity rate as the liquid/solid ratio increase s shows that the reaction which leads to HA p uses the necessar y amount of water and the excess of water occupies an interstitial position in the solid. The measuremen t of excess of liquid consists in measuring the residual humidity rate. The results achieve d by nitrogen and mercury pore size analyzers are compatible with the theory: in each sample the rate of open porosity depends on liquid/solid ratio and is included betwee n 45 and 60%. The decrease of mechanica l resistance is correlate d with porosity [5,6]. Another fact can contribute to the decrease of the mechanica l propertie s : a large amount of liquid slows down the crystallizatio n by modifying the different calcium phosphate s successivel y formed. The mercury pore size analyze of samples shows that there is two pore families for each liquid/solid ratio : the main family is situated near 1000 Angstroms diamete r (figure 1) and increase s as the amount of liquid increases . The vision of figure 2is a SEM photograph (x 20000) illustrating the porosity and the needle s of HA p inside the cement (Ca/P =1.63 and liquid/solid ratio = 0.43). CONCLUSIO N The liquid/solid ratio is an important paramete r in the preparation of ionic cements . The amount of water required is determine d by the Ca/P ratio. Excess liquid is not used for the reaction but occupies an interstitia l position in the solid. When the excess of liquid increases , the porosity and the hardening time increase and the tensile strength decreases . To achieve sufficient porosity while maintaining correct mechanica l properties , the formulation of those cement s must be optimize d and the precise conditions of their utilization must be determined .
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ACKNOWLEDGMEN T We thank TEKNIMED (B.P. 60, Vic en Bigorre, France) for financial support of the work and for providing of powder raw materials. One of the authors (F.B.) wishes to gratefully acknowledge Pr J.L. Lacout for encouragement and support.
REFERENCES
1. E.W. BROWN, L.C. CHOW, J. Dent Res.,62, 1983, 672. 2. A. A MIRTCHI, J. LEMAITRE, E. MUNTING, Biomaterials,11, 1990, 83-88. 3. J.L. LACOUT, E. MEJDOUBI. Procede d’obtention d’hydroxyapatite phosphocalcique, application au comblement osseux ou au moulage de pieces et produits utilises. Brevet Fr92.09019/PCT/FR. 4. J.L. LACOUT, E. MEJDOUBI, M. HAMAD, J. Mater Sci, Mater.Med, 7, 1996, 371374. 5. R.WmCE, J. Am. Ceram. Soc, 79, 1993, 1801-1808. 6. O. BERMUDEZ, M.G. BOLTONG, F.C.M. DRIESSENS, J.A. PLANELL, J. Mater. Sci. Mater.Med, 4, 1993, 389-393.
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
SINTERIN G BIOCERAMIC
AND S
THERMA L
DECOMPOSITIO N
OF
HYDROX Y APATIT E
J. Cihlar and M. Trunec Department of Ceramics, Institute of Materials Engineering, Brno Technical University, Technicka 2, 616 69 Brno, Czech Republic
ABSTRAC T In the course of high temperature treatment of injection moulded hydroxyapatite ceramics (HA) sintering, grain growth and thermal decomposition of HA to tricalcium phosphate took place. The sintering was finished at 1573 K. The grain growth started at 1500 K and the thermal decomposition started at 1623 K. The activation energy of grain growth was 215" kJ/mol, that of thermal decomposition 283.5 kJ/mol. The optimum sintering temperature was found at 1473 K. KEYWORD S Hydroxyapatite ceramics, thermal decomposition, gram growth, sintering, kinetics INTRODUCTIO N Properties of sintered hydroxyapatite ceramics, namely their mechanical and biochemical properties depend on the physical and chemical structure of HA [1]. This structure is dependent on processing parameters of HA ceramics, namely on conditions of thermal treatment. In the course of thermal treatment sintering, grain growth and thermal decomposition of HA take place [2]. Data published about the physical and chemical behaviour of HA ceramics in the course of thermal treatment are not consistent. Nonuniformity namely has to do with an optimal sintering temperature and mechanism of thermal decomposition of HA ceramics [3, 4, 5]. In this contribution the authors try to make a kinetics and mechanism of thermal decomposition more clear and to give optimal tempei-ature of HA sintering. MATERIAL S AND METHOD S The samples of HA ceramics were prepared by ceramic injection moulding. Materials used and sample processing had been published [2]. Thermal treatment of HA samples was made in a superkanthal furnace in air atmosphere of 25% relative humidity at the temperature range from 1373 K to 1773 K for 1 to 22 hours. The microstructure and microanalysis of HA specimens were determined by SEM on a JXA-840 microscope equipped with an energy dispersion analysator (Link). The phase composition of HA ceramics was established by X-ray diffraction analysis on a D-500 diffractometer (Siemens). For quantitative phase analysis the part of spectrum in the range of 30 to 50 for 26 was used. The diffraction of (121) + (211), (300) and (301) planes were used for HA content determination, the diffraction of (123) + (254), (434) + (264) + (401) and (400) planes were used for the determination of tricalcium phosphate (TCP) content. 183
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RESULTS AND DISCUSSION Sintering The dependence of linear shrinkage of hydroxyapatite ceramics on sintering temperature is given in Figure 1. In the temperature range from 1373 to 1573 K linear shrinkage increased with increasing temperature. A maximum shrinkage (16%) was obtained at the temperature of 1573 K. With this maximum linear shrinkage the hydroxyapatite ceramics had relative density of 98%. Sintering at the temperature above 1573 K (in the air atmosphere) did not result in further increase of the density of hydroxyapatite ceramics.
1100
1200
1300
1400
SINTERING TEMPERATURE [K]
Figure 1 The dependance of the shrinkage of HA ceramics on sintering temperature
Figure 2a Microstructure of injection moulded HA ceramics sintered at 1373 K for 1.5 hour SINTERING TIME [hour] ’ \m Y
2
4
1
1
’
6
1
-I
8
T
1
Sintering Time Sintering Temperatur e
1
\
iy
1
10
A
Y-
0 1000
* t-^t
1200
1
1
-1 5.
1400
\
1600
SINTERING TEMPERATURE [K]
Figure 2b Microstructure of injection moulded HA ceramics sintered at 1773 K for 1.5 hour
Figure 3 The dependence of the grain size of HA ceramics on sintering time and temperature
Sintering and ThermalDecompositionof Hydroxyapatite Bioceramics:J. Cihlar and M. Trunec 185
Grain Growth The microstructure of the sintered HA ceramics is shown in Figure 2. The average grain size of the HA ceramics sintered at 1373 K for 1.5 h was about 1 |Lim (figure d). The HA ceramics sintered at 1773 K for 1.5 h contained grains of the average size of 16 fin (^Figure 2b). The growth of grain size depended, above all, on sintering temperature. The most pronounced grain growth was observed in the temperature range from 1573 K to 1673 K (see Figure 3). Thermal decomposition The loss in weight (due to the loss of water) of hydroxyapatite ceramics started ai the temperature of 1373 K. A negligible shift of diffraction lines of HA iccompanied b> weight loss of HA was caused by formation of oxyapatite (OA). HA-OA system (termed as hydroxyoxyapatite [4]) was stabile for 15 h at temperature 1573 K. The presence of crystalline a-TCP was detected until sintering for 2 hours at 1623 K. The course of thermal decomposition is perceptible from Figure 4.
Figure 4 Section by the layer of TCP on the surface of HA ceramics sintered at a) 1573 K for 22 hour b) 1773 K for 8 hour %
200
1 1 1 1 1 1 1
1
LU
150 Q. O
^
CO CO
UJ
z "^ o X
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1
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-4 ^ J
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^
/ Oi [ 0
1
-5
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Ep = 283. 5 kJ/mo l
^^^^"^^ 1
100
O
r - - i-
I ’ 1 ’ 1 1
1
1 1
I
.
I
.
I
.
200 400 600 800 100012001400 SINTERING TIME [min]
Figure 5 The dependence of the thickness of TCP layer on the sintering time
1
0,56
0,57
.
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0.58
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I
0,59
.
I
0,60
.
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lOOOyT [1/K] Figure 6 The temperature dependence of the rate constant of HA decomposition
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The thermal decomposition started on the surface of HA ceramics. At first, islands of TCP appeared (Figure 4a). These islands were connected together in the compact TCP layer growing into the inside of HA-OA ceramics (Figure 4b). The growth of thickness of TCP layer (x) with time (t) was described by parabolic rate law x=(Kt)^^^ (see Figure 5) [6], where K is a factor of proportionality. The rate of the thermal decompositicm of HA-OA is then controlled by diffusion of reaction products (water) through the layer of TCP. The activation energy of thermal decomposition of HA-OA was 283,5 kJ/mol (see Figure 6). In the temperature range from 1623 to 1773 K, thermal decomposition products of HA ceramics contained only TCP. Traces of tetracalciumphosphate appeared at 1773 K. The thermal decomposition of HA ceramics could be described by equations: 2 Ca5(P04)30H o Caio(P04)60 + H ^ (1373 - 1723K)
(1)
Ca,o(P04)60 <^ 3 Ca3(P04)2 + CaO
(1623 - 1723 K)
(2)
3 Ca3(P04)2 + CaO o Ca4(P04)20
(above 1773 K)
(3)
CONCLUSION S In the course of thermal treatment of HA ceramics thermal decomposition of HA took place at the temperature higher than 1623 K. The layer of reaction product - TCP was formed on the surface of HA samples. The kinetics of thermal decomposition was described by parabolic rate law. The activitation energy of thermal decomposition was 283.5 kJ/mol. Optimum sintering temperature of injection moulded HA ceramics was found at 1473 K Acknowledgements The autors thank Mrs Drahoslava Janova for providing the SEM photographs and Dr. Antonin Buchal for providing X-ray diffraction analyses. This project was supported by Czech Ministry of Education Fund. REFERENCE S 1. LeGeros, R.Z. and LeGeros, J.P. In: An Introduction to Bioceramics, World Scientific, London 1993,139 (Eds.: Hench, L.L. and Wilson. J.). 2. Cihlar, J. and Trunec, M., Biomaterials1996, 17, 1905-1911. 3. Zhou, J. Zhang, X. Chen, J Yeng, S. and De Grot, K., J. Mater.ScL, Mater.Med. 1990, 4, 83-85 4. Wang, P.E. and Chaki, T.K., Biomaterials:Materialsand Application,ACerS, Ohiol995, 225 ( Eds.: Fischman, G. and all.) 5. Locardi, B., Pazzaglia, V.E., Gabbi, C. and Profilo, B.,Biomaterials14, Oxford 1993, 437441 6. Kingery, W.D., Bowen, H.K. and Uhhnann, D.R., Introductionto Ceramics,New York 1976, Wiley, 386
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
AN ELABORATIO N OF THE NEW DISSOLUTIO N MECHANIS M FOR APATIT E Sergey V. Dorozhkin Research Institute of Fertilisers and Insectoflingicides Kudrinskaja sq. 1-155,123242 Moscow D-242, Russia
ABSTRAC T Used methods of scanning electron microscopy, Auger-electron spectroscopy and IR-reflection spectroscopy, a new information about chemical dissolution of fluorapatite in phosphoric acid was obtained and analysed. The dissolution was found to occur incongruently on the very thin surface layer of fluorapatite, via previous leaving of fluorine followed by calcium. The experimental results obtained were extended with literature analysis. A new chemical mechanism for apatite dissolution in acids was proposed as a result. The mechanism for the first time described the chemical irregularity of the dissolution process at the atomic (ionic) level. KEYWORDS : apatite, dissolution, chemical mechanism. INTRODUCTIO N A pure calcium apatite is widely used as a bioceramics in medicine, because its chemical composition is very close to the inorganic part of animal and human bones and teeth. Unlike natural bones, the artificial bioceramic substitutes based on apatite (or covered with apatite) are not alive; so they are in constant interaction with different biological and chemical substances inside a human body. That is wiiy, the fundamentals of apatite dissolution are so important for the correct simulation of processes occurring with the substitutes in vivo. There are a large number investigations, devoted to apatite dissolution kinetics in the acidic medium [1]. Different models for the apatite dissolution process were proposed as a result. The polynuclear dissolution model [2] seems to be the most likely dissolution mechanism among them. But non of the models described chemical transformations occurring on the apatite crystal surface during dissolution. For nowadays the only equation, used for apatite dissolution in acidic medium, is the following chemical reaction [1,2]: Ca5(P04)3F + TKT = 5Ca^’^ + 3H2P04" + HF
(1)
Reaction (1) means that seven hydrogen ions are necessary to dissolve one « molecule » of apatite and reports nothing about a possible chemical mechanism. But as soon as reaction (1) occurs, the chemical mechanism at the atomic (ionic) level must exist. MATERIAL S AND METHOD S 187
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Crystals of natural fluorapatite and pure phosphoric acid were used for the experiments. 0.1 - 1 g of crystals were dissolved in 200 - 300 ml of 2 - 7 M H3PO4 solution during 3 - 20 s under temperature of 50 - 90 C and Reynolds hydrodynamics (Re) of 2000 - 3000. After being slightly dissolved, the crystals were quickly separated from the acid solution by filtration through a filter with porosity of 1.0 jim, followed by washing them in acetone during 20 - 25 s and drying in air at room temperature. After being washed and dried, the crystals were studied with a scanning electron microscope (SEM) JSM 35CF (secondary electron mode, acceleration voltage of 15 kV), IRreflection spectroscope Specord 75 (in the range of 2000 - 200 cm"^) and Auger electron spectroscope (AES) Jump 10 with the standard measurement techniques [3], published elsewhere [4]. After being investigated the crystals were washed in water during 2 - 5 min and the above acidic treatment followed again. The results obtained for the treated crystals were compared with the same of initial fluorapatite. RESULT S AND DISCUSSIO N Table 1 shows position of some adsorption bands of fluorapatite in the IR spectra for the initial crystals (row 1), the crystals treated in acid followed by washing them in acetone (row 2), and the same as row 2, but with additional washing in water during 1 min (row 3). Three new absorption bands at 1100, 1046 and 668 cm"^ were found for the treated crystals (Table 1, row 2). Two of them (1100 and 1046 cm’^) were found in the libration range of phosphate groups [1, 5]. They pointed out to some chemical changes occurred with the bands at 1116 and 1078 cm"^ respectively. The third band (668 cm’^) was new. It appeared in the libration range of hydroxyl groups [1,5]. Similar results of AES are given in Table 2. The very thin surface layer, equal to the Augerelectrons formation and scattering (approximately 1-3 nm), was found to have lost all fluorine and atomic ratio of Ca : P decreased from 1.67 – 0.05 (initial apatite) to 1.30 – 0.05 for the apatite treated [4]. The latter value appeared to be close to that of octacalcium phosphate Ca8H2(P04)6’5H20 (Ca : P = 1.33) [1]. Thus, both methods have pointed out to some chemical changes happened on the surface of fluorapatite treated. The results of SEM showed clear evidence for the appearance of a conducting layer on the surface of apatite treated in acid [4]. Unlike initial apatite (it is dielectric), the apatite treated was found to have enough surface conductivity to be studied in the SEM directly [5]. On the other hand, crystals of acidic calcium phosphates CaHP04 and Ca(H2P04)2 were also found to have surface conductivity good enough to be studied in the SEM directly. Thus, the results of SEM also pointed out to some changes occurred on the crystal surface of apatite during dissolution. To extend information about the surface layer obtained, some additional results were taken from literature. Apatite was found to be charged positively in acidic mediums [6, 7] as a result of chemosorption of protons [6] or calcium and protons [7] from the solution. Another results of the
Table 1. The wave numbers (iij, cm"^) of some IR spectra bands for fluorapatite rowl. Initial apatite
1116
-
1078
row 2. Etched apatite washed in acetone
-
1100
-
1046
668
-
1078
-
668
row 3. Etched apatite washed in acetone and water 1116
An Elaboration of the New Dissolution Mechanismfor Apatite: S.V. Dorozhkin 189
Table 2. AES results of Ca/P ratio on the surface of fluorapatite (according to the rows in Table 1) row 1
row 2
row 3
1.67 – 0.05
1.30 – 0.05
1.65 – 0.10
sequence of atomic (ionic) dissolution of fluorapatite were also taken from literature. Apatite was found to dissolve incongruently; moreover, the following sequence of ionic dissolution was established: first fluorine, next calcium and afterwards phosphate [8, 9]. DISSOLUTIO N MECHANIS M As soon as apatite was found to be dissolved in acidic mediums mainly [1], the role of protons was noteworthy. According to the above [8, 9], fluorine dissolved first. But the results of IR-spectroscopy (Table 1, row 2) pointed out to hydroxyl incorporation into the crystal lattice of fluorapatite. Hydroxyl is known to be easily replaced with fluorine and back in the channels parallel to the c axis [1]. So, it is logical if a first chemical equation describes the exchange of fluorine with hydroxyl or water: Ca5(P04)3F + H2O + H" = Ca5(P04)3(H20)’’ + HF
(2)
Here incorporation of water is chosen instead that of hydroxyl, because in acidic medium investigated incorporation of basic hydroxyl is very unusual. Protons are supposed to be previously chemosorbed on the surface of apatite crystals (positive charge formation [6, 7]) and used as a catalyst [2] for removing of fluorine and water incorporation instead. A possibility of such substitution is taken from literature [1, 10]. Molecule of water consists of two 0-H bonds; each of them is believed to be able to form the libration band at 668 cm’* (the results of IR-spectroscopy). As soon as a local positive charge on apatite has already appeared (2), an interaction of other protons becomes more difficult. So, dissolution of one calcium cation might be supposed: 2Ca5(P04)3(H20)’’ = 3Ca3(P04)2 + Ca^’’ + H20
(3)
Reaction (3) corresponds completely with the results of [8, 9] (calcium dissolves ahead of phosphate) and partly with the results of AES, because for intermediate Ca3(P04)2 obtained, Ca : P atomic ratio is equal to 1.50 (i.e.the ratio is already not 1.67 – 0.05, but it is still not 1.30 – 0.05). A fiirther interaction between protons and apatite surface might be described as follow: Ca3(P04)2 + 2H^ = Ca^"’ + 2CaHP04
(4)
Here next calcium is replaced with two protons. This equation completely corresponds with the results of [8, 9], SEM (a conductive layer appears, because CaHP04 is formed) and IRspectroscopy (changes with phosphate bands at 1116 and 1078 cm"* have happened) and partly with ones of AES, because for CaHP04 the Ca : P ratio is afready 1 : 1 (i.e.below 1.30 – 0.05). A reasonable explanation of this difference is easy: a mixture of CaHP04 and Ca3(P04)2 is formed on the surface of apatite treated in acid. Both of these substances were proposed as precursors of apatite formation [11]. So, for the above reasons reaction (4) contradicts to nothing. At last, phosphate ions also dissolve according to one of the following reactions:
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CaHP04 + 2H’’ = Ca^"’ + H3P04 CaHP04 + tf = Ca^^ + H2P04’
(5) (6)
As soon as the above dissolution occurs in the water medium, all the ions and molecules mentioned are hydrated, but the hydration effect is omitted here for simplicity. SUMMAR Y Thus, a new chemical mechanism for acidic dissolution of apatite has been elaborated. It is based on the chemical logic and takes into consideration all experimental results available for nowadays [1-11] and references therein. It consists of five successive chemical reactions (2) - (6) to be used for description of the dissolution chemistry of apatite instead of reaction (1). The mechanism for the first time describes a chemical irregularity (incongruence) of dissolution for different ions at the atomic (ionic) level. The above mechanism should be also usefiil for chemical simulation of dental caries process. REFERENCE S 1. EUiott, J.C. Structureand Chemistry of the Apatites and OtherCalcium Orthophosphates. Springer. Amsterdam - London - New York - Tokyo, 1994, p. 389. 2. Christoffersen, J., Christoffersen, M.R. and Johansen, T. J. Cryst. Growth 1996,163,304310. 3. Goldstein, J. and Yakowitz, H., Eds. Practical Scanning ElectronMicroscopy.FlQvmm.^Qw York, 1975,356. 4. Dorozhkin, S.V., Nikolaev, A.L., Melikhov, I.V., Saparin, G.V. and BUadze V.G. Scanning 1992,14,112-117. 5. Knubovets, R.G. Reviewersin Chem.Engin. 1993, 9, 112 -140. 6. Bell, L.C., Posner, A.M. and Quirk, J.P. J. Colloid InterfaceSci. 1973, 42, 250 - 261. 7. Chander, S. and Fuerstenau, D.W. J. Colloid InterfaceSci. 1979, 70, 506 - 516. 8. Krivoputskaya, L.M., Lemina, L.M. and Gusev, G.M. Trans.SiberiaBranch Acad.Sci. USSR, ChemicalSeries.1978, 4 Issue 2, 65 - 72. 9. Tarantsova, M.L., Kulikov, B.A., Chaikina, M.V., Kolosov, A.S. and Boldyrev V.V. Trans. Siberia Branch Acad Sci. USSR, ChemicalSeries.1980, 9, 55 - 62. 10. LeGeros, R.Z., Bond, G. and Legros, R. Calcif Tiss.Res. 1978, 26, 111 - 118. 11. Francis, M.D. and Webb, N.C. Calcif Tiss.Res. 1971, 6, 335 - 342.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
ULTRASTRUCTURA L STUD Y OF LON G TER M IMPLANTE D CA- P PARTICULAT E MATERIAL S INT O RABBI T BONE S Dupraz A., Rohanizadeh R., Delecrin J., Pilet ?*., Passuti N., Daculsi G. UPRES EA 2159, Faculte de chirurgie dentaire, Place A. Ricordeau, 44042 Nantes Cedex, France * Service commun de microscopic et de microanalyse, ensemble sante - Nantes.
ABSTRAC T Long term implantation study of Ca-P particles showed that bone integration of biodegradable particulate materials was not completed at 78 weeks. Many ceramic particles were totally incorporated in bone, but some remained loose in the bone marrow or were phagocytized by mononucleated or multinucleated histiocytic cells. The particles totally integrated in bone were protected from a degradation process, because intimately linked to the bony tissue. While the ones with contact to medullary cavities showed a dissolution fringe, with less mineral content at their periphery, as observed by SEM. TEM confirmed that the grain core was more compact because of presence of precipitated microcrystals, whereas no precipitation was observed at the grain periphery. Microprobe analysis and microdiflfraction identified the crystals of the grain periphery as HA, suggesting that P-TCP was already dissolved. Moreover, microdiffraction patterns indicated that the phagocytized particles were also HA. KEYWORD S Material degradation in vivo,particulate calcium phosphate, electron microscopy INTRODUCTIO N The development of minimal invasive surgical techniques requires the need of injectable bone substitutes. In this purpose, it could be beneficial to associate a viscous polymeric carrier with bioactive ceramic granules to make them injectable. Although the literature refers many cases, where implantation of particulate materials gave rise to severe inflammatory reaction and caused rejection of the implants [1], it was reported that bioactive HA particles did not elicit such adverse cellular reactivity, when implanted in bone [2]. Thus, it was of particular interest to study the behaviour of an injectable composite paste, made of bioactive Ca-P particles and of a cellulosic vehicle, after long term bone implantation. This paper focuses on the biodegradation/bone substitution process that the calcium phosphate particulate paste undergoes during implantation. The cellular reaction as well as the bone ingrowth in regard to these implants have been described elsewhere [3]. MATERIAL S AND METHOD S Biphasic Calcium Phosphate (BCP) powder (60% HA, 40% P-TCP) was obtained by crushing macroporous blocks and was fiirther sieved mechanically between 80 and 200 |Lim. It was then used as the mineral phase for an injectable biomaterial. The composite paste made of 60% BCP 191
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Figure 1 : SEM picture of the new formed bone incorporating BCP particles
Figure 2 : SEM picture of a BCP grain in contact with a medullary cavity
powder and 40 % HydroxyPropylMethylCellulose (HPMC) aqueous solution (2% w/w) was steam sterilised (121 C, 20 min, 1,05 mBar) and injected into femoral cylindrical cavities (6 x 10 mm) of 36 adult rabbits. Six animals were sacrificed at 78 weeks after surgery. Seven micrometers thick undecalcified sections were stained by Hematoxylin-Eosin and by Solochrome Cyanine to assess the cell reaction and the bone ingrowth towards the implants . Material degradation/integration process was evaluated by electron microscopies, SEM (Jeol, JSM 6300 operating at 15 kV) and TEM (Jeol, JEM 200CX, operating at 200 kV). X-ray microanalysis was used to assess the chemical changes occurring in the ceramic grains between before and after implantation and to evaluate Ca/P ratios, whereas electron diffraction indicated the modifications of the crystalline structure. RESULT S AND DISCUSSIO N After 78 weeks implantation, the resorption/bone substitution process of the injectable biodegradable composite was still not completed. The Ca-P particles resorption was at different stages, depending if the ceramic particles were : 1. integrated in new bone trabeculae, 2. loose in the bone marrow, 3. or phagocytized by histiocytic cells.
Figure 3 : TEM picture of a ceramic grain, loose in the bone marrow, (c : grain core; p : periphery)
Figure 4 : Electron diffraction pattern of the apatitic precipitated microcrystals in the core of the ceramic grain
UltrastructuralStudy of Long Term ImplantedCa-P ParticulateMaterials: A. Dupraz et al.
193
Ca/P weight ratio BCP before implantation Periphery
1.95 –0.16 1.88 –0.14
BCP after implantation Centre 2.00 – 0.04 Table 1 : Weight Ca/P ratios obtained by electron microprobe analysis 1. The resorption process of the ceramic was limited in case of bone apposition. At 78 weeks, bone ingrowth has occurred throughout all the cavity injected with the composite paste, integrating ceramic particles in its structure. The new formed bone was tightly bonded to the ceramic particles (figure 1), confirming the osteocoalescence of the Diphasic Calcium Phosphate [4]. Even particulate, BCP ceramic could act as scaffold for cell maturation and bone formation, just as what was reported for HA particles [2, 5]. 2. Two kinds of reactions have been observed around the material grains, in contact to medullary cavities. First case, they were under high degradation process, either cell mediated or solution mediated. Solubilized regions on the ceramic grains were identified by backscattered electron microscopy as a less dense zone located on the material surface surrounding a dense core of ceramic (figure 2). TEM evaluation corroborated that the core of the grain was more dense because of the presence of precipitated microcrystals in the intercrystalline spaces, whereas no precipitation was observed at the periphery (figure 3). Through selective area electron diffraction, it could be demonstrated that the microprecipitates consisted of apatitic crystals (figure 4). Two hypothesis could be proposed to explain the absence of microprecipitations at the periphery of the ceramic grain. Based on the assessment of Daculsi et al. [6], that a dissolution/precipitation process occurred in bioactive BCP ceramic prior to bone apposition, these precipitates should be present through all the grain and thus a peripheral dissolution would have occurred, the grain periphery being more exposed to a cleaning by physiological fluids. Another hypothesis could be that a mineralisation process started in the intercrystalline spaces of the core of the ceramic grains and proceeded peripherally. Moreover electron diffiaction patterns identified most of the crystals of the ceramic grain (90%) as HA, suggesting that most of the P-TCP crystals would have been already dissolved or
Figure 5 : TEM picture of a macrophagic cell, phagocytizing ceramic crystals (cer)
Figure 6 : Electron diffraction pattern on a phagocytized ceramic grain, denoting its HA structure
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maturated into apatite crystals after 78 weeks implantation. This would be consistent with the literature about in vivo behaviour of (J-TCP [7, 8]. Table 1 shows the weight Ca/P ratios, calculated on selective areas (grain core, periphery) by electron microprobe analysis. The mean of ten measurements on individual randomly selected crystals revealed that the Ca/P ratio tended to decrease in the periphery of the grain (compared to the Ca/P ratio obtained on BCP powder before implantation), whereas it showed a slight increase in the core. The low value of the Ca/P at the grain periphery resulted actually fi-om the mean of a few individual Ca/P ratios around 1,60, whereas the others Ca/P were equal to 2,00 (just as in the core of the grain). This would indicate calcium deficient apatite crystals at the surface of the grain and would be consistent with the surface studies made on Ca-P ceramics in aqueous environment [9]. The slight increase of the Ca/P in the core of the grain corroborated that the ceramic grain was composed essentially of apatite crystals at 78 weeks implantation. Second case, some loose grains were surrounded by plump osteoblast cells or even covered by an osteoid layer. They corresponded thus to the centre of a calcified nodule, template for bone deposition. The both phenomena observed confirmed that dissolution and bone apposition processes are inter-dependent, many authors suggesting that dissolution took place before a mineralised bone matrix was laid down [6, 10]. 3. Finally, active degradation of the BCP granules also occurred through a phagocytic process, (jiant cells and mononucleated macrophages presenting intracellular vacuoles containing numerous phagocytized synthetic ceramic crystals, were observed in the bone marrow (figure 5). All electron patterns obtained on these phagocytized crystals indicated HA structure (figure 6). ACKNOWLEDGEMENTS Tha authors thank P. Weiss to have taken part in the preparation of the injectable biomaterial and S. Couillaud and M. Cottrel for their helpftil technical assistance. REFERENCES 1. Goodman SB., Aspenberg P., Wang JS., Acta Orthop.Scand, 1993, 64, 627. 2. Wang JS., Goodman S. and Aspenberg P., Clin. Orthop.Relat.Res., 1994, 304, 272-279. 3. Dupraz A., Delecrin J., Moreau A., Pilet P., Passuti N., "Long term bone response to particulate injectable ceramic", submitted for publication. 4. Nery EB., LeGeros RZ., Lynch KL. and Kalbfleisch J., J. Periodont.,1992, 63, 729-735. 5. Zaner DJ., and Yukna RA., J. Periodont.,1984, 55, 406-409. 6. Daculsi G., LeGeros RZ., Heughebaert M. and Barbieux, Calcif. TissueInt., 1990, 46, 20-27. 7. Daculsi G, LeGeros RZ, Nery E, Lynch K and Kerebel B, J. Biomed Mater. Res. 1989, 23, 883-894. 8. Klein CPAT., Driessen AA., de Groot K., van den Hoof A., J. Biomed Mater. Res. , 1983, 17, 769-784. 9. Amrah-Bouali S., Rey C , Lebugle A. and Bernache D., Biomaterials,1994, 15, 269-272. lO.Hashimoto-Uoshima M., Ishikawa I., Kinoshita A., Weng HT. and Oda S., Int. J. Periodontics& RestorativeDentistry,1995, 15, 205-213.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BIOACTIV E GLASS - AND GLASS-CERAMI C COMPOSITE S AND COATING S M.Ferraris^ E. V e m e *\ A. Ravaglioli^, A. Krajewski^, L. Pa^acchini^ J. Vogel^ G. Carl^, C . Jana"*
^Materials Science and Chemical Engineering Dept., Polytechnic of Torino C.so Duca degli Abruzzi 24 -110129 Torino - Italy. ^CNR - IRTEC, Via Granarolo 64 - 48018 Faenza - Italy. ^Joint Research Centre, Institute for Advanced Materials, Ispra (VA), Italy. "^Friedrich Schiller University Jena/Otto Schott Institute, Jena, Germany. ABSTRAC T The preparation of several biocomposites is described and three methods are proposed: pressureless sintering of bulk biocomposites, vacuum plasma spray of biocomposite powders and in situvacuum plasma spray biocomposite deposition on Ti-6Al-4Valloy. The biocomposites are tougher than the correspondent matrices and still showed the same in vitrobioactivity of their matrices. KEYWORDS : biocomposites, sintering, vacuum plasma spray INTRODUCTIO N Several bioactive glasses and glass-ceramics have been studied in the last five years as matrices for the preparation of titanium particle reinforced bioactive composites: the aim of the whole project was the toughening of bioactive glasses and glass ceramics to widen their application field as bulk materials for orthopaedic applications. The well known bioactivity of silicate, and phosphate glasses as SCB [1], AP40 [2], RKKP [3], Bioverit I fiand Bioverit IIIfi [4] was combined with the toughening properties of titanium particles; titanium is known to be one of the most biocompatible metal for prosthetic devices. The obtained bioactive composites still showed the same bioactivity of the glass or glass-ceramic matrix, but they are tougher than the pure matrices. The preparation of bioactive composites was carefully studied, always taking into account that the bioactivity of the glass or glass-ceramic matrix is strictly related to its starting composition, which should not change during the composite preparation. Bioactive composites were prepared by a pressureless viscous flow sintering method to obtain bulk composites, or by vacuum plasma spray (VPS) to obtain composite coatings; in the latter case, two methods were compared: the VPS of presintered powdered composites or the preparation of in situcomposites by spraying mixtures of titanium particles plus powdered glass matrices. Biocomposite bulks and coatings gave interesting results in terms of in vitro bioactivity and mechanical strength, compared to the pure matrices and to hydroxyapatite coatings, respectively. Moreover, the proposed preparation methods are very simple and can be easily modified to obtain graded bioactive composites for the design of specific prosthetic devices. MATERIAL S AND METHOD S The composition of the bioactive glasses used to prepare biocomposites are shown in Table 1. The glasses and the composites were prepared by melting reagents in a platinum crucible; they were then powdered and sieved; the powders were mixed with 15 % volume of Ti particles (Plasma Technik 99,99% purity) to prepare composites by a "pressureless sintering method, as described 195
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Table 1. Composition (wt%) of the biocomposite s (each one with 15 % volume Ti particles) biocomposite s SCB T TSCB T
Si02 50.2 49.1 44.3 44.3 30.5
CaO Na20 Ca3(P04)2 46.9 47.9 24.5 4.6 18.6 APT 24.5 4.6 RKKP T 18.6 14.4 2.3 BI T 14.0 15.0 B3T * contains 0.1 % LaiOa and 1.0 % Td^Os
P2O5
AI2O3
11.4 51.0
15.9 10.0
others balance balance balance jalance* balance balance
referenc e
1 1 2 3 4
in a previous work [5]. The sintered composite s were then powdere d up to 50-100 microns and deposite d by vacuum plasma spray (VPS, Plasma Technik AG ) on Ti-6A1-4V substrates . When the sintering of composite s was not successful , a in situVPS procedure was chosen: i.e. the bioactive glass e powder were mixed with 15 % vol of Ti particles directly in the plasma torch. Each biocomposit (bulk or layer) was characterise d by optical and scanning electro n microscopy (SEM , Philips 525 M) , compositiona l analysis (EDS - Philips 9100), Vickers micro-indentation s and X-ray diffraction (XRD , Philips PW 1710).[l,2,5-7 ] The bioactivity of the composite s was investigate d in vitroby soaking three samples of each one in 50 ml of a simulated body fluid (SBF) [1] for 30 days, then by analysing the morphological and compositiona l changes on their surface by SEM , ED S and XRD ; furthermore , on some biocomposites , anunal fibroblastcell growth was observed after 48 hours at 37 C. [1,2,5-7 ] m n 3N \o
Figure 1. SEM micrograph of a polished cross-sectio n of APT biocomposit e prepared by pressureles s sintering.
Figure 2. SEM micrograph of a polished CTOss-section of BI T biocomposit e deposite d on Ti-6A1-4V by in situ VPS .
RESULT S AN D DISCUSSIO N The aun of this work was the preparation of biocomposite s (as bulk or coating), tougher than the parent glass or glass ceramic matrices. Titanium particles were chosen as toughening particles because of the well known titanium biocompatibilit y and its tougher behaviour, compared to glasses. On the other hand, the preparation of new materials made of glass and metal, required a careful attentio n for the preparation process. Any compositiona l modificatio n of the starting bioactive glass due to the preparation of the composite , i.e. to a reactivity with titanium, could lead to a lack or to a decrease of bioactivity .
Bioactive Glass-and Glass-Ceramic Compositesand Coatings: M. Ferraris et al.
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Firstly, a simple and low-cost pressureles s sintering method [5] was chosen to prepare bulk biocomposites : the linear shrinkage of biocomposit e greens was measured vs temperatur e and vs time, to optimise the sintering process. To fully utilise the softenin g property of the glass matrices, the lowest possible sintering temperatur e was chosen, to avoid any crystallisation , which is known to be detrimenta l for obtaining high density samples. Several biocomposite s with a density of 90-98% (compared to the theoretica l one) were sintered by this method: APT (Figure l),RKKPTandB3T . Secondly, these biocomposite s were powdere d and used as "biocomposite powder" to coat by VPS , a Ti-6A1-4V alloy (widely used for biomedical applications , but not bioactive) : the spraying of biocomposit e powders, i.e. glass coated titanium particles, allowed the use of a low temperatur e VP S process because of the softenin g propertie s of the glass. In this way, the Ti particles did not melt and their glass coating only soften, avoiding problems like, e.g., compositiona l change of the glass matrix and consequentiy , detrimenta l effect s on the bioactivity . Each biocomposit e layer was analysed by EDS , to verify any compositiona l change of its glass matrix after VPS : in all cases, the glass matrix of the biocomposit e did not show any compositiona l variation from the starting one. In some cases, the linear shrinkage of the composite green was too low to obtain an acceptabl e density by a pressureles s sintering method, in other words, the viscosity of these composite glass matrices at the process temperatur e was too high. This was the case of SCBT , TSCB T and BIT . Another preparation method was chosen to obtain these biocomposites : glass and titanium powders were directly mixed in the plasma torch. It was known that this was not the best way to produce heterogeneou s VPS coatings, because of the different density and plasma frequency absorption of the two powders, but for these cases the method resulted quite effective : figure 2 shows a SEM micrograph of a polished cross-sectio n of in situVPS BI T biocomposite . The titanium particles changed their morphology and looked like filaments, probably due to the high temperatur e of the VPS in this case; anyway, the compositio n of the glass matrix did not seem to differ from the starting one. For each bulk- and VPS-biocomposite , the crack propagation due to Vickers micro-indentatio n (0.5-1 kg) was observed and compared to the behaviour of the pure glass or glass-ceramic matrix, prepared, as referenc e material in the same conditions ; figure 3 shows the typical result found for each biocomposite : the crack started from an edge of the tip and was stopped by a Ti particle. By comparison, the referenc e materials showed smaller tip areas and longer cracks, both signals of a lower toughness of the pure matrices compared to the biocomposites . Indentations were also made at the interface betwee n the VPS biocomposit e coatings and the Ti6A1-4V substrate: the coatings were not detache d from the substrate and the cracks propagated in the coating, being deviate d by Ti particles or filaments, in some cases [1], the shear strength of the interface was measured and compared to that of VPS hydroxyapatit e coatings on Ti-6A1-4V: the values are comparable (about 30 MPa). One of the most interestin g propertie s of the prepared bioccMnposite s (bulk or coating) is their bioactivity: if the matrix compositio n does not change during the composite preparation, the resulting material is equally bioactive . In fact, in vitrobioactivity tests on the composite s (by soaking them in SBF, (see refs. 1,2,6,7 ) showed the formation of a calcium and phosphorus rich silica gel layer after a period ranging from few hours to 30 days. It is well known that the formation of such layer is one fundamenta l step of the complex mechanism of bonding to soft and hard tissues, typical of bioactive materials [8]. The two glass-ceramic matrix biocomposite s labelled as BI T and B3T, showed different behaviour in terms of bioactivity , compared to the others, and it was impossible to detect any surface modificatio n after soaking them in SBF, due to the presence of alumina in these compositions ; therefore , their bioactivity was studied by animal
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fibroblast cell growth. Figure 4 shows a completely covered surface of B3T after 48 hours at 37 C in the culture: the cells formed a thick "carpet" which totally hidden the composite surface (compare to Figure 1).
Figure 3. Optical micrograph of a Vickers indented B3T bulk composite: the crack is stopped by a Ti particle
Figure 4. B3T bulk composite after 48 hours at 37 C in the culture: the fibroblast cells form a thick "carpet" which totally hidden the composite morphology (compare to Figure 1).
SUMMAR Y Several biocomposites were prepared by pressureless sintering or by VPS; biocomposite powders or mixed glass and titanium powders were used fw the VPS process. The biocomposites were tougher than the correspondent matrices and still showed the same in vitrobioactivity of their matrices. These materials could be proposed as small bone substitutes or tough bioactive coatings for titanium alloys. REFERENCES 1. M.Ferraris, P.Rabajoli, L.Paracchini,F.Brossa "Vacuum Plasma Spray Deposition of Titanium Particle/Glass-Ceramic Matrix Biocomposites" Journal of American Ceramic Society 79,6 (1996) 1515-20. 2. E. Verne’, M. Ferraris, A. Ventrella, A. Krajewski. A. Ravaglioli, "Bioactive glass-matrix composite coatings on titanium alloy", "Ceramic in Oral Surgery" Edited by A. Ravaglioli and A. Krajewski, Gruppo Editoriale Faenza Editrice 1995, p.231 3. A. Krajewski, R. Malavolti and A. Piancastelli, "Albumin Adhesion on some Biological and Non-biological Glasses and Connection with their Z-Potentials", Biomaterials, 17(1996)53-60 4. W. Hoeland and W. Vogel, "Machineable and Phosphate Glass-Ceramics", in "An introduction to Bioceramics", vol.1, cap.8, ed. by L.L.Hench and J. Wilson, World Scientific, 1993, p.l25. 5. M.Ferraris, E.Veme* "Viscous Phase Sintering of Particle-Reinforced Glass Matrix Composites" Journal of the European Ceramic Society, 16 (1996) p.421 6. M.Ferraris, P.Rabajoli, L.Paracchini,F.Brossa "Vacuum Plasma Spray Deposiiion of Titanium Particle/Glass-Ceramic Matrix Biocomposites" Journal of American Cerainic Society 79,6 (1996) 1515-20. 7. E.Verne’, M.Ferraris "Characterisation of Vacuum Plasma Sprayed Biocomposites", in "Bioceramics", Edited by A.Ravaglioli, vol. 8, p.l47 (1995) 8. "An Introduction to Bioceramics", Vol.1, ed. by L. L. Hench and J. Wilson, World Scientific Pub.,1993.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MECHANICA L CHARACTERISATIO
N OF BIOACTTV E COATING S ON ZIRCONI A
E. Vem6\ M. FerrarisS C. Moisescu^, A. Ravaglioli^ A. Krajewski^ ^Materials Science and Chemical Engineering Dept., Polytechnic of Torino - C.so Duca degli Abruzzi 24 - 10129 Torino - Italy. ^CNR - IRTEC, Via Granarolo 64 - 48018 Faenza - Italy. KEYWORDS : Bioactive glasses, glass and glass-ceramic coatings, mechanical tests. ABSTRAC T In order to join the mechanical properties of a high-strength inert ceramic, (yttria-stabilised zirconia, 2i02-3%Y203) with the properties of glasses, a bioactive calcium-phosphate glass and glass-ceramic (AP40), was used to coat zirconia substrates (named "bioactive zirconia" in this paper). The coatings were prepared by a simple and low cost firing method; their microstructure was characterised by q)tical and Scanning Electron Miaoscopy (SEM) and compositional analysis (EDS). The adhesion of the coatings to zirconia was tested by three different methods: Vickers indentations at the coating/zirconia interface, shear tests on sandwich structures, and four point bending tests on butt-joints, giving encouraging results. INTRODUCTIO N Zirconia is one of the newest biocompatible ceramic, proposed for the human body parts substitutions [1]. The applications of zirconia are generally focused on the field of bone surgery, for the realisation of prostheses that need good mechanical properties. When implanted, it shows a morphological fixation with the surrounding tissues, without any chemical or biological bonding; for that reason it could not be considered a bioactive material, but a biocompatible "inert ceramic". Several bioactive glasses and glass-ceramics have been developed fOT biomedical s^plications due to their ability of forming a strong bond to hard and soft tissues [2]; they found a large employ when the replaced part has not to be load bearing [3]. In order to combine the mechanical properties of a high-strength material, as zirconia, with the peculiar properties of bioactive materials, AP40 [4-6] glass and glass-ceramic thin coatings on zirconia have been successfully prepared by a simple and low cost firing method, after an accurate thermal treatment optimisation, as described in previous works [7,8]. In this work, we evaluated the adhesion strength between the zirconia substrate and the glass or glass-ceramic coatings by using different mechanical test methods [9,10] in order to propose the bioactive zirconia for clinical applications. MATERIAL S AND METHOD S Zirconia substrates were obtained by sintering 3% Y2O3 - stabilised zirconium oxide powders. Before ^plying the glass powders, the substrate was ultrasonically cleaned in acetone. Two kinds of substrates, prepared at Irtec- Faenza (Italy) were used: as sintered zirconia and a laser textured surface zirconia. An optimised thermal treatment process was developed to coat zirconia substrates by an amorphous AP40 layer: zirconia substrates were covered by dry glass powders (having a 199
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granulometry up to 100 microns), then heated at temperature s just above the melting temperatur e (1300 C), obtaining 100-30 0 micron thick layers. After the heating treatment , the coatings were simply annealed to obtain an amorphous coating, or thermally treated with a nucleatio n and growth process to obtain a glass-ceramic coating. [7,8] Glass Zirconia ^ ^ ^
Zirconia a) Vickers indentatio n
b) Shear test
Zirconia c) Four point bending test
Fig. 1.: Different mechanica l tests used in this work. The interface adherenc e betwee n the coatings and zirconia was firstly determine d by performing some Vickers indentatio n tests (500 N) at the interface coating/substrat e (fig. l.a). In order to perform the shear tests [9] and the four point bending tests [10], some zirconia substrates were joined by a glass layer, using the same thermal treatmen t develope d to obtain the amorphous coatings^ The size of the zirconia substrate s was of 5x5x2 nun^ and 50x5x5 nmi^, and their shape is described in figure l.b and l.c, respectively . Each sample was characterise d by optical and scanning electro n microscopy (SEM - Philips 525 M ) and compositiona l analysis (EDS ) (Model ED AX 9100, Philips) before and after the mechanica l characterisation . RESULT S AN D DISCUSSIO N After the preparation process, each cross-sectione d sample showed three different layers: the starting zirconia substrate, a "composite" layer, (rf an average thickness of 25 |im, made of zirconia granules surrounded by the glass, and the glassy or glass-ceramic AP40 layer. As already discussed [7,8], this "composite" allows a gradual variation of the thermal and mechanica l properties from the zirconia substrate to the glass coating; ED S analyses performed in different regions of the coatings, demonstrate d that the glass or glass-ceramic layers do not change their starting composition ; this feature is fundamenta l to guarantee the bioactivity of the coated zirconia [7,8]. Vickers indentation s (500 N loads) were made at the interface betwee n the composite layer (zirconia + glass) and the glass coating (fig. 2.a) or the glass-ceramic coating (fig. 3). The cracks propagated mainly into the amorphous coatings (fig. 2.a, b) because the fracture energy of the interface is higher than that of the glass coating. When the cracks propagated through the composite layer, they were immediatel y deviate d by the zirconia granules, without any detachmen t of the coating from the substrate , proving that this is a tough layer, fundamenta l for the mechanica l propertie s of the coated zirconia. Some cracks propagated into the glass-ceramic coatings, but never at the interface betwee n the coating and the composite (fig. 3, see arrows). A tougher behaviour was observed for the glass-ceramic coating than iocthe glassy one, as shown by figure 2and 3: cracks propagated straightly through the glass coating while the glass-ceramic coating only showed the Vickers tip with short aacks at the edges, readily stopped by the crystalline phases.
Mechanical Characterisationof Bioactive Coatings on Zirconia: E. Verneet al.
201
The shear tests were preliminary performed both on the as sintered and textured zirconia/glass/zirconia sandwiches (three samples of each one): the shear strength (x) was of about 62 MPa for the textured samples and about 30 MPa for the as sintered ones. It seems to be evident that a surfacial roughness of the substrate can increase the coating adhesion by a mechanical bonding. Moreover the fracture surfaces morphology showed that the cracks propagated on different planes and not on preferential directions, i.e. the interface substrate/glass, or the glass itself. The four point bending test, preliminary carried out only on as sintered zirconia/glass/zirconia butt-joints, gave a bending strength (a) of about 170 MPa. Also in this case the fracture surface morphology showed a multiplanar cracks propagation and it is expected to obtain higher bending strength results using the textured zirconia substrates. SUMMAR Y Coatings of bioactive glass and glass-ceramic AP40 on TsOi substrates having two different surface roughness, were prepared by a low cost firing method: the glass and glassceramic coatings did not change their starting composition after the preparation process; in this way "bioactive zirconia" was obtained. The mechanical characterisations, carried out by means of three different methods, showed in each case a strong adherence of the coatings to the substrates, particularly for the textured zirconia samples. The same mechanical characterisation is in progress on glass-ceramic coated zirconia samples.
Figure 2. Cross section of an AP40 glass coating/zirconia interface, after Vickers indentation tests (500 N). a): general view; b) particular of the crack path. ACKNOWLEDGEMEN T The authors are mdebted to the Fiat Research Centre (To, Italy) fcM* SEM-EDS analyses, and to the "Associazione per lo Sviluppo Scientifico e Tecnologico del Piemonte" (ASP) for financial supports.
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Figure 3. Cross section of an AP40 glass-ceramic coating on zirconia substrate, after Vickers indentatio n tests (500 N). REFERENCE S 1. S.F.Hulbert, "The Use of Alumina and Zirconia in Surgical Implants", in "A M Introduction to Bioceramics’\ Advances Series in Ceramics-vol.l , p.25, World Scientific Publishing Co. Pteitd. (1993) edited by L.L.Hench and J.Wilson. 2. W.Cao, L.L.Hench,"Bioactive Materials", Ceramics Internationals, 22 (1996) 493-507 . 3. J.Wilson, A.Yli-Urpo and H.Risto-Pekka, "Bioactive glasses: Clinical Applications", in "A n Introduction to Bioceramics’\ Advances Series in Ceramics-vol.l , pag. 63, World Scientific Publishing Co. Pte.Ltd. (1993) edited by L.L.Hench and J.Wilson. 4. F.Wihsmann, G. Berger, G. Bochynek, V. Thime, H. Hoftnann, S. Kohler, K. Retemeyer , "Bioaktive Implantate auf der Basis von Vitrokerammaterialien" , Wiss.ZtschnFSUJena,2/3 (1983), 553-569 . 5. G. Berger, R. Gildenhaar, "Long - Term Stable Bioactive Glass Ceramic as Implant Material Ten years of Clynical Experience", Fourth World Biomaterials Congress, April 24-28, 1992, Federal Republic of Germany, Berlin, p.33. 6. A. Ravaglioli, A. Krajewski, A. Piancastelli , G. Berger, K.Adam, R. Gildenhaar, "Th e Influence of Alumina Porosity in Glass Ceramic Adhesion" Interceram, vol. 41, n.2, 1992 p.69. 7. M.Ferraris, E. Vem6, P. Appendino C. Moisescu A. Krajewski, A. Ravaglioli, A. Piancastelli , "Bioactive Glass and Glass-Ceramic Coatings on Zirconia", submitted to the J.Am.Cer.Soc. 8. M.Ferraris, E. Verne’, C. Moisescu, A. Ravaglioli, "Bioactive Coatings on AI2O3 and Zr02", 3rd Meeting and Seminar on Ceramics, Cells and Tissues, Faenza, May 2 - 4 1996. 9. J.Lin and H. Kato, "Interfacial Structure and Strength of Silicon-Silicon Diffusion Bonding", Mat.Sci,andTecK H (1995) 1035-40 . 10. A.J.Moorhead and H.E. Kim, "Joining of Oxide Ceramics", in ’’ShortCourseon ceramic Joining’\98th Annual Meeting and Exposition Of the American Ceramic Society, Indianapolis, In. (U.S.A) April 1996, Sec.7, p. 511.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
VACUU M PLASM A SPRAYE D TITANIU M AND HYDROX Y APATIT E COATING S ON CARBO N FIBE R REINFORCE D POLYETHERETHERKETON E (PEEK ) S.-W. H a \ A. Gisep\ H. Gruner^, J. Mayer^ E. Wintermantel^ ^ Chair of Biocompatible Materials Science and Engineering, Department of Materials, ETH Zurich, Wagistr. 23, CH-8952 Schlieren ^ Medicoat AG, Gewerbe Nord, CH-5506 Magenwil
ABSTRAC T In this study, carbon fiber reinforced polyethetheretherketone (CF-PEEK) was coated with tita› nium (Ti) and hydroxyapatite (HA) by vacuum plasma spraying (VPS). The coatings were charac› terized using scanning electron microscopy (SEM), ground section analysis, optical profilometry and X-ray diffraction (XRD). The obtained VPS coatings were compact, crack-free and had a rough topography. As main component hydroxyapatite was detected by XRX). Additionally, small amounts of p-tricalcium phosphate were found. Cross section analysis revealed a close contact between sub› strate and VPS coating. The thermoplastic composite substrate showed no voids or cracks. It is therefore concluded that after adequate adaption of the process parameter, the VPS technique has no adverse effects on the substrate material and may be a suitable method for obtaining osteoconductive coatings on carbon fiber reinforced PEEK implants. KEYWORDS Vacuum plasma spraying, carbon fiber reinforced PEEK, titanium, hydroxyapatite INTRODUCTIO N Carbon fiber reinforced PEEK is currently being investigated as implant material for orthopaedic implants [1,2]. According to the literature, very few studies have been performed to deposit calcium phosphate coatings on thermoplastic polymer materials by plasma spraying. Therefore, a systematic approach to apply the VPS process on CF-PEEK was performed in the present investigation. Exist› ing technique, which had been established for metallic substrates [3] was used as a basis for adapt› ing the VPS technique to the new substrate material. The objectives of this study were (a) development of an improved process setup for vacuum plasma spraying Ti and HA on CFPEEK, and (b) topographical and chemical characterization as well as microstructural analysis of the interface between composite substrate and VPS coating. MATERIAL S AND METHOD S CF-PEEK disks (Ensinger GmbH, Germany) with a diameter of 10 mm and a height of 7 mm were sandblasted with alumina, cleaned with ethanol and deionized water and dried in a vacuum oven at 200 C for at least 7 days. VPS was performed in a plasma spraying chamber (Medicoat AG, CH-Magenwil). Ti coating was produced with fine Ti powder having an average grain size of 203
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d5o"" 25 |Lim. Subsequently, coarse Ti powder (d5o = 120 jim) was applied. HA coated specimens were first coated with Ti as described above and subsequently with HA powder (d5o =115 |Lim). HA coatings of three different thicknesses were prepared. After VPS coating, the specimens were gently rinsed in distilled water and dried in air. Coating morphology and cross section analysis was performed using scanning electron micros› copy (SEM, Hitachi S-2500C), Roughness values and profile length ratio values of the VPS-coat› ings were determined with an optical autofocusing profilometer (Laser UBM). X-ray diffraction of the VPS HA coated samples was performed with a PADX diffractometer (Scintag, USA) with CuK(3^-radiation in the Bragg-Brentano mode. The X-ray data was collected in the 20 range of 1555 in steps of 0.02 . RESULTS AND DISCUSSION The Ti coating showed a rough and porous topography built-up by coarse Ti particles (fig. 1 left). The backscattering electron (BSE) image of the Ti coating revealed a dense Ti layer (fig. 1 middle) and it is assumed that the beforehand applied layer of fine Ti particles completely cover the underly› ing composite substrate surface. Fig. 1 (right) shows a HA coating, which displays a with a smoother topography compared to the rough and porous Ti coating. Cross sections of the VPS coat› ings showed a close contact between substrate and Ti coating (fig. 3). The Ti layer is completely covered with HA and the close contact between Ti and HA indicates a good mechanical interlocking at the HA/Ti interface. The substrate showed no formation of voids and cracks indicating that the coating process had no adverse effects on the substrate material. The various roughness values of VPS Ti and HA coatings measured with an optical profilometer (Laser UBM) are shown in table 1. The rough surfaces of the VPS Ti coatings showed average roughness values of R^ = 28.29 – 3.07 |im, Rq = 36.61 – 2.92 |Lim, R^ = 145.35 – 9.88 and R^ax "" 179.21 – 9.30 (fig. 2). With increasing HA coating thickness roughness values decreased to Ra= 12.25 – 2.45 |Lim, Rq = 15.30 – 2.95 |im, for the HA coatings with a thickness of 150 |Lim. Pro› file length ratio L^ significantly increased from a value of 1.08 to 1.45 after coating the sandblasted CF-PEEK substrate with Ti. Increasing the HA coating thickness resulted in a reduction of the L^ values. The 150 L | Lm thick HA coating showed an L^ value of 1.09.
Figure 1
SEM image (left) and corresponding BSE image (middle) of a VPS-Ti coating and SEM image of a VPS-HA coating (right).
Vaccuum Plasma SprayedTitanium andHydroxyapatiteCoatings: S.-W. Ha et al. 205 Table 1
Roughness values Ra, Rq, R^ and R^^^x ^"^ profile length ratio L^. of CF-PEEK substrates before and after coating with Ti and HA using VPS (n=5).
Specimen
RaU^m]
Rq[/im]
Rzl/im]
Rmax [/^m]
Lr
CF-PEEK
4.64 – 0.69
6.68 – 1.06
62.71 – 4.49
53.65 – 9.02
1.08 – 0.01
Ti
28.29 – 3.07
36.61 – 2.92
145.35 – 9.88
179.21 – 9.30
1.45 – 0.05
HASOjL/ m
20.18 – 2.31
26.30 – 2.65
99.94 – 8.24
147.19 – 17.73
1.16 – 0.02
HA100jL/ m
13.25 – 2.95
15.24 – 1.19
62.71 – 13.26
80.08 – 13.41
1.12 – 0.01
HA150jL/ m
12.25 – 2.45
15.30 – 2.95
62.71 – 7.80
79.08 – 16.30
1.09 – 0.01
H Ra 40.00
X
J
35.00
] E ] ] g 20.00 §* 15.00 ] O ] ^ 10.00 (0 (0
5.00 0.00
im xii
!
11 1 M m
, -, 30.00
Rq
-
200.00
^
TI
1.
^ T .^ i
mt\
;
HA 50
11 HA 100
HA 150
Ref
Ti
HA 50 urn
HAIOO^m
HA 150 urn
Figure 2
Roughness values R^, Rq, R^ and R^ax of VPS-Ti and HA coatings measured with an optical autofocusing profilometer (Laser UBM; HA 50 = 50 )im, HA 100 = 100 |Lim, HA 150 = 150 |Lim).
Figure 3
SEM image (left) and BSE image (right) of a cross section of VPS-HA coating on CF-PEEK.
The chemical composition of the as-received sample was investigated by XRD. The spectrum showed hydroxyapatite (PDF # 9-432) as main component and a small p-tricalcium phosphate (p-TCP) peak at 20 = 31.16 (fig. 4). No other phases were detected. The obtained spectrum is com› parable with those of the VPS HA coatings sprayed on titanium substrates [4]. It is therefore con› cluded that adapting the VPS process for CF-PEEK as substrate material did not affect the chemical composition of the HA-coating.
206 Bioceramics Volume10 14000 12000 10000
Figure 4
XRD-spectrum of the vacuum plasma sprayed HA-coating showing HA (x) as main component and a small fraction of ^TCR
CONCLUSIO N With a VPS coating setup optimized for CF-PEEK substrates, compact and crack-free Ti and HA coatings were obtained. The used VPS process paramaters lead to the desired coatings on the com› posite material. Cross section analysis showed that the Ti layer completely covered the CF-PEEK substrates, which is assumed to be an important factor for the biological performance of the compos› ite material. The achieved coating thickness of the Ti layer varied between 10 and 150 |Lim. With this underlying rough Ti layer, HA coatings of various roughness and profile length ratio values were obtained, depending on coating thickness. No physical substrate deterioration and a good interlock› ing between Ti intermediate layer and CF-PEEK substrate was observed. It is assumed, that the observed interlocking will result in suitable mechanical adhesion strength. This assumption has to be confirmed in further investigations. From the results obtained in this study it is concluded that VPS is a suitable method for manufacturing HA coatings on carbon fiber reinforced PEEK implants.
REFERENCES [1] [2] [3] [4]
Widmer M., et al., OberflachenWerkstoffe, 5, 1995, 34-37. Heiner A.D., et al., Biomaterials,17, 1996, 2211-2217. Gruner H., Pulvermetallurgiein Wissenschaftund Praxis, 8, 1992, 82-101. Ha S.-W., et al., Journal of theAmericanCeramicSociety,1997, submitted.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
LO W TEMPERATUR E CRYSTALLIZATIO FILM S IN AN AUTOCLAV E
N OF HYDROX Y APATIT E SPUTTERE D
J. Hamagami, K. Nakamura, Y. Sekine, K. Yamashita, and T. Umegaki Department of Industrial Chemistry, Faculty of Engineering, Tokyo Metropolitan University, Minami-Osawa, Hachioji, Tokyo 192-03, Japan ABSTRAC T Coatings of crystalline hydroxyapatite (Caio(P04)6(OH)2, HAp) thin fihns with the thickness of 2 |Lim on a-Al203 ceramics, Inconel 600 alloy or Ti metal plates were carried out by annealing sputtered films at low temperature (-HO^C) under high water vapor pressure in an autoclave. The sputtering was carried out m an argon atmosphere using the calcium phosphate glass target with a Ca/P molar ratio of 0.6, much lower than that of the stoichiometric HAp value of 1.67. Assputtered film was annealed in an autoclave at temperature of 130 to 180 C for 5 to 146 hr under the saturated water vapor pressure. HAp phase was crystallized at a temperature as low as 140 C. The present annealing method was also proved effective for the bonelike apatite coating of metal substrate. [KEYWORDS ] Hydroxyapatite coatmg, r.f magnetron sputtering. Low temperature crystallization. Calcium phosphate glass target. Titanium metal INTRODUCTIO N The hydroxyapatite fihns coated onto ceramics and metals with excellent mechanical properties have currently been developed for the implant applications[l-7]. The main advantage of the coating by sputtering is a high adhesion strength between films and substrates. The biocompatibility of sputtered HAp layers onto titanium has been proved by Jansen et al.[4]. We have recently succeeded in the coating of single phase HAp onto alumina and zirconia ceramics using the glass target of the calcium phosphates with much lower Ca/P ratios than the stoichiometric value of HAp (=1.67) [1]. In general, the crystallization of the amorphous fihns are performed at a temperature above 600 C [1,7]. However, high temperature annealing gives rise to some cracks on the coating, oxidation of metal or alloy substrates, and partially peeling of films from substrate because of the high temperature annealing process. The thermal shock of the films and substrates also become more serious with increasing the annealing temperature. This paper mainly reports the development of the coating technique of single phase HAp at low annealing temperature through sputtermg. Good biocompatibility of the postannealed coatings will also be by in vitroexperiments of the bone-like crystal growth m a simulated body fluid. MATERIAL S AND METHOD S The precursor fihns were prepared by the r. f magnetron sputtering apparatus (SP87-005, Vacuum Metallurgy Co.). The targets for sputtering were calcium phosphate glasses with 0.6 of a Ca/P molar ratio much smaller than 1.67 for stoichiometric HAp. These glass targets of calcium phosphate was made by melt quenching method according to previous work[l]. The substrates 207
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used were alumina ceramic, Inconei 600 alloy, and titanium metal plates. Sputtering was carried out in a 0.67 Pa argon atmosphere. The applied r.f. input power was fixed at 7.6 W/cm^ operating at 13.56 MHz. The sputtered fihns had the thickness of 2 |Lim. These as-sputtered films were crystallized by postannealing using an autoclave (Figure 1). The annealing was carried out at 130-180 C for 5 to 146 hr under the saturated water vapor pressure. In order to estimate of the biocompatibility, the postannealed films were examined as follows; the specimens were soaked for several days at 37.5 C in a simulated body fluid (SBF) with pH=7.25 [8]. These as-sputtered, postannealed, and soaked specimens were characterized by X-ray diffraction (XRD), Fourier transform infrared (FT-IR) analyses and scanning electron microscopy (SEM) with energy dispersive x-ray (EDX) analysis. RESULTS AND DISCUSSION Since the sputtering was carried out under no intentional heating of the substrate, all assputtered films were confirmed to be almost completely in amorphous state. Figure 2 shows the XRD pattern of the film onto the alumina substrate annealed at 140 C for 20 hr under high water vapor pressure. Since the peaks observed in the specimen were assigned to those of HAp Ca5(P04)3(OH) (JCPDS #9-432) and a-A^Os (JCPDS #46-1212), the fibn prepared by annealing at a temperature as low as 140 C was confirmed as single phase HAp. Another related calcium phosphates such as tetracalcium phosphate were not be detected in the annealed films. The crystallinity of the postannealed films depends on temperature for annealing. The shnilar XRD results were also obtained in the case of Inconei 600 alloy and titanium metal substrates.
r
Teflon packing
[Test tube containin g distilled water \ !
T
H-Stainles s Steel Teflon liner
Specime n
Figure 1. Schematic illustration of an autoclave for annealing. autoclave with a test tube containing distilled water.
A sputtered specimen is set in an
Low TemperatureCrystallizationof HAP SputteredFilms in an Autoclave:J-I. Hamagami et al.
209
3 (0
0)
c
0)
30 40 20 (Cu-KJ / degre e
50
60
Figure 2. X-ray diffraction pattern of the film on the a-Al203 substrate annealed at 140 C for 20 hr under high water vapor pressure. In order to study the local structure concerning the ion groups of P04^" and OH", multiple internal reflection IR spectroscopic analysis was carried out. Figure 3 shows the IR absorption spectrum of the specimen onto Inconel 600 postannealed at 140 C for 20 hr under high water vapor pressure m the wavenumber region 400 to 4000 cm’\ Inconel 600 alloy used as the substrate showed no absorption peak in this wavenumber range. The IR absorption peaks at 962, 1035, and 568 cm"^ were assigned to the ion groups of tetragonal P04^’ in HAp structure. On the other hand, the absorption peaks at 3550 and 630 cm’^ due to the OH" ions m HAp were too weak to be confirmed. The structure of the present HAp films was assumed to be nonstoichiometric oxyhydroxyapatite, possibly expressed as Caio(P04)6((OH)2.2xOxnx), where D represents the vacancy at the OH lattice site. The biocompatibility of the postannealed specimen was carried out in vitro experiments. Figure 4 shows the SEM photographs of the surface of HAp fihns deposited onto titanium metal before (a) and after (b) soaking the fihn in SBF for 6 days.
0)
o c
1(/) c
(0
4000
3500
3000 2500 2000 1500 Wavenumber / cm-^
1000
500
Figure 3. Infrared absorption spectrum of the specimen onto Inconel 600 postannealed at 140 C for 20 hr under high water vapor pressure.
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(b) After soaking in SBF for 6 days i
:.;i’:’"^^^^^^^^^
Figure 4.
v^.^ ... v^
’
^
Scanning electron micrographs of the surface of HAp fibn deposited on titanium metal before (a) and after (b) soaking the fihn in simulated body fluid after 6 days.
The postannealed specimens did not have any cracks on the surface with SEM observation (Figure 4 (a)). As seen in Figure 4 (b), the crystal growth on the surface of the specimens is observed. The grown layers on the surface were confirmed to be bonelike apatite by XRD and JR. This result is important from the practical viewpoint that the mechanically excellent ceramics and metal can be provided with bioactivity by this low temperature annealing method. CONCLUSION S The polycrystalline hydroxyapatite coatings onto a-Al203 ceramics, Inconel 600 alloy, and titanium metal substrates were prepared by an r.f magnetron sputtering followed by low temperature annealing in an autoclave. Formation of single phase HAp thin fihns were obtained by postannealing at temperature as low as 140 C for 20 h under high H2O vapor pressure. The postannealed fihns were confirmed to be oxyhydroxyapatite by XRD and FT-IR measurements. ACKNOWLEDGMEN T The present work was supported by Grant-in-Aid for Scientific Research, the Ministry of Education, Science, Sports and Culture of Japan. REFERENCE S 1. Yamashita, K., Arashi, T., Kitagaki, K., Yamada, S., Umegaki, T., and Ogawa, K., J. Am. Ceram.Soc. 1994, 77, 2401-2407. 2. Wolke, J. G. C , Van Dijk, K., Schaeken, H. G., De Groot, K., and Jansen J. A., J. Biomed Mat. Res 1994, 28, 1477-1484. 3. Van Dijk, K., Schaeken, H. G., Wolke, J. G. C , Maree C. H. M., Habraken, F. H. P. M., Verhoeven, J., and Jansen, J. A., J. Biomed.Mat. Res. 1995, 29, 269-276. 4. Hushoff, J. E. G., Van Dijk, K., van der Waerden, J. P. C. M., Wolke, J. G. C , Kalk, W., and Jansen, J. A., J. Biomed Mat. Res. 1995, 29, 967-975. 5. Yamashita, K., Yagi, T., and Umegaki, T., J. Am. Ceram.Soc. 1996, 79, 3313-3316. 6. Yamashita, K., Yagi, T., Hamagami, J., and Umegaki, T., Bioceramics1996, 9, 337-340. 7. Van Dijk, K., Schaeken, H. G., Wolke, J. G. C , and Jansen J. A., Biomaterials1996, 17, 405-410. 8. Tanahashi, M., Kokubo, T., Nakamura, T., and Yamamoto, T., J. Am.,Ceram.Soc. 1995, 78, 1049-1053.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
FABRICATIO N OF IN-CERA M COR E BY SHEE T FORMIN G PROCES S Dae-Joon Kim\ Myung-Hyun Lee^ and Chang-Eun Kim^ ^Ceramics Division, Korea Institute of Science and Technology, 39-1 Hawolgok-dong, Seoul 136791, Korea ^Department of Ceramic Engineering, Yonsei University, Seoul 120-794, Korea
ABSTRAC T A novel fabrication process of In-Ceram crown core for all-ceramic dental restorations has been developed utilizing ceramic tapes prepared by the doctor-blade method. Optimum tape properties were achieved by mixing alumina powder with polymers in the weight ratio of 40 to 9. The polymers consisted of polyvinyl butyral as a binder and dibutyl phthalate as a plasticizer in the weight ratio of 4 to 5. The core was formed by wrapping the tape around a tooth duplicate plaster die and by isostatically pressing at 10 MPa in a silicon oil bath at 80 C. After sintering and glass infiltration, accuracy of fit of the ceramic restoration was comparable to that produced from a slipcast alumina powder in the In-Ceram crown fabrication process. KEYWORDS : dental ceramic, all-ceramic crown, ceramic sheet INTRODUCTIO N Materials for dental restoration should meet three essential criteria, that is, strength, esthetics, and biocompatibility. Although traditional porcelain-fused-to-metal restorations are widely practiced, the quest for esthetic and biocompatible restorative materials has led to developments of various all-ceramic systems for fixed dental restorations [1,2]. Among them In-Ceram is relatively recent all-ceramic restoration with good esthetics, high strength, and accurate fit [3], which are derived from a sintering of slip-cast alumina at relatively low temperature of 1120 C and a subsequent infiltration of the porous structure with molten glass, whose main components are Si02, B2O3, AI2O3, La203, and CaO [4]. Despite the promising properties, the forming process of the In-Ceram crown core, which involves a slip preparation, building up the slip on a plaster model by brushing, drying, and a control of the core thickness, is somewhat troublesome and timeconsuming to be executed in dental laboratories. In the present study the core was shaped using ceramic tape and tape properties required for the forming process were investigated. MATERIAL S AND METHOD S The ceramic sheets were prepared by a tape casting using the doctor-blade technique where a ceramic suspension passes beneath a blade to form a layer of controlled thickness [5]. The slurry was composed of alumina powder (AL-32, Sumitomo), solvent (MEK 66 wt% + Etoh 34 wt%), dispersant (Solsperse 24000, ICI), binder (polyvinyl butyral), and plasticizer (dibutyl phthalate). 211
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Not formable
74 0.76 0.78 0.80 0.82 0.84 0.86 c/(c+p) Fig. 1. Formable composition region at room temperature as depicted by the weight ratios of ceramic powder/(ceramic powder+polymers) and binder/(binder+plasticizer). The amount of the dispersant was determined to be 0.5 wt% with respect to the powder from measurements of viscosity versus dispersant concentration in the alumina suspensions. The suspensions were prepared by ball-milling for 4 h. Subsequently, binder and plasticizer were added, and the batches were ball-milled for an additional 20 h. The contents of alumina powder, binder, and plasticizer were varied to optimize the flexibility of tape and the ceramic powder loading. The batches were de-aired and the tapes were cast at a thickness of 0.5 mm on a Mylar film. Mechanical properties of the sheets were characterized by a tensile test of dog-bone-shaped specimens, prepared according to ASTM D 638-95 [6], at the loading rate of 50 mm/min. The tape was wrapped around a tooth duplicate plaster die to shape the crovm coping. The adherence of tape to a contour of the die and at a folded part was improved by an isostatic pressing of 10 MPa at 80 C. The coping was sintered on the duplicate plaster die with the following heating schedule: 25 - 600 C at rate of l C/min, isothermal hold for 1 h; 600 - 1120 C, 3 C/min, isothermal hold for 2 h. The resuhed porous structure was strengthened by mfiltration of the glass powder following the preparation procedure of In-Ceram crown core [3]. RESULTS AND DISCUSSION The requirements of ceramic sheets for forming the In-Ceram crown core are a high flexibility for an easy shaping and a high ceramic powder loading for maintaining an integrity of the framework during sintering. The flexibility can be obtained by increasing the polymer content, especially plasticizer, but at the same time this causes to a low powder loading in tapes. In an effort to optimize the flexibility and powder loading, a composition region, in which the core can be shaped from tapes, was evaluated by means of folding tapes having various contents of the alumina powder, binder, and plasticizer. Tapes that were not suffered a rupture at the folded area were classified as formable ones and their composition region was depicted in Fig. 1. For the tapes prepared with the formable compositions in Fig. 1, the influence of powder content on strength and strain to failure was investigated and the results are shown in Fig. 2, where the weight ratio of binder to binder plus plasticizer (b/(b+pl)) was fixed to 0.444. The tensile strength and the strain to failure increased and decreased, respectively, with increasing the
Fabrication of In-Ceram Core by Sheet Forming Process: D-J. Kim et al.
,4
213
A
.
% B g
c/(c+p) Fig. 2. Effect of ceramic/(ceramic+polymers) weight ratio on tensile strength and strain to failure of tapes containing a fixed weight ratio of binder to plasticizer at 0.444.
b/(b+pl) Fig. 3. Effect of binder/(binder+plasticizer) weight ratio on tensile strength and strain to failure of tapes containing a fixed weight ratio of ceramic powder to polymers at 0.816.
powder loading in the mixture of the alumina powder and the polymers consisted of the binder and the plasticizer. As the weight ratio of the powder to polymers (c/(c+p)) was higher than 0.82, the strength decreased abruptly and the tapes were fractured in a brittle fashion. Consequently, the tapes could not be used for forming the framework as indicated in Fig. 1. For the tape having the limiting composition, the strain to failure was determined to be about 0.6 indicating that the ceramic tapes should posses the strain to failure higher than 0.6 for a successful shaping of the allceramic crown core. The influence of relative content of the binder and the plasticizer, as designated by the weight ratio of b/(b+pl), on the mechanical properties of alumina sheets is exhibited in Fig. 3, where the ratio of c/(c+p) was fixed at 0.816. Compared with the effect of powder loading in Fig. 2, the change in the strength is more pronounced but the variation in the strain to failure is less significant. The strain to failure dropped below 0.6 as b/(b+pl) increased above 0.5 in Fig. 3. The tapes having the strain to failure smaller than 0.6 lost an adhesiveness so that a lamination of tape at an overlapping area became difficult during forming the crown core. Among the sheets fractured at the strain to failure higher than 0.6, the core structure, shaped with the tapes having b/(b+pl) smaller than 0.4, could not sustain when the polymers were decomposed during the sintering process. Accordingly, the optimum composition for the fabrication of all-ceramic crown core resides in 0.81 < c/(c+p) < 0.82 and 0.4 < b/(b+pl) < 0.5 in the composition map in Fig. 1. It is noteworthy that the strain to failure and the tensile strength are strongly influenced by the ceramic powder content and the weight ratio of binder to plasticizer, respectively, as indicated in Figs. 2 and 3. Fig. 4(a) demonstrates the tapes prepared from the slurry containing the alumina powder and the mixture of the binder and the plasticizer, which was mixed in the ratio of 0.444, in the weight ratio of 0.816. The tape, wrapped around the duplicate plaster die and isostatically pressed, is exhibited in Fig. 4(b). Fig. 4(c) shows sintered ceramic core after trimming the shaped framework. The porous core was strengthened by infiltration of the molten lanthanum
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(a)
(b)
(c )
(d)
(e)
Fig. 4. All-Ceramic dental restoration fabricated using ceramic tape: (a) alumina tapes; (b) crown core shaped using tape; (c) sintered core; (d) glass infiltrated alumina core; (e) all-ceramic crown. borosilicate glass powder (Fig. 4(d)). The glass infiltrated core surface was sandblasted and then a aluminous porcelain powder was applied on the alumina coping and fused to form the all-ceramic crown in Fig. 4(e). The all ceramic crowns and copings fitted well with marginal openings.
REFERENCES
Piddock, V. and Qualtrough, A.J.E., 1 Dent, 1990, 18, 227-35. McLean, J.W., OperativeDent, 1991, 16, 149-56. Probster, L. and Diehl, J., QuintessenceInt.,1992, 23, 25-31. Tyszblat, M., U.S. Pat. No. 4772436, 1987. Reed, J.S. In: Principles of Ceramic Processing 2nd Ed., John Wiley & Sons, Inc., New York, 1995,525-41. 6. ASTM D 638-95, In: 1996 Annual Book of ASTM Standards vol. 08.01, Sect 8, American Society for Testing and Materials, Philadelphia, PA, 1996.
1. 2. 3. 4. 5.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
SURFAC E STRUCTUR E OF BIOACTIV E TITANIU M PREPARE D BY CHEMICA L TREATMEN T ^H.M. Kim, ^F. Miyaji, ^T. Kokubo, ^T. Suzuki, ^F. Itoh, ^S. Nishiguchi and ^T. Nakamura ^Department of Material Chemistry, Faculty of Engineering, Kyoto University, Kyoto 606-01, Japan, ^Chemical, Polymer, Bio, Technology Laboratory, Kobe Steel Ltd., Kobe 651 -22, Japan, department of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, Kyoto 606-01, Japan
ABSTRAC T Surface structural changes of titanium metal due to NaOH and heat treatments to induce its bioactivity and subsequent exposure to a simulated body fluid were investigated by Auger electron and X-ray photoelectron spectroscopies. A graded surface structure from outermost amorphous sodium titanate to metal substrate via titanium oxide, was formed on the metal by the above surface treatments. After subsequent soaking in a simulated body fluid, another graded structure from top apatite to the metal substrate via hydrated titania and titanium oxide, was formed on the surface of metal. These indicate that the sodiimi titanate released Na^ ions and converted into hydrated titania abimdant in Ti-OH group to induce the apatite formation m the shnulated body fluid. The graded structure across the interface between the apatite and the metal is believed to provide a strong bonding of the metal to the living bone through the apatite layer. KEYWORD S titanium, chemical treatment, bioactivity, surface structure, simulated body fluid (SBF), apatite INTRODUCTIO N In the past three decades, only a limited kind of ceramics, called bioactive ceramics, have been known to bond to living bone. However, the present authors clarified recently that bioactive metals, which spontaneously bond to living bone via formation of a bonelike apatite on their surfaces in the body as the bioactive ceramics do, can be prepared by subjecting titaniimi and its alloys to NaOH and subsequent heat treatments [1,2]. Because of intrinsically high fracture resistance and spontaneous bone-bonding ability, bioactive metals in this type are expected to provide significant advantages over ceramics and metals lacking in either of the above properties for load-bearing implant applications. In the present study, surface of titanium metal was investigated with special interests on compositional and structural changes at the near-surface cross-section in processes of the NaOH and heat treatments to induce its bioactivity and subsequent exposure to body environment. MATERIAL S AND METHOD S Commercially pure titaniimi (Ti) substrates (Ti>99.8%, Kobe Steel Ltd., Japan) were 215
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subjected to 5M-NaOH treatment at 60 C for 24 h and subsequent heat treatment at 600 C for 1 h. Thus surface-treated Ti substrates were soaked m an acellular simulated body fluid (SBF) with pH (7.40) and ion concentrations (H?t 142.0, K"^ 5.0, Ca^^ 2.5, Mg’" 1.5, Cr 147.8, HC03- 4.2, HPO/ 1.0, SO^^ 0.5 mM) nearly equal to those of human blood plasma. Compositions and structures of the Ti substrates after the surface treatments and subsequent soaking in SBF were analyzed at the siufaces and near-surface cross-sections by thin-fihn X-ray diffraction (TF-XRD; 265 lAl, Rigaku, Japan; incident angle 1 ), scanning electron microscopy (SEM; S2500CX, Hitachi, Japan), Auger electron spectroscopy (AES; ULVAC-PHI4300, Perkin-Elmer, England; primary electron beam voltage 5 keV with Ar-sputtering at a rate 48 nm/min and take-off angle 45 ) and X-ray photoelectron spectroscopy (XPS; ULVAC-PHI5500, Perkin-Elmer, England; X-ray source MgKa with Xe-sputtering at a rate 25 nm/min and take-off angle 45 ). RESULT S AND DISCUSSIO N Figure 1 shows SEM photographs and TF-XRD patterns of the surfaces of Ti substrates (A) after the NaOH and heat treatments and (B) after subsequent soaking in SBF for 3 days. A broad XRD pattern containing small peaks ascribed to sodium titanate (Na2Ti50^i) and rutile (Ti02) and porous network morphology in SEM picture indicate that essentially an amorphous sodixun titanate was formed on the metal after the NaOH and heat treatments. The XRD pattem and SEM picture after the subsequent soaking in SBF indicate that an apatite layer was formed on the surface-treated metal by the SBF-soaking. O: Apatite S: Na2Ti50ii R: Rutile (Ti02)
O
20
\
TI I Ti
’
1
30 40 26^/degre e
(B )
50
Figure 1. SEM photographs (left) and TF-XRD patterns (right) of the surfaces of Ti substrates (A) after the NaOH and heat treatments and (B) after subsequent soaking in SBF for 3 days.
Surface Structureof Bioactive Titanium Prepared by Chemical Treatment:H.M. Kim et al.
5J100 ^ ^^TTTT
217
1 ^10^ ^ 20-1
I 80. S 60-JS12H S
I 40 u 20
1 5o
oJ< 0
5000
10000
Depth/A
15000 ^
0
10000
20000
Depth/A
30000
Figure 2. AES depths profiles of the Ti substrates (A) after the NaOH and heat treatments and (B) after subsequent soaking in SBF for 3 days. Figure 2 shows AES depths profiles of the Ti substrates (A) after the NaOH and heat treatments and (B) after subsequent soaking in SBF for 3 days. The depth profile after the surface treatments shows that Na and 0 concentrations gradually decrease from the siuface to the interior while Ti concentration gradually increases toward the same direction. This indicates that a graded surface structure from the top amorphous sodium titanate to the metal substrate via titaniimi oxide, was formed on the metal by the surface treatment. The depth profile after the subsequent soaking in SBF shows that Ca, P and O concentrations
T\2p
01s I 455
r-
475
I
362
Ca2p I
I
I
r
-
34 2
P2p
W Wv^
T 146
1 126
Titanate OH H2O Oxide Pliosphat e OH H2O Oxide 53 8
I
534
Binding energy/e V
530
Figure 3. XPS spectra of the surface-treated Ti substrates after the subsequent soaking in SBF for 3 days.
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gradxxally decrease from the surface to the interior while Ti concentration gradually increases toward the same direction. This indicates that a graded structure from top apatite to metal substrate via titaniimi oxide, was formed on the metal via Na-release from the sodiimi titanate. Figure 3 shows XPS spectra of the the surface-treated Ti substrates after the subsequent soaking in SBF for 3 days. Gaussian deconvolution of 01s spectrum [3] in Fig. 3 indicates that a phosphate (solid line), i.e. the apatite, coexists with a titanate (dotted line) abimdant in Ti-OH group (535.3eV). Such titanate is assumed as a hydrated titania, which was converted from the sodiimi titanate via ion exchange of Na^ with H30^ to induce the apatite formation in the SBF [4]. As indicated above, the formation of hydrated titania, and thus the formation of apatite to be induced by it, are determined by the release of Na^ ion from the sodium titanate with a graded structure formed on the metal by the NaOH and heat treatments. This graded structure is assumed to enable the sodium titanate to be changed gradually into hydrated titania inducing the apatite formation from the surface to the interior. Lti this process, another gradient structure, which changes from the outermost apatite to the Ti metal substrate via hydrated titania and titanium oxide, is spontaneously formed on the surface of metal. The present authors reported that bonding of the apatite to the above surface-treated Ti metal istso tight that tensile fracture by an ASTM C633-type of adhesion test using Aralditefi glue occurs at the apatite-glue interface or in the glue with adhesive strength over 28 MPa [5]. Such tight bonding of the apatite to the metal is attributed to the graded structure from the outermost apatite to the metal substrate, which may enable a proper load transfer along the apatite-metal interface. CONCLUSION S In the process of a bonelike apatite formation on the titanium metal subjected to NaOH and subsequent heat treatments in body environment, a graded surface structure, which changes from an outermost apatite to the metal substrate via hydrated titania and titanium oxide with no distinct boundary, is spontaneously established on the surface of metal. This graded structure is believed to provide a strong bonding of the metal to living bone through the apatite layer. ACKNOWLEDGMEN T This study was supported by Grant-in-Aid for Scientific Research, the Ministry of Education, Science, Sports, and Culture, J^an. REFERENCES 1. Kim, H.M., Miyaji, F., Kokubo, T. and Nakamura, T. J. Biomed.Mater. Res., 1996, 32,409-417. 2. Yan, W.Q., Nakamura, T., Kobayashi, M., Kim, H.M., Miyaji, F. and Kokubo, T. J. Biomed Mater.Res., 1997,37, 1-11. 3. Asami, K. and Hashimoto, K. Corros.&/., 1977,17, 713-23. 4. Li, P., Ohtsuki, C , Kokubo, T., Nakanishi, K., Soga, N., Nakamura, T., Yamamuro, T. and de Groot, K. J. Biomed Mater.Res., 1994,28, 7-15. 5. Kim, H.M., Miyaji, F., Kokubo, T. and Nakamura, T. In: Bioceramics Volume 9, Elsevier, Oxford 1996, 301-304.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PREFABRICATE D BIOLOGICA L APATIT E FORMATIO N ON A BIOACTIV E GLASS-CERAMI C PROMOTE S IN VITR O DIFFERENTIATIO N OF FETA L RA T CHONDROCYTE S C. Loty\ S. Loty^ T. Kokubo^ N. Forest\ and J.M. Sautier^ Laboratoire de Biologie-Odontologie, Faculte de Chirurgie Dentaire, Institut Biomedical des Cordeliers, Universite Paris 7, 15-21 rue de I’Ecole de Medecine, F-75270 Paris Cedex 06, France, and ^Kyoto University, Faculty of Engeneering, Division of Material Chemistry, Yoshida, SA Kyo-Ku, Kyoto 606-01, Japan.
ABSTRAC T Rat chondrocyte cultures were carried out on three different substrata: glass-ceramic with and without a prefabricated surface apatite layer and glass-coverslips. A culture period of two weeks in the presence of ascorbic acid and 6-glycerophosphate resulted in chondrocyte differentiation and the expression of type II collagen and chondroitin sulfate. In addition, alkaline phosphatase activity was found to increase gradually during the culture period but the enzyme activity on the apatite-formed glass-ceramic was higher than on the ceramic without surface apatite layer and about twice that of the control glass-coverslips. This data indicates that the glass-ceramic surfece promotes chondrocyte differentiation mainly when a biological apatite layer was prefabricated. KEYWORDS : Chondrocyte - Ceramic-surface - In vitro - Differentiation INTRODUCTIO N Bioactive materials, such as calcium-phosphate ceramics, bioglasses, bioactive glassceramics are known to form a strong chemical bond with bone. Among bioactive ceramics, glassceramic containing apatite and woUastonite (A.W.G.C.) have been reported to have a relatively high mechanical strength and bond to living bone [1]. When the bioactive ceramic is implanted in bone, an apatite layer is naturally formed on the original ceramic surface which seems essential for the bone bonding. This work was undertaken to investigate the significance of ceramic surfaces on chondrocyte differentiation and metabolism. We examined the behavior of fetal rat chondrocytes cultured on three different substrata: A.W.G.C, A.W.G.C. on which an apatite layer was formed by immersing them in a simulated body fluid for 2 weeks, and glass coverslips. MATERIAL S AND METHOD S AW glass ceramic was synthesized by a method previously described [2]. Mirror-poUshed disks (15 mm in diameter, 1 mm thick) were provided by Nippon Electric Glass Co, Ltd (Japan) and some disks were incubated in a simulated body fluid to generate a surface apatite layer with an in vitro method developed by Kokubo et al. [1]. Chondrocytes were enzymatically isolated fix)m nasal cartilage of 21- day-old fetal Sprague Dawley rats, as previously described by Sautier et al. [3]. Briefly, nasal septa were aseptically dissected and cartilage fragments incubated in coUagenase (0.25%) and hyaluronidase (0.50%) for 2 h at 37 C. Then the cells were plated at 4 x 10"^ cells/cm^ directly onto A.W.G.C. disks. The chondrocyte phenotype was monitored by immunocytochemistry after 10 days of culture for type II collagen and chondroitin sulfate. Estimation of protein content was carried out using the Pierce BCA Protein Assay Kit (Pierce 219
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Chemicals). The specific activity of alkaline phosphatase was assayed in the cell layers as the released of p-nitrophenolfi-omp-nitrophenolphosphate [4]. RESULT S AND DISCUSSIO N Kokubo et al. [1] have previously demonstrated that when A.W.G.C. was immersed in simulated body fluids for two weeks at 36.5 C the suface of the ceramic was completely covered with a layer of apatite which consisted of carbonate-containing hydroxyapatite with compositional characteristics of biological bone apatite. Rat chondrocytes enzymatically released fromthe nasal cartilage of 21-day-old fetuses were directly seeded onto A.W.G.C. disks with an without the surface apatite layer. In addition, plastic glass coverslips were used as controls. During the first days of culture, chondrocytes attached to all substates and reached confluency on day 4 of culture as previously reported [5]. Immunostaining with anti-type 11 collagen (Fig. la) and chondro’itin sulfate (Fig. lb) performed on day 10 of culture showed strong fluorescence associated with clusters of round cells on disks on which an apatite layer was formed but also on others substrates (data not shown). This data indicated that on all substrates tested the cells maintained their phenotype and synthesized type n collagen and chondro’itin sulfate, both halhnarks of the chondrocyte phenotype. In the same way, Li Vecchi et al. [6] reported a high percentage of type II collagen when articular chondrocytes were grown on a porous polymer. Protein content slowly increased to the thirtheeths day on each substrates but then the number of proteins synthesized on glass coverslips increased up to day 22 of culture (Fig. 2a). In contrast, on the glass-ceramic with and without an apatite layer protein content no more increased during the following days. These discrepancies, may be due to diflfCTences in proliferation rates. Calcium phosphate ceramics have been shown to both stimulate [7] and inhibit [8] proUferation of osteoblasts. The mhibitory effect on proUferation may depend on the form in which the ceramics were used. For example, when calcium phosphate was used as chips, the material surfece inhibited costochondral chondrocyte proliferation [9], whereas when used in solution form the same material was competent growfli promoter for osteoblasts [10, 11],
Figure 1. Immimolocalization of type n collagen (a) and chondroitin sulfate (b) on day 10 of cultures on A.W.G.C. with a biological apatite layer, x 200.
PrefabricatedBiological Apatite Formation on a Bioactive Glass-Ceramic: C. Loty et al.
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Figure 2. Kinetics of protein content (a) and time course of alkaline phosphatase activity (b) (Control = glass coverslips ; A.Wa = A.W. glass-ceramic with a prefahricated apatite layer ; A.W. = A.W. glass-ceramic without prefahricated apatite layer.) Alkaline phosphatase (ALP) activity also gradually increased after the 3rd day of culture and was comparable on each surfaces still day 14 (Fig. 2b). In addition, on day 22 of culture, the activity of the enzyme on A.W.G.C showed a 40% increase compared to that of the control glass coverslip and the ALP activity on A.W.G.C with the prefabricated apatite layer showed about twice that of control. Our results are in accordance with those of Ohgushi et al. [12] whose demonstrate the positive effect of apatite formation on osteogenic differentiation in rat marrow cell cultures. In conclusion, it was shown in this study that the glass-ceramic surface promotes chondrocyte differentiation and that the promotion can be fiirther enhanced by the formation of an apatite layer on the surface of the ceramic, very similar to natural bone mineral. ACKNOWLEGDMEN T This work was supported by a grantfix)mCANAM - Assistance Publique - Hopitaux de Paris.
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REFERENCE S 1. Kokubo, T., Kitsugi, T. and Yamamuro, T. CaP-rich layer formed on high-strength bioactive galss-ceramic A-W. J. Biomed Mater.Res. 1990, 24, 331-343. 2.
Kokubo, T., Shigematsu, M., Nagashima, Y., Tashiro, M., Nakamura, T., Yamamuro, T. and Higashi, S. Apatite and woUastonite containing glass-ceramic for prosthetic application. Bull. Inst. Chem.Res. Kyoto Univ. 1982, 60, 260-268.
3.
Sautier, J.M., Nefussi, J.R. and Forest, N. In vitro differentiation and mineralization d cartilaginous nodules fromenzymatically released rat nasal cartilage cells. Biol Cell 1993, 78, 181-189.
4.
Engstoom, E. Aspects of molecular structure of bone. In: The Biochemistry and Physiology of Bone Volume 1, 2nd Ed., Bourne, G., ed.. Academic Press, New York 1972, 785-793.
5.
Loty, C , Forest, N., Boulekbache, H., Kokubo, T. and Sautier J.M. Behavior of fetal rat chondrocytes cultured on a bioactive glass-ceramic. J. BiomedMater.Res 1997, in press.
6.
Li Vecchi, A.B., Tombes, R.M. and LaBerge M. In vitro chondrocyte collagen deposition within porous HDPE: Substrate microstructure and wettability effects. J. Biomed Mater.Res. 1994, 28, 839-850.
7.
Gregoire, M., Orly, I. and Menanteau, J. The influence of calcium phosphate biomaterials on human bone cell activities: an in vitro approach. Biomed. Mater. Res. 1990, 24, 489494.
8.
Lanzer, W.L., Crane, G., Howard, G.A., Davidson J.A. The effect of implant wear debris on human bone cell proliferation in vitro. In: 16th Annual Meeting of the Society for Biomaterials1990, 293.
9.
Hambleton, J., Scharwt, Z., Khare, A., Windeler, S.W., Luna, M., Brooks, B.P., Dean, D.D. and Boyan, B.D. Culture surfaces coated with various implant materials affect chondrocyte growth and metabolism. J. OrthopaedicRes. 1994, 12, 542-552.
10.
Cheung, H.S. and McCarty, D.J. Mitogenesis indiced by calcium-containing crystals: role of intracellular dissolution. Exp. Cell Res. 1985, 157, 63-70.
11.
Cheung, H.S. and Haak, M.H. Growth of osteoblasts on porous calcium phosphate ceramic: an in vitro model for biocompatibility study. Biomaterials1989, 10, 63-67.
12.
Ohushi, H., Dohi, Y., Tamai, T., Tabata, S., Okunaga, K. and Shobuya, T. Osteogenic differentiation of cultured marrow stem cells on the surfece of bioactive glass-ceramic. Biomed.Mater.Res. 1996, 32, 341-348.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
THE EFFEC T ON MECHANICA L PROPERTIE S BY OSTEOBLASTI C CEL L INGROWT H IN MACROPOROU S SYNTHETI C HYDROXYAPATIT E AND INTERPOR E 200 E. Nordstrom*, H. Ohgushi, H. Yoshinari, S. Tamai, T. Yokobori* *Department of Materials Science and Engineering, School of Science and Engineering, Teikyo University, Toyosatodai 1-1, Utsunomiya 320, Japan and Department of Orthopaedic Surgery, Nara Medical University, Kashihara, Nara 634, Japan
ABSTRAC T In the subculture condition, in the presence of beta-glycerophosphate(BGP), ascorbic acid, and dexamethasone, osteoblastic cells began to appear at about one week post culturing, and many macroscopic mineralized nodules were observed on each culture substrata at two weeks. Osteoblastic cell ingrowth was detected in coral derived hydroxyapatite. However, the effect of the ingrowth on the mechanical strength was only marginal. KEYWORD S - Hydroxy(l)apatite - Bioceramics - Cells - Strength INTRODUCTIO N It seems that the fundamental phenomenon that leads to tissue bonding is surface transformation, which can influence surrounding cellular activity and result in stimulatory effects on cell differentiation. In this investigation we use tissue culture methods to investigate osteogenic cell differentiation in macropores of synthetic and coral derived hydroxyapatite (Interpore 200 ). Hydroxyapatite, both dense and porous and glass ceramics, shows bioactive behavior and tissue bonding[l-13]. This is considered as true bone bonding when implanted into biological bone. The first person to discover the bone bonding phenomena was L. L. Hench[4]. After this many types of bioactive ceramics has been developed[l-8,14,15]. These materials show surface changes, including dissolution and precipitation[16-23]. Whereas nonbioactive ceramics show neither precipitation nor bone bonding[l-3,19,20]. Ohgushi et al. has previously demonstrated surface dependent osteoblastic differentiation on bioactive ceramics[ 13,24]. Even if the surface change is favorable for tissue bonding and promoting cell differentiation, it was not sure, until Ohgushi et al. published their results on osteogenetic differentiation of cultured marrow stromal stem cells on the surface of bioactive glass ceramics, whether or not in vitroprefabrication of apatitic layer on the ceramic surface can stimulate cell differentiation[25]. The aim is to investigate the surface properties of the chemically surface active materials in a biological environment. The present paper will focus surface activity of porous materials and how the osteoblastic activity in forms of ingrowth into pores affect the mechanical properties of the ceramics. 223
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MATERIAL S AND METHOD S The materials used for the investigation were prepared from commercially available products. The materials used v^ere Interpore 200 and macroporous hydroxyapatite (HA, Merck, Germany) which was prepared from large round shaped presintered granules of the size 150 250 jim. The granules were sintered together using hot pressing (Shimadzu Co. Ltd., Japan), a constant pressure corresponding to about 1 kbar was applied and the temperature at which the load was introduced was 800 C. The heating rate was 20 C min"^ and maximum temperature was 1400 C. The holding time at maximum temperature was about 5 min. The HA/mica composite was prepared as described earlier[14]. The materials origin from commercial products. The HA grain size was less than 6 |im (Mitsubishi Materials Co. Ltd., Japan) and the mica grain size was less than 45 |im (Norwegian Talc Minerals AS, Norway). The method used for sintering was hot pressing. The constant pressure of about 60 MPa was applied on the material at around 800 C at each sintering cycle. The total amount of the cycles were three. The first two cycles had a maximum temperature of 1125 C and the last one 1150 C. The holding time at maximum temperature was one hour and the heating rate was 20 Cmin\ (a)
CUT O F F - >
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if-(
^
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span length Figure 1. The picture (a) shows how the samples for the 3-point bend test were prepared. Length A - B = 75 nmi, width C - D = 40 mm and thickness E - F = 20 mm. In the figure beneath (b) the 3-point bend appratus is described. Span length = 5 mm ( h = 20 mm and b = 40 mm) and crosshead speed used was 0.1 mm/s (for all specimens).
Effect on Mechanical Propertiesby OsteoblasticCell Ingrowth: E. Nordstrom et al.
225
The materials were produced into cylindrical shape in a graphite dye. The diameter of the cylinders were 8 mm. Each specimen was then sliced into 2 mm thin discs using diamond saw blade. After slicing the specimens were cut to bars by removing the edges of the discs (Fig. 1 a). The discs were used for the cell tests. A total of 40 samples of each material were tested by 3point bend test (Fig. 1 b), before and after biological testing. For the marrow cell preparation and culturing, the primary cell culture was done by using marrow cells from the bone shaft of a rat femora. The medium consisted of Eagle Minimal Essential Medium (MEM) and 15% fetal calf serum (FCS). The cell culture conditions essentially were the same as those reported earlier, with minor modifications[26,27]. Ascorbic acid and pglycerophosphate were used in the culture medium because they appear to be a requirement for formation and mineralization of bone-like extracellular matrix in organ and tissue culture [26,27]. While dexamethasone is not a requirement for osteogenesis by fetal cells in vitro , osteogenesis appears to be enhanced when appropriate concentrations of dexamethasone are added to the culture medium[26,27]. For the surface analysis, the surfaces of the materials were subjected to scanning electron microscopy (SEM). In the SEM observation, platinum film was coated on the surface on the specimen by ion sputtering and scanning electron microscope was used for imaging. SEM (JEOL, Japan) was employed under 3.0 kV low voltage. RESULT S AND DISCUSSIO N In the subculture condition, osteoblastic cells began to appear at about one week ( -- 6 days ) postculture, and many macroscopic nodules were observed on each culture substrata at two weeks ( - 1 3 days). It was found that in Interpore 200"^*^ the osteoblastic cells grew into the pore space. This in turn gave a sligth increase of the 3-point bend strength values. Table 1 presents the 3-point bend Table 1. Bend strength and share stress values (kg/mm^) before and after in vitrotesting.
Sample
in vitrotest
Bending Stress, Av. +/- S.D.
Share Stress, Av. +/- S.D.
Fairy
before
0.5438 +/- 0.2304
0.1365 +/- 0.0586
Fairy (dex+)
after
0.5481 +/- 0.1382
0.1372 +/- 0.0382
Fairy (dex-)
after
0.6724 +/- 0.2125
0.1694 +/- 0.0593
before
1.6286 +/- 0.3409
0.4673 +/- 0.0882
after
1.6238 +/- 0.1912
0.4050 +/- 0.0482
before
3.9161 +/- 0.3519
1.0568 +/- 0.0866
1 Angel Angel
1 Devil
j
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Strength results. From the results it can be seen that the macroporous synthetic hydroxyapatite {Angel) showed the same bending strength before and after in vitro test. In the coral derived samples (Fairy) a sligth increase in 3-point bend strength was observed after the in vitro test. However, we have to consider the differences in porosity between the two sample types. The dex (-) specimens, without dexamethasone, showed higher bend strength values than the dex (+) specimens. However , all the differences were marginal and at maximum the difference was about one fourth for the dex (-) specimens compared with dex (+) and test results before the in vitro test. For the HA/mica (Devil) only before values are available, because this type of sample were discarded from the test series before in vitro.It was considered that the pore size ( only about 10 jim), was too small for osteoblastic cell ingrowth. SUMMAR Y In conclusion the present work showed that macroporous synthetic hydroxyapatite was not penetrated by osteoblasts and therefore an increased mechanical strength could not be detected. In the coral derived hydroxyapatite some osteoblastic differentiation could be obtained in the pores, which in turn lead to a small increase of the mechanical properties. ACKNOWLEDGEMEN T The authors would like to thank The Japan Society for the Promotion of Science for their support. REFERENCE S 1. Hench, L. L., Ann. NY Acad. ScL, 1988, 523, 54. 2. Hench, L. L , in "The Bone-Biomaterial Interface", edited by J. E. Davies (University of Toronto Press), 1991, p. 33. 3. Gross, U. M., Schultz, H. J., and Strnz, V., Ann. NY Acad. Sci., 1988, 523, 211. 4. Hench, L. L., Sphner, R. J., Allen, W. C , and Greenlee, T. K., J. Biomed. Mater. Res. Symp., 1971,2,117. 5. Kokubo, T., Ito, S., Sakka, S., and Yamamuro, T., J. Mater. Sci, 1986, 21, 536. 6. Nakamura, T., Yamamuro, T., Higashi, S., Kokubo, T., and Ito, S., J. Biomed. Mater. Res., 1985, 19, 685. 7. Nordstrom, E. G., Niemi, L., and Miettinen, J., in "Ceramics in Substitutive and Reconstructive Surgery", edited by P. Vincenzini (Elsevier, Amsterdam, 1991) Materials Science Monographs, 69, p. 335. 8. Nordstrom, E. G., Niemi, L., and Miettinen, J., Bio-Med. Mater.Eng.,1992, 2, 115. 9. Jarcho, M., Clin. Orthop.,1981,157, 259. 10. Kato, K., Aoki, H., Tabata, T., and Ogiso, M., Biomat.Med. Devices Artif.Organs, 1979, 7, 291. 11. Ohgushi, H., Goldberg, V. M., and Caplan, A. I., J. Orthop.Res., 1989, 7, 568. 12. Ohgushi, H., Okumura, M., Tamai, S., and Shors, E. C , J. Biomed.Mater. Res., 1990, 24, 1563. 13. Okumura, M., Ohgushi, H., and Tamai, S., Biomaterials,1991,12,411. 14. Nordstrom, E. G., Her0, H., and J0rgensen, R.B., Bio-Med. Mater.Eng., 1994, 4[4], 309. 15. Nordstrom, E. G., ElectronMicroscopy, 1992, 2, 441. 16. LeGeros, R. G., Orly, I., Gregoire, M., and Daculsi, G., in "The Bone Biomaterial Interface," edited by J. E. Davies (University of Toronto Press), 1991, p. 76.
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17. Radin, S., and Ducheyne, P., 7. Biomed.Mater.Res,, 1993, 27, 35. 18. Radin, S., and Ducheyne, P., J. Biomed.Mater.Res., 1994, 28, 1303. 19. Kokubo, T., Ito, S., Huang, T., Hayashi, T., Sakka, S., Kitugi ,S., and Yamamura, T., J. Biomed.Mater.Res., 1990, 24, 331. 20. Kokubo, T., Kushitani, H., Sakka, S., Kitugi, T., and Yamamuro, T., J. Biomed. Mater. Res, 1990,24,721. 21. Nordstrom, E. G., Hara, T., and Her0, H., Bio-Med. Mater. Eng., 1996, 6[4], 73. 22. Nordstrom, E. G. ,and Karlsson, K. H., Bio- Med. Mater.Eng., 1992, 2, 185. 23. Nordstrom, E. G. ,and Karlsson, K. H. ,7. Mater.Sci. Mater.Med., 1990,1, 182. 24. Ohgushi, H., Dohi, Y., Tamai, S., and Tabata, S.,J.Biomed.Mater. Res., 1993, 27, 1401. 25. Ohgushi, H., Dohi, Y., Yoshikawa, T., Tamai, S., Tabata, S., Okunaga, K., and Shibuya, T., J. Biomed.Mater.Res. 32 in press (production No 1940). 26. Ohgushi, H. , Dohi, Y., Katuda, T. , Tamai, S. , Tabata, S., and Suwa, Y., J. Biomed. Mater. Res. 32 in press (production No 1939). 27. Maniatopoulos, C., Sodek, J., and Melcher, A. H., Cell TissueRes., 1988, 254,317.
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Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CHARACTERIZATIO N AN D CELL REACTIO N OF a-TCP- AN D HAp-COATING S ON TITANIU M PLAT E
M.Ohgaki, S.Nakamura, and M.Akao Institute for Medical and Dental Engineering, Tokyo Medical and Dental University. 2-3-10 Kanda Surugadai,Chiyoda-ku,Tokyo 101, JAPAN. ABSTRAC T Titanium plates coated with a-tricalcium phosphate(a-TCP) and coated with hydroxyapatite (HAp) were prepared by plasma spraying techniques. The characterizations for the a-TCP and HAp coatings were investigated by X-ray diffraction analysis, SEM observation, X-ray fluorescence analysis, and surface roughness measurements. In the HAp coatings on titanium plate, broad amorphous-like peaks were observed for the interior layers of coating near to the substrata, while crystalline peaks were observed for the surface layer. In the a-TCP coating, crystalline peaks were observed for the whole coating layers. Cytotoxicity of these materials were evaluated in vitro. Results from the proliferation tests of L-cells and MC3T3-E1 cells on the coating layers indicated that both the a-TCP- and HAp-coatings on titanium plate had better biocompatibility than the titanium plate without coatings. KEYWORD S : a-TCP, hydroxyapatite, plasma coating, characterization, cytotoxicity INTRODUCTIO N a-tricalcium phosphate (a-TCP, Ca^(?0^)^ ) and hydroxyapatite (HAp, C2i^Q(FO^)^(OU\ ) have good biocompatibility. But, the brittleness of bulk ceramics of a-TCP and HAp makes them incapable of being used in load bearing. To obtain a bioactive biomaterial with excellent mechanical properties for load bearing, bioactive HAp have been applied on to titanium implant by plasma-spraying technique[l]. Recently, titanium coated with a-TCP was prepared by plasma spraying P-TCP powders on titanium plates by authors[2]. Plasma spraying will induce changes in the phase composition, structure and other properties of the starting materials[3-4]. There are reports of in vivo and in vitroexperiments in the presence of HAp[4-8]. In the present study, the a-TCP coated titanium and HAp coated titanium were investigated on the crystal structure and characterizations. Cytotoxicity of the a-TCP coated titanium plate was evaluated by cell culture techniques. MATERIAL S AND METHOD S Preparation of a-TCP and HA p coatings on Titanium plate Titanium plates ( 22.0x22.0x2.0mm^), conforming to JIS II, were used as substrata. Before plasma spraying, the substrate plate surfaces were blasted with alumina grit to roughen the 229
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surface and cleaned in acetone bath to remove organic contaminants. For a-TCP coatings, ^TCP powders, heated at 1100’ C, were used as spraying materials. For HAp coatings, HAp powders, heated at 1150* C, were used. All coatings used in this study were applied by means of an atmospheric plasma spray technique (MCN System, Metteco Co.Ltd.). The plasma flame was composed of plasmatized Ar-Hj gas using an arc method. The thickness of the coated layers were about lOOjum. Chemical analysis, X-ray deiffractometr y and SEM observatio n The chemical properties of a-TCP and HAp coated layers were investigated by X-ray diffraction, scanning electron microscope(SEM), X-ray fluorescence analysis, and surface roughness measurements. The crystallinity of coating layers, particularly distinction between the surface layer and the interior layer, were investigated by the thin film X-ray diffraction analysis and the SEM observation for cross sections of the coating layers. Thin film X-ray diffraction patterns were measured with various fixed angles of incident beam, and X-ray irradiation area on sample face was fixed for every measurement. Compositions and densities of those coating layers were measured by X-ray fluorescence analysis. Cell proliferation The a-TCP- and HAp-coated titanium plates were sterilized by heating at 180* C, for Ih and also by exposure to ultraviolet light for Ih. L-cells were grown in Eagle’s minimum essential medium (MEM) supplemented with 10% foetal bovine serum (FBS); MC3T3-E1 cells, in a-MEM supplemented with 10% FBS. The cells (6.6x10^^ cells/ml for L-cells, 5.0x10^ cells/ml for MC3T3-Elcells) in 2ml of the medium were added to 35mm culture dishes and incubated at 37* C in 5% CO2 atmosphere. The sample plates were applied to the culture dishes and placed in the centre of the dishes. Cytotoxicities of coating plates in vitrohave been evaluated by cell growth measurements and morphological observations. Cell multiplication in the cultures were measured by counting the cells at days 1, 2, 3 and 4. Cultures on culture dish and on titanium plate were used as controls.
29 (CuKa)/degree (a)
29(CuKa)/degree (b)
Figure 1 Thin film X-ray diffraction patterns of a-TCP coatings (a), and HAp coatings (b).
Characterizationand Cell Reaction of a-TCP- and HAp-Coatings on Titanium:M. Ohgaki et al.
Titanium plate
HAp coating layer
Titanium plate
HAp coating layer
N
’<
231
(b)
(a)
Figure 2 SEM photographs of cross sections for HAp coatings (a), and HAp layers etched with acetic acid (b). RESULT S AND DISCUSSIO N Characterization The starting p-TCP and HAp powders, used for plasma splaying, had respective single phase under X-ray diffraction examination. Phases in both coating layers before and after the cell culture remained unchanged. Densities of a-TCP coating had about the same value as that of HA coating, 2.60g/cm^ for a-TCP layer and 2.59g/cm^ for HAp layer, determined by X-ray fluorescence analysis. The surface roughness (Ra) for a-TCP plasma coating was about 9.59 |im (0.04 \xxafor plastic culture dish, 0.24 \x,mfor titanium plate). The diffraction intensities of a-TCP coating and HAp coating, obtained by the thin film X-ray diffraction method, become larger in proportion to background as that measured with higher angles of incident X-ray paths. For a-TCP coating, the interior layer of coating had the same crystallinity as the surface layer, considered by peak profiles(Figure 1(a)). On the other hand, for the HAp coating, broad amorphous-like peaks appeared at higher angles of fixed incident Xrays(Figure 1(b)). Thus, HAp coatings had less crystallinity in the interior layers of coating near to substructure than the surface layers. The broad peaks remarkably enlarged beyond 7.0** of incident X-ray angle, which angle could be estimated to be 40 |xm of depth participated in diffraction patterns, calculated by X-ray path lengths of the coating layer.
O : Control X : a-TCP coating titanium plate # : Ttanium plate
4
5
Days Figure 3 Cell proliferation curves of MC3T3-E1 cells (a) and L-cells (b), cultured on a-TCP coatings.
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Table 1 Relative growth rates of MC3T3-E1 cells at 4 days proliferated on titanium plate with and without HAp coatings. Titanium plate with HAp coating Titanium plate
89% 76%
In fact, two types of coating layer were observed by SEM analysis in cross section of sample plate(Figure 2(a)). The thickness of each coating layer of the two types, estimated by SEM observations, was in agreement with the results from the thin film X-ray diffraction method. The interior layer was exposed by etching with acetic acid(Figure 2(b)). The’ amorphous phase, existing in the plasma-splayed HAp coatings, was regard to be caused by rapid solidification or by quenching in striking a cool substrate. Cytotoxicity The morphological evaluation (phase-contrast micrographs and SEM observations) indicated that both L cells and MC3T3-E1 cells proliferated well in every cultured dishes. For the a-TCP coated titanium plate, both cells were well grown and sticked to the edges of samples. Cell proliferations indicated that the a-TCP coated plate had better biocompatibility than the titanium plate without coating(Figure 3). The relative growth rates of MC3T3-E1 cells at 4 days were 64% for the a-TCP coatings and 47% for the titanium plate. On the other hand, the growth rates of L-cells at 4 days were 51% for the a-TCP coatings and 44% for the titanium plate. The growth rates for the a-TCP coatings plate were significantly higher than those for the titanium plate, especially on the osteoblastic MC3T3-E1 cells. In case of titanium plate with HAp coatings, cytotoxic effects were similar to those of a-TCP coating plate(Table 1).
ACKNOWLEDGMEN T The authors are grateful to Isao Yamaji at Philips Japan Ltd. for his permission to use thin film X-ray diffractometor. REFERENCE S 1. de Groot, K., Geesink, R., Klein, C.P.A.T. and Serekain, P. J. Biomed.Mater.Res. 1987, 21, 1375-1381. 2. Kuroyama, Y., Aoki, H., Higashikata, M., Yoshizawa, K., Nakamura, S., Ohgaki, M. and Akao, M. J. JapaneseSoc. Dent.Mater.Dev. 1993,12, 528-534. 3. Tampieri, A., Celotti, G., Szontagh, F. and Landi, E. /. Mater.Sci. Mater.Med. 1997, 8, 29-37. 4. Yang, C.Y., Wang, B.C., Chang, W.J., Chang, E. and Wu, J.D. J. Mater.Sci. Mater. Med 1996,7,167-174. 5. Weng, J., Liu, X., Zhang, X., Ma, Z., Ji, X. and Zyman, Z. Biomaterials1993,14, 578-582. 6. AlUot-Licht, B., Gregorie, M., Orly, I. and Menanteau, J. Biomaterials1991, 12, 752-756. 7. Gomi, K., Lowenberg, B., Shapiro,G. and Davies, J.E. Biomaterials1993,14, 91-96. 8. Kuroyama, Y. J. JapaneseSoc. Dent.Mater.Dev. 1993, 12, 773-789.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TH E ECTOPI C OSTEOCONDUCTIO N MODE L (Periosteum/H A composite implantation at subcutaneou s site) H, Ohgushi, M. Okumura, T. Yoshikawa, H. Ishida, H. Yajima and S. Tamai Department of Orthopedics, Nara Medical University, Kashihara city, Nara 634, Japan ABSTRAC T Vascuiarlized rat periosteum was wrapped around porous hydroxyapatite ceramics (HA) and transferred to a rat subcutaneous site. The periosteum/HA composites were harvested 4 weeks post implantation and used for histological analysis. New bone formation was detected outside of HA (original location of the periosteum) as well as m the pore areas away from the location of periosteum. In the pore areas, the bone formation started on the HA surface and fibrous tissue interposition was not observed between thus formed bone and HA. The results imply that the osteogenic capacity of the periosteum engaged in the bone formation outside of the HA, whereas undifferentiated mesenchymal cells migrated from invading tissue into the pore areas initiated bone formation directly on the HA surface. These findings indicate the importance of the HA surface for osteoblastic differentiation and we propose the subcutaneous implantation of the composite as an ectopic osteoconduction model. KE Y WORD S Hydroxyapatite, Periosteum, Marrow cells, Bone formation INTRODUCTIO N The osteogenic capacity of stromal stem cells that reside in marrow cell population is well known and we reported that the subcutaneous implantation of marrow/HA[l-7] and marrow/porous alumina [6,7]composites can show new bone formation. Interestingly, the new bone formation always starts on the HA surface whereas it starts away from the alumina surface. Therefore, we have defined that the surface of bioactive materials can support osteoblastic phenotype expression, whereas the surface of non-bioactive materials can not support the differentiation[5,6]. The osteogenic capacity of the thin tissue covering bone (periosteum) is also known and we reported that subcutaneous implantation of periosteum/HA composites resuhs in new bone formation. Characteristics of the new bone formation were analyzed histologically as well as biochemically usmg assay of alkaline phosphatase and bone Gla protem (osteocalcin) content[8,9]. Histological analysis also confirmed that in the periosteum/HA composites, the de novo bone appeared outside as well as inside of HA. This indicated the HA surface dependent and independent osteoblastic differentiation [9]. We have reanalyzed the bone formed in the periosteum/HA composites, and propose that the composite implantation at subcutaneous sites can be useful to analyze the osteogenic response of the implants as an ahemative osteoconduction model at ectopic sites, MATERIAL S AND METHOD S Ceramics: Coralline hydroxyapatite ceramics (fully interconnected pores, measuring 190 to 230 jLun in diameter, with an average pore volume of 50 to 60 %; Interpore International, Irvine, Calif 233
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USA) were cut into a cube (20mg) and used for this study. Periosteum/HA composite : A medial portion of tibial periosteum (size 3 x 15mm) of 7 week old Fischer rat was elevated on a vascular pedicle of the saphenous vessels. The periosteum was wrapped around a HA and transferred to a subcutaneous pouch on the medial aspect of the thigh [8]. The implants were harvested at 4 weeks post implantation. RESULT S AND DISCUSSIO N A composite of periosteum and HA was a viable osteogenic implant because the composite was made with vascularized periosteum. As shown in Fig. 1, the composite showed new bone formation outside of the HA implants. The periosteum was wrapped around the HA and therefore the periosteimi was located outside of HA implants. The coincidence of the periosteum and the new bone location suggest that the bone was derived from the periosteum. However, bone formation could also be detected in many pore areas (Fig. 2). As the bone formed in pore areas located away from the original location of the periosteum. An unportant question is, what is the origin of the bone in the pore areas. As the pore interconnection is complete for HA used in this study, it seems the bone in the pore areas was derived from the growing of bone which was located outside the HA into the pore regions. However, as shown in Fig. 2, there was no intervening fibrous tissue between thus formed bone in the pore areas and HA surface, and in the vicinity of the bone, osteoblastic cells appeared on the HA surface. This histological feature indicates that the bone in many pore areas began on the HA surface and therefore not from ingrowth of the bone which was located outside of HA. Although, the bone ingrowth could have occurred in very restricted pore areas adjacent to the HA surface. In this regard, we reported that subcutaneous implantation of marrow cell/HA composites showed new bone formation. The bone was derived from marrow cells and the formation began on the HA surface through the cascade of osteoblastic differentiation [5]. The HA surface dependent bone formation was also detected in the composites of osteoinductive protein factors (bone morphogenetic proteins; BMPs) and HA [10]. These our previous results indicate the undifferentiated cells in the pore areas can attach on the HA surface followed by osteoblastic differentiation. If the cells in the pore areas are marrow stromal cells (recently the cells are called as mesenchymal stem cells)[l 1], the cells by themselves can differentiate into osteoblasts and if the cells are undifferentiated fibroblastic cells, the osteoblastic differentiation requires an osteo-
Fig. 1 Periosteum/HA composite at 4 weeks post implantation (x 250). Bone formation (B) is detected outside of the HA. Intervening fibrous tissue(arrows) is seen between the bone (B) and HA.
The Ectopic OsteoconductionModel: H. Ohgushi et al.
235
Fig. 2a Periosteum/HA composite at 4 weeks post implantation ( x 20). Bone formation (B) is detected ; many pore areas.
k’>
HA
HA
Fig. 2b Higher power magnification of the rectangular area seen in Fig. 2a. There is no intervening fibrous tissue between bone (B) and HA surface. Osteoblastic cells (arrows) are seen on the HA surface.
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inductive action. Because the mesenchymal stem cells reside in periosteum, the stem cells migrated from the periosteum into the pore regions can show new bone formation on the HA surface. It is also well known that the existence of BMPs in bone tissue (bone matrix), and as described above, exogenously added BMPs can induce osteoblastic differentiation primarily on the HA surface. Based on these results, we suppose mesenchymal stem cells from surrounding periosteum and protein factors (BMPs) in newly formed bone tissue can participate in the osteogenesis in HA pore areas. When bioactive materials such as HA was implanted in bony defects, Osbom reported[ 12] that bone bonding was accomplished through the cascade of bonding osteogenesis. That is, the new bone formation occurs on the bioactive material’s surface. Though the mechanism of the bonding osteogenesis is not ftiUy understood, the preexistence of bone tissue near the implantation sites is the source of osteogenic cells or osteoinductive factors to show new bone formation. The phenomenon of the new bone formation is also known as osteoconduction. The term of osteoconduction implies the appearance of new bone tissue around implanted materials. When the implanted materials are bioactive, as reported by Osbom, the bone bonding occurs by the cascade of bonding osteogenesis and when the materials are bioinert, the bone contact occurs by the cascade of contact osteogenesis [12]. Therefore, osteoconduction is the phenomena at orthotopic sites which leads to new bone formation around implanted materials not related to the materials properties. The present experimental model of periosteum/HA composite which was implanted at subcutaneous sites also showed new bone formation around the implanted materials. The bone was derived from the periosteum which surrounded the HA and therefore, present subcutaneous implantation of periosteum/HA can be regarded as an osteoconduction model at ectopic site. Importantly, the bone formation in the pore regions initially occurs on the HA surface (bonding osteogenesis) and suggest that present osteoconduction model can be available to evaluate the materials properties regarding bone bonding. REFERENCES 1. Ohgushi, H.,Okumura, M.,Tamai S.and Shors,E.C. J. Biomed.Mat.7^^^. 1990,24,1563-1570 2. Okumura, M., Ohgushi H. and Tamai, S,Biomatenals1991,12, 411-416 3. Ohgushi, H., Okumura,M., Yoshikawa, T., Inoue, K., Senpuku, N. and Tamai, S., J. Biomed Mat. /?e5.1992,26,885-896 4. Yoshikawa, T., Ohgushi, H. Okumura, M., Tamai, S., Dohi, Y. and Moriyama, T., Calcified TissueInternational1992 50,184-188. 5. Ohgushi, H.,Dohi, Y.,Tamai, S.and Tabata, SJ.BiomedMat. Res.1993.27,1401-1407. 6. Ohgushi, H., Okumura, M. Yoshikawa, T. Tamai, S. Tabata, S and Dohi, Y. in BonebondingBiomaterials,Helthcare Comm. Publ, the Netherland, 1992,pp.47-56. 7. Takaoka, T., Okumura, M., Ohgushi, H., Inoue, K., Takakura Y. and Tamai, S. Biomaterials1996,17,1499-1505 8. Ishida, H, Tamai, S., Yajima, H., Inoue, K., Ohgushi, H. and Dohi, Y., Plast. Reconstr. Surg. 1996,97,512-518 9. Ishida, H., Ohgushi, H., Inoue, K., Yoshikawa T., Yajima, H., Tamai, S. and Y. Dohi. Bioceramics 1996,9,73-76 10. Ohgushi, H., Okumura, M., Inoue, K., Dohi, Y.,Tamai, S.Murata, M., and Kuboki,Y. Bioceramics 1995,8,61-67 11. Caplan, A.I. Clin. Orthop. 1990,261,257-267 12. Osbom, J.F. and Newesely, H. Biomaterials 1980, (John Wiley and Sons, Ltd, edts: G.D.Wmter, D.F. Gibbons and H. Plenk Jr.) 1982, 51-58
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CONDITION S OF THE COPRECIPITATIO N OF CALCIU M HYDROXYAPATIT WIT H ZrOi , ZrOi+YjOa , AI2O3 FROM AQUEOU S SOLUTION S USIN G AMMONI A
E
V.P. Orlovskii, Zh. A. Ezhova and E.M. Koval Institute of General and Inorganic Chemistry of Russian academy of sciences, Moscow, Leninskii pr. 31, 117907 Russia
ABSTRAC T The systems CaCl2-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20, CaCl2-AlCl3-(NH4)2HP04-NH3-H20, CaCl2-YCl3-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20 are studied at the 25 C using the solubility method (Tananaev method of residual concentrations). The conditions of coprecipitation of calcium hydroxyapatite (HA) with hydroxides of metals (Zr; Al; Zr with addition of Y) are determined. Using the chemical analysis, XRDA and IRS methods it is shown that after the coprecipitated phases calcined at 900 C the homogenous mixture of HA and oxides (Zr02; AI2O3; Zr02+Y203) is formed. Zirconium dioxide is crystallized in tetragonal modification. KEYWORDS : hydroxyapatite, zirconium dioxide, aluminium oxide, coprecipitation INTRODUCTIO N The preparation and detailed physical-chemical investigations of new phases based on HA of various dispersity degree and AI2O3, Zr02, Zr02+Y203 and other is one of fundamental problem of HA chemistry and technology. The addition of appropriate additives improves the mechanical characteristics of HA bioceramics without decreasing their biocompatibility. Previously, it was used the solubility method to study interaction in the CaCl2-(NH4)2HP04-NH3-H20 system at 25 C and determined the optimum conditions for obtaining HA. The major factors involved in securing pure HA (free of tricalcium phosphate TCP and other calcium phosphates) are the ratio between the initial components (Ca/P), pH, the time of attainment of equilibrium, and the order of mixing solutions [1]. A high value (^10) favors the coprecipitation of HA and zirconium and yttrium hydroxides [1,2]. In present work interaction is studied by the solubility method (residual concentrations) in following systems at 25 C: CaCl2-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20; CaCl2-YCl3-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20; CaCl2-AlCl3-(NH4)2HP04-NH3-H20. 237
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MATERIAL S AND METHOD S Aqueous solutions of CaCb, AICI3, YCI3, Zr(OH)2Cl2-8H20 (twice recrystallized from water-alcohol solutions of hydrochloric acid), (NH4)2HP04, and ammonia were used as initial components. The solutions were prepared with the use of boiled twice-distilled water. Interaction in the system was studied at a constant ammonium phosphate concentration (0.025 mol/1), and variable amounts of calcium, yttrium, aluminium and zirconium ions. Ratios in the initial mixture were: ni=CaCl2/(NH4)2HP04=1.67^1.75 n2=Zr(OH)2Cl2[(Zr(OH)2Cl2+2YCl3);2AlCl3]/(NH4)2HPO4=0.033^.2 n3=Zr(OH)2Cl2/(Zr(OH)2Cl2+2YCl3)=0.85 and 0.97 n4=2YCl3/(Zr(OH)2Cl2+2YCl3)=0.15 and 0.03 n5=NH4OH/AlCl3=3.0-J-4.0. Systems with Zr(OH)2’^-ions was investigated at pH-lO, which attained by adding to each specimen a concentrated aqueous solution of ammonia. The total volume of the mixtures was 200 ml. The mixtures were vigorously stirred at 25–0.1 C until equilibrium was attained. Then solutions were filtered, the pH was measured and the chemical analysis of the liquid and solid phases were carried out. RESULT S AND DISCUSSIO N I. Systems CaCl2-Zr02Cl2-(NH4)2HP04-NH3-H20 and CaCl2-YCl3-Zr(OH)2Cl211(NH4)2HP04-NH3-H20. To the solution containing CaCl2 and concentrated ammonia were added at first the (NH4)2HP04 and then the Zr(OH)2Cl2 (Zr(OH)2Cl2+YCl3) solutions. In this case the precipitate formed immediately. The equilibrium was reached for 7 days. After the stirring of mixture for one day in liquid phase were detected the considerable amounts of calcium ions; the phosphate and zirconium (zirconium and yttrium) ions practically absent. Further stirring of mixtures for 7 days leads to a gradual quantitative transition of calcium ions into the solid phase. The Ca/P ratio in the solid phase nsoiid=1.67 (pH=9.8) indicates the formation of mixed calcium hydroxyapatite and zirconium (zirconium and yttrium) hydroxide(s) phases. The composition of resulting solid phases is described by the general formulas: Caio(P04)6(OH)2-Zr(OH)4-xH20 Caio(P04)6(OH)2-m{ [Zr(OH)4]o 97[2Y(OH)3]o 03}xH2O Caio(P04)6(OH)2.m{[Zr(OH)4]o85[2Y(OH)3]o,5}-xH20, where m=0.2-^1.2; x=6-12. Heating of hydrated mixed phases of HA and zirconium (yttrium) hydroxide to 900 C, as shown the IRS and XRDA data, leads to the gradual water removing. HA not decomposed and not interacted with the formed zirconium dioxide and yttrium oxide. The HA structure is conserved. Homogeneous mixtures of HA and zirconium dioxide, HA, zirconium dioxide and yttrium oxide with mixing composition formed after calcination at 900 C: Caio(P04)6(OH)2-mZr02 Caio(P04)6(OH)2-m[(Zr02)n(Y203)i-n], where m=0.2^1.2; n=0.85; 0.97.
Coprecipitationof HA With Zr02, Zr02 + Y2O3, Al20sfrom Aquoeus Solutions: V.P. Orlovskii ti al.
4000
2000 1000 Wavenumber (cm’^) Figure. XR D pattern (a) and IR spectrum (b) of 3Ca3(PO)4Al203. Marks : the lines due to AI2O3 (solid sircle)
10
20
30
40
50
60
239
400
X-ray powder diffraction data showed that in all calcinate d solid phases of HA (without the TC P and CaO impurities) and Zr02 in the tetragonal syngony presented . The IR absorption spectra analysis shows that in obtained phases the PO^-ions are strongly distorted, as is indicated by a significant frequenc y split. The spectra show a clearly defined, narrow band of stretching vibrations of the Oir"-groups v(OH>-3570 cm’\ After the heating at 900C the water bending vibrations 5(H2O)=1650 cm"^ disappear, whereas the absorption bands of PO J and OH" remain practically unchanged. The IR absorption spectra of dehydrate d phases are analogous to the IR spectra of HA [3]. II . System CaCl2-AlCl3-(NH4)2HP04-NH3-H20 (25 C). The amphoteric propertie s of Al were considere d in the investigatio n of the system with AICI3 . Ammonia was added in the system in summary amounts neede d for HA formation of the reaction: 10CaCl2+6(NH4)2HPO4+8NH4OH -> iCaio(P04)6(OH)2+20NH4Cl+6H2 0 (CjgH3react~^-^^32 5 mol/1
-
coust)
and
in
agreemen t
with
the
variable
ratio
n5=OH/Al=3.0-^4.0 . With the exceptio n of AIPO 4 formation, alumunium chloride was added in end of precipitation . Interaction is system depends on pH precipitatio n and goes in several studies. 1. At the points of the system with nl=1.67-^1.75 , n2=0.166 , n5=3.0-^3. 2 and pH=7.64-8. 0 at the stirring of initial component s for 14 days (and more) in liquid phase was detecte d a significant quantity of calcium ions; aluminium and phosphate ions practically absent. The Ca/P ratio in the solid phase correspond s to nsoiid=1.5 . In IR absorption spectra of isolated in this region solid phases the stretchin g vibrations of OH~-group v(OH) at 3570 cm"^ is absent. Therefore HA on this precipitatio n study not formed and reaction betwee n CaC b and (NH4)2HP04 proceede d with the TCP formation. Precipitate d hydrated phase consiste d of TCP and A1(0H)3 with the small amount of CI" - ions after the calcination. If calcined at 900 C, this phase formed the mixed phase that consist of TCP and AI2O 3 of the compositio n 3Ca3(P04)2-Al203 . The X-ray powder diffraction data (Fig.a) and IR-spectrum of 3Ca3(P04)2-Al20 3 (Fig.b) are analogous to that of >S-Ca3(P04)2 synthesize d by as [4].
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On the diffraction pattern all peaks characteristics for >S-Ca3(P04)2 are present. Absence of band v(OH) at 3570 cm"^ and splitting of stretching (V3) and bending (V4) vibrations of P04’ tetragon are characteristic to IR-spectra of this phase. The band of P-0 antisymmetric valence vibrations V3 is splitted on some components with a maximum 1116, 1090, 1080, 1038, 1020 cm\ The band of vj vibrations are 969 and 940 cm"\ The band of bending vibrations 0-P-O are V4 - 600, 590, 549, 540 cm’\ 2. At the increasing of precipitation’s pH 8.0l. At pH>9.55 (n5=40) and stirring of initial components for 14 days calcium and phosphate ion were not detected in liquid phase, however aluminium presented in significant amount. Ratio nsoiid=Ca/P=1.67 therefore in solid phase HA presented without the TCP impurity. Aluminium precipitated in form of variable composition’s basic salts. For example at ni=1.67;n2=0.083-H0.2; n5=4.0; pH=9.55 the precipitates of composition Caio(P04)6(OH)2-2[(NH4)x-yAl(OH)3y+x]-zH20, y=0.3-i-1.0 formed in solid phase. If calcined at 900 C,the solid phases had the composition Caio(P04)6(OH)2-yAl203, where y=0.3^1.0. CONCLUSION S Interactions in the CaCl2-Zr(OH)2Cl-(NH4)2HP04-NH3-H20, CaCl2-YCl3-Zr(OH)2Cl2(NH4)2HP04-NH3-H20 systems at 25 C are studied by the method of residual concentrations variant. The investigation of these systems [5, 6] are given the wide information about the optimal conditions of HA and covalent metals hydroxides coprecipitation, the compositions of precipitated phases and the calcination products. The results of these investigations may be used for the development of bioceramics with assigned and determined mechanical properties. REFERENCE S 1. Orlovskii, V.P., Ezhova, Zh.A., Rodicheva, G.V., Koval’,E.M., Sukhanova, G.E., and Tezikova, L.A., Zh.Neorg, Khim.1992, 37, 4, 881. 2. Ezhova, Zh.A., Rodicheva, G.V., Koval’,E.M., and Orlovskii, V.P., Zh,Neorg. Khim. 1991, 36, 10, 2494. 3. Chumaevskii, H.A., Orlovskii, V.P., Ezhova, Zh.A., Minaeva, N.A., Rodicheva, G.V., Steblevskii, A.V., and Sukhanova, G.E., Zh. Neorg. Khim.1992, 37, 6, 1455. 4. Ezhova, Zh.A., Orlovskii, V.P., and Koval’,E.M.,., Zh. Neorg. Khim.1997, 42 (in press) 5. Ezhova, Zh.A., Orlovskii, V.P., Koval’,E.M., and Kozhenkova, E.B., Zh.Neorg. Khim.1996, 41, 11, 1686. 6. Ezhova, Zh.A., Orlovskii, V.P., and Koval’,E.M.,., Zh. Neorg. Khim.1995, 40, 10,1563.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TRANSFORMATION OF a -TCP TO HYDROXYAPATITE IN ORGANIC MEDIA Kiyoko Sakamoto^, Shunro Yamaguchi^, Atsushi Nakahira^, Kazunori Kijima^, and Masayuki Okazaki"^ ^ Department of Chemistry, Osaka Sangyo University, Nakagaito, Daito, Osaka 574, Japan, ^ ISIR, Osaka University, Mihogaoka, Ibaraki, Osaka 567, Japan, ^ Kyoto Institute of Technology, Goshokaidocho, Matsugasaki, Sakyo-ku, Kyoto 606, Japan, "* Osaka University Faculty of Dentistry, Yamadaoka, Suita, Osaka 565, Japan.
ABSTRACT The transformation of ot-tricalcium phosphate (ot -TCP ) to hydroxyapatite ( HAp ) in organic media has been investigated. Hydrolyses of ot-TCP in a series of aliphatic alcohols were carried out under the control of pH and the reaction temperature. The transformation rates and microstructures of HAp were influenced by the hydrophobicity of the aliphatic alcohols. The formation rates of HAp increased with increasing hydrophobicity of alcohols. The rates in 1octanol and 1-hexanol were 5-times faster than that in the hydrophilic alcohols and were compatible to that in the absence of alcohols. The microstructures of HAp prepared in the hydrophobic alcohols were the needle-like particle (length ; l.O’^Z.O U m) and differed from HAp prepared in hydrophilic alcohols and in the absence of alcohol. KEYWORD S ; OL -tricalcium phosphate, hydroxyapatite, transformation, aliphatic alcohols. INTRODUCTIO N There has been considerable interest in transformation of ot-tricalcium phosphates ( a -TCP ) to hydroxyapatite ( HAp ) due to the convenient control of crystal formation. The formation of HAp in the hydration and hardening of a-TCP has been extensively investigated by H. Monma et al. [1,2,3]. While, in organism the formation process of HAp is extremely complication due to concern with organic compounds. It is known that other inorganic phosphates were effectively prepared in organic media [4]. Therefore, the transformation process of OL -TCP to HAp in organic media was examined in detail. In this study, hydrolyses of ot-TCP in a series of aliphatic alcohols were carried out. EXPERIMENTA L PROCEDUR E ot -TCP was provided by Taihei Chemical Industrial Co. Ltd. Hydrolyses of a -TCP in a series of aliphatic alcohols ( ethanol, 1-butanol, 1-hexanol, and 1-octanol ) were carried out ; the mixture of ot-TCP (0.01 mol) and 0.1 M ammonium aqueous solution (36ml) in aliphatic alcohols ( 50ml) was stirring for 2^^120 hours at 70 C. The initial pH value was adjusted to about 11.0 241
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with ammonium solution. The reaction products were filtered off, washed with distilled water, and dried in air at 40 C for 5 hours. The obtained products were identified by X-ray diffractometry ( XRD ; Rigaku Geigerflex RAD I A ). The microstructures of HAp were observed by the scanning electron microscopy (SEM ; Hitachi FESEM H800) and transmission electron microscopy (TEM ; Hitachi H8000, 200kV). RESULT S AND DISCUSSIO N Hydrolyses of a -TCP in each aliphatic alcohols were carried out. Except that of a -TCP in ethanol, the reactions proceeded biphasically. Fig. 1 shows the X-ray diffraction patterns of the products ( reaction time ; 4 hours, reaction temperature ; 70 C ). In this reaction condition otTCP was partially transformed to HAp. Thus, the peaks corresponding to a -TCP and HAp were observed. Based on the relative intensity of the peak for a-TCP and HAp, the rates of transformation were compared. With increasing the hydrophobicity, the rates of hydrolyses
(0
c o
IE
CO
25
30
35
26 / degre e Figure 1. The X-ray diffraction patterns of the products prepared in each aliphatic alcohols (reaction time ; 4 hours, reaction temperature ; 70 C). a ; absence of alcohol, b ; 1octanol, c ; 1-hexanol, d ; 1-butanol, e ; ethanol, ; HAp, O ; a-TCP.
Transformationof a-TCP to Hydroxyapatite in Organic Media: K. Sakamoto et al.
30 /i m
243
1 lim
Figure 2. The SEM (left) and TEM ( right) photographs of a -TCP ( a ), HAp prepared in the absence of alcohol ( b ), and HAp prepared in ethanol ( c ) and 1-octanol ( d ).
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increased as the following sequence ; 1-octanol > 1-hexanol > 1-butanol > ethanol. Even if at 40 C, Q:-TCP were partially transformed to HAp in these solvent systems. In 1-octanol and 1-hexanol, the complete conversions of a-TCP to HAp were achieved for stirring 24 hours at 70 C. The transformation rate in 1-octanol was compatible to that in the alcohol-free system. In 1-butanol ot -TCP was completely transformed to HAp after stirring for 72 hours at 70 C. However, in hydrophilic alcohol (ethanol) the complete conversion was not achieved after 96 hours at 70 C. Fig. 2 shows the SEM and TEM photographs of ot-TCP, and HAp prepared in each aliphatic alcohols. The particles of a-TCP were the smooth surfaces and irregular form. On the other hand, the figures of obtained HAp were a fiber like. The microstructures of HAp were influenced by hydrophobic or hydrophilic of aliphatic alcohols. HAp prepared in ethanol and 1-butanol were as well as that in the absence of alcohols. They were the mixture of platelet-form ( width ; 0.5 U m ) and fine needles ( length ; about 1 M m ). The products in 1-octanol and 1-hexanol were the needle-like particle (length ; 1.0^2.0 U m, width ; 0.1 M m ). Solubilities of the aliphatic alcohols in water at 70 C are as follows : ethanol ; <^, 1-butanol; 7.9 g/lOOg, 1-hexanol ; 0.6 g/lOOg, 1-octanol ; insoluble. The similar rate of hydrolysis in 1octanol to that in water appeared that a-TCP might be hydrolyzed in water phase and that 1octanol did not affected the hydrolytic process. However, the microstructure of obtained HAp in 1octanol was considerably different from that in water and similar to that by hydrothermal reaction. It is noticed that long needle-like HAp should be prepared under mild reaction condition. CONCLUSIO N The transformation rates and microstructures of HAp were influenced by the hydrophobicity of the aliphatic alcohols. In the hydrophobic aliphatic alcohols such as 1-octanol and 1-hexanol, hydrolysis of ot-TCP were effectively carried out, furthermore, the obtained HAp were the uniform needle-like particle ( length ; 1.0^2.0 U m, width ; 0.1 U m ). These resuhs imply that hydrophobicity of organic compounds might control the hydroxyapatite-formation in biosystems. ACKNOWLEDGEMEN T The authors thank Taihei Chemical Industrial Co. Ltd. for providing ot-TCP.
REFERENCES 1. 2. 3. 4.
Monma, H., and Kanazawa, T., Yogyo-Kyokai-Shi,1976, 84, 209-213. Monma, H., Ueno, S., and Tsutsumi, M., Gypsum&Lime, 1978,156, 6-11. Monma, H., Goto, M., and Khomura,T., Gypsum&Lime, 1984,188,1 1-16. Sakamoto, K., Tsuhako, M., Murakami, M., and Tanaka, K., Nippon Kagaku Kaishi, 1995, 1995, 681-688.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
STRUCTUR E AND SOLVATIO N EFFECT S OF P04^, HPO42, H2PO4- AND H3PO4 FRO M AM I AND PM 3 A. J. Salinas^ A. Serretl, M. Vallet-Regi^ andL.L. Hench^ 1. Departamento de Quimica Inorganica y Bioinorganica; Facultad de Farmacia; Universidad Complutense; E-28040 Madrid; Spain. 2. Deptartment of Materials. Imperial College of Science, Technology and Medicine. Prince Consort Road. London SW7 2BP. England.
ABSTRAC T To understand the role of phosphate ions in interfacial reactions that occur on the surface of bioactive materials, the structure and solvation effects of P04^", YiPO^’ and H2P04" were modelled using the Molecular Orbital semi-empirical methods AMI and PM3. For comparison, H3PO4 was also studied. Models were made for both the isolated species and with water as simulate solvent, allowing the study of the influence of solvent on geometries and ionic charges: partial charges on eveiy atom of phosphate ions, which could play an important role in the formation of the calcium phosphate layer on the siuface of bioactive materials, were calculated. The most stable structures were obtained with AMI in the presence of water. KEYWORDS : MO modelling, phosphate ions, AMI, PM3, solvation effects. INTRODUCTIO N Certain compositions of glasses and glass-ceramics have the ability to bond with living tissues [1]. A conunon characteristic of such bioactive materials is the formation of a calcium phosphate rich layer on their surface when exposed to physiological solutions [2]. The reaction layer is initially amorphous calcium phosphate (a-CaP) which, after nucleation and growth, crystallises to hydroxycarbonate apatite (HCA). In bioactive glasses and glass-ceramics, a high surface area silica gel layer is formed by partial network dissolution and surface polycondesation reactions. The silica gel and the HCA layers provide adsorption sites for cellular growth factors generated by macrophages and stem cells. The final result is the formation of a bond between the tissues and the bioactive ceramics. With the objective to study the interfacial reactions on the surface of bioactive materials, in this paper, ions P04^", VS?0/^’ and H2P04’» that can be involved in formation of the a-CaP and HCA layers, were modelled. For comparison H3PO4 was also studied. Since classical molecular mechanics methods are not suitable to calculate reaction pathways and ab initio quantum mechanical methods require very high CPU time [3], we have used the semi-empirical methods AMI (Austin Model 1) [4] and PM3 (Parametric Method 3) [5] for calculations. EXPERIMENTA L DETAIL S All calculations were made using CAChe Worksystem [6], 3.8 release, running on a Power Macintosh which offers an interface for several semi-empirical methods such as MOP AC a 245
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Bioceramics Volume10
molecular orbital (MO) package optimized for the study of chemical structures and reactions. We have used MOPAC’s semi-empirical methods AMI and PM3 which provides two approximations for the solution of SchrGdinger’s equation. The process for geometry optimization was: the chemical species were first generated in CAChe Editor, locating the charge of the ions on the oxygen atoms, and then valence, configuration and geometry were beautified. Atomic positions obtained were the starting point for the AMI and PM3 geometry optimizations. To understand the influence of an aqueous medium, MO modelling was made for the isolated species (gas-phase) and with water as simulate solvent. RESULT S AND DISCUSSIO N Because of their importance in biological systems, the structure and properties of phosphorus-oxygen ions are the subject of this investigation. Properties like heats of formation, bond lengths and partial charges were calculated. The optimum geometry obtained by AMI and PM3 is the spatial disposition of atoms for which a minimum in the calculated value of heat of formation is reached. In Table 1 are shown the heats of formation obtained by AMI and PM3 and the experimental valuesfi*omliterature. Table 1. Heats of formation (Kcal/mol) of PO43", HPO42-, H2P04- and H3PO4. AM I PM 3 gas-phase solvent: water Experimental* solvent: water gas-phase
P04^HPO42H2PO4H3PO4
44.08 -205.04 -312.92 -287.13
-587.25 -513.13 -429.68 -334.47
79.34 -172.50 -290.95 -269.72
* Experimental values from literature for aqueous ions [7] and phosphoric acid [8],
-533.92 -467.91 -393.38 -306.80
-305.60 -309.12 -310.12 -306.2
For every specie, in the same conditions, the lower value of the heat of formation (higher stability) is obtained by AMI. The presence of water as simulate solvent of the envirorunent of the species increases the stability of ions and molecules. The maximum decrease of the heat of formation due to the presence of water, -600 Kcal/mol, is obtained for the phosphate ions where the charge is -3. The solvent stabilization effect decrease with the decreasing of the charge. So, for the neutral phosphoric acid the solvation effect is only --50 Kcal/mol. For H3PO4, a good agreement between experimental and calculated values of heat of formation is obtained. For the phosphate ions with water, theoretical values are lower than the experimental. In Figure 1 are depicted the structures at their lowest-energy state, obtained for AMI and PM3 after geometry optimisation using a simulated solvent (water) and without any solvent (gas). The bond lengths are included in Table 2. The distances P - O obtained by AMI are always lower than those obtained by PM3. However, the AMI and PM3 O - H distances are almost identical. The influence of water as solvent is as follows. For ion P04^’ stabilization due to the molecules of water produces a decrease in the four equal bond lengths. For the ion Y{PO/^’^water influence can be detected mainly by shortening of the P(l) - 0(4), that is the largest of the four P - O distances due to the influence of H(6) bonded to 0(4). In the ion H2P04’ there are two groups of P - O lengths depending whether the oxygen atom is bonded or not to a hydrogen. For the first group, P(l) - 0(4) and P(l) - 0(5), the bond length decrease for the influence of solvent. For the second group, P(l) - 0(2) and P(l) - 0(3), the bond length is near that of a double bond and increases when water molecules are present. Finally, in phosphoric acid there is only one oxygen that is not bonded to a hydrogen. This oxygen is double bonded to the phosphorous and for this
Structure andSolvation Effectsof PO4’-, HPO4^-.HiPO^ andH3PO4:A.J. Salinas et al. 247
gas-phas *
AMI
solvwit: walar
gat*phat «
PM3
solvant : watar
P043 -
HP042 -
H2PO4-
H3PO4
Figure 1.- Plan view of AMI and PM3 optimized species. Table 2.- Bond lengths (Angstrom) of P04’-, WO^; H2PO4-, H3PO4 obtained in gas phase and using water as simulate solvent by PM3 and AMI. AM I PM 3 POd-*P ( l ) - 0 ( 2) P ( l ) . 0 ( 3) P ( l ) . 0 ( 4) P(1)-CX5) HPOd^ P ( l ) - 0 ( 2) P ( l ) . 0 ( 3) P ( l ) . 0 ( 4) P ( l ) - 0 ( 5) 0(4).H(6) H2PO4P ( l ) - 0 ( 2) P(l)-0(3) P(l)-0(4) P ( l ) . 0 ( 5) 0(4)-H(6) 0(5)-H(7) H1PO 4 P ( l ) . 0 ( 2) P ( l ) . 0 ( 3) P ( l ) . 0 ( 4) P ( l ) - 0 ( 5) 0(3).H(6) 0(4).H(7) Q(3)-H(8)
gas-phase
solvent: water
gas-phase
solvent: water
1.556 1.555 1.555 1.556
1.531 1.530 1.530 1.530
1.662 1.663 1.663 1.663
1.612 1.614 1.613 1.613
1.509 1.515 1.702 1.509 0.948
1.511 1.510 1.639 1.510 0.954
1.575 1.585 1.822 1.575 0.943
1.575 1.575 1.745 1.575 0.945
1.479 1.475 1.643 1.631 0.946 0.950
1.491 1.491 1.612 1.610 0.958 0.958
1.504 1.504 1.744 1.744 0.943 0.943
1.528 1.529 1.712 1.712 0.946 0.946
1.451 1.592 1.590 1.584 0.953 0.954 0.956
1.474 1.580 1.582 1.584 0.962 0.962 0.963
1.450 1.677 1.679 1.660 0.943 0.944 0.944
1.487 1.667 1.666 1.666 0.949 0.948 0.948
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reason is the shortest bond length. The influence of water molecules in the environment of H3PO4 is veiy small and only can be detected as a weak increase of the double bond distance. The influence of the solvent in the O - H bond lengths is very weak and, in all cases, only a very small increase of the calculated distances can be detected. The MO optimization allows one to obtain the partial charges on every atom. This aspect may be important to explain the bonding of phosphate ions at the surface of bioactive materials. When water is used as simulated solvent, higher values of the partial charge values are obtained in all cases. With AMI the partial charge of phosphorousatom ranges from +2.74 (P04^’) to +2.86 (H3PO4) respectively. However, PM3 yields lower values of the charge of phosphorus ranging from +1.88 to +2.36. As expected, the charge on the oxygenatoms is strongly influenced by the presence or not of hydrogen bonded to the oxygen atoms. So, for the P04^" group all the charges on the oxygen are -1.43 for AMI and -1.22 for PM3. For the other species the charge ranges from -1.38 to -1.27 (AMI) and from -1.17 to -1.06 (PM3) for the oxygens that are not bonded to a hydrogen, and from -0.90 to -0.86 (AMI) andfrom-0.78 to -0.72 (PM3) for oxygens bonded to hydrogen atoms. Finally, charge on the hydrogenatoms ranges from +0.27 to +0.33 (AMI) and from +0.24 to +0.29 (PM3). CONCLUSION S MO modelling, using the semi-empirical methods AMI and PM3, is a useful tool for understanding the interfacial reactions in the siuface of bioactive materials where phosphate ions are involved. For chemical species modelled P04^", W^0/^\H2P04" and H3PO4, the most stable structures are obtained by AMI, and when water is used like simulate solvent. The presence of water in the environment exerts a weak influence in the O - H bond lengths. However, water has a large effect on the P - O distances. In most cases where water was used like solvent, the distances of the double bonds P - O decrease while the single bonds distances increase. Finally, the data obtained for partial charges on every atom of P04^", HPO42", H2PO4" and H3PO4 are presented. However, additional theoretical and experimental work must be done before the role of P ions on bioactive bonding is understood. ACKNOWLEDGEMENT S We thank Dr. J. West of the University of Florida for helpful discussions. CICYT for financial support through research projects MAT94-0424-C02-01 and MAT96-0919 is also acknowledged. One of the authors (A.S.) thamks NATO Scientific Comitee for a grant. REFERENCE S 1. L.L.Hench, R.J.Splinter, W.C.Allen, T.K.Greenlee, J. BiomedMat. Res.,1971, 2 117-141. 2. Lany L. Hench, J. Am. Ceram. Soc.,1991, 74 [7] 1487-510 . 3. K.B. Lipkowitz and D.B. Boyd, (Eds.), Reviews in Computational Chemistry.VCH Publishers, New York, 1990. 4. M. Dewar., Zoebisch, Eve.G. Healy, E.F. Stewart, J.J.P." JAm.Chem.Soc, 1985,107, 39023909. 5. J.J.P. Stewart, J.Comput.Chem.,1989,10, pp. 209 ibid,pp 221. 6. Cache Worksystem, Version 3.8, CaChe Scientific Inc. Oregon. 7. D.D.Wagman W.H.Evans, V.B. Parker, R.H. Schum, I. Halow, S.M. Bailey, K.L. Chumey, R.L. Nuttall, J. Phys. Chem.Ref.Data,1982, Vol 11, Suppl 2. 8. J.C. Bailar, HEmeleus, R. Nyholm, A.F. Trotman-Dickenson (Eds.) ComprensiveInorganic Chemistry,Vol. 2, Pergamon Press, Oxford, 1973.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
THE DETAILE D CONFIGURATIO N OF CARBONAT E ION S STRUCTUR E DETERMINE D BY POLARIZE D IR SPECTROSCOP Y
IN
APATIT E
Y. Suetsugu^ L Shimoya^ and J. Tanaka* 1
National Institute for Research in Inorganic Materials, 1-1 Namiki, Tsukuba, Ibaraki 305, Japan. ^ School of Science and Engineering, Waseda University, 3-4-1 Okubo, Shinjuku, Tokyo 160, Japan. ABSTRAC T In order to elucidate the configuration of carbonate ions in apatite structure, carbonate apatite single crystals were grown by a CaCOa flux method. Carbonate ions could substitute both OH (A-site) and PO4 (B-site); especially A-sites were completely occupied by carbonate ions. The space group of the crystal obtained was determined to be P^ (hexagonal) using a four-circle X-ray diffractometer, and cell parameters were a=0.948nm and c=0.689nm. From the angular dependence of polarized IR spectra for out-of-plane and in-plane bending vibrations of CO3 ions, it was indicated that the triangular plane of CO3 ion was parallel to the c-axis for A-site substitution and perpendicular to the c-axis for B-site substitution. KEYWORD S Carbonate apatite, X-ray diffraction, space group, polarized IR spectroscopy. INTRODUCTIO N For the comprehension of inorganic component in biological hard tissue and the development of new artificial bone materials, it is important to elucidate the crystal structure of carbonate apatite (CAp). Carbonate ions can replace both OH (A-site) and PO4 (B-site) in hydroxyapatite structure [1]. However, the detailed structure of CAp has been unknown since no single crystal of CAp with enough size for direct structure analyses has been prepared. In this paper, single crystals of CAp were grown by afluxmethod and the configuration of carbonate ions was determined. EXPERIMENTA L PROCEDUR E Mixed dry powders of calcium carbonate and a-tricalcium phosphate (a-TCP) were sealed in a platinum capsule; CaCOs worked as a flux whose content was 40wt%. The capsule was heated up to \SO(fC under Ar gas pressure of 50MPa. The growth of single crystals was performed by lowering the temperature by St/hour. Residual CaCOa flux was eliminated using 10wt% of EDTA-Na2 aqueous solution. The crystals obtained were analyzed by an electron probe microanalyzer (JEOL JXA-8600MX) and a carbon analyzer (Horiba EMIA 511) which was equipped with an IR detector for the quantitative analysis of CO2 gas desorbed from samples. Xray diffraction was measured by a four-circle X-ray diffractometer (Rigaku AFC-5R) with AgKa (40kV, 180mA, monochrometer: HOPG, detector slit width: 1/2fi). FT-IR spectra were measured by Spectrum 2000 (Perkin Elmer) equipped with a microscope. A single crystal was put for its caxis to be perpendicular to the direction of incident beam and the angular dependence of reflected IR spectra was measured using polarized beam. 249
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Figure 1. Intensity of X-ray diffraction scanned along AOl, 0*1 and -hhl. Extremely strong peaks are removed from data, corresponding to gaps. RESULT S AN D DISCUSSIO N The single crystals grown were hexagonal prismatic and long along the c-axis; their typical size was 0.2 X 0.2 X I m m ^ From chemical analyses it was found that the crystals had the formula of Ca9 8[(P04)5 6(C03)o.4](C03)i 0 in which A-sites were completely occupied by carbonate ions and about 7% of B-sites were substituted by carbonate ions. In order to determine the space group of carbonate apatite, the X-ray diffraction intensity change was measured along AOl, 0*1 and -hhl on the first layer in a reciprocal lattice by a four-circle diflfractometer; the results are shown in Fig. 1. N o diffraction peak was found at A, AP=1/2, 3/2, 5/2 and 7/2, such peaks should be observed when the space group is Pb (monoclinic) as was reported for A-type of CAp on the basis of powder X-ray diffraction [2]. However the present crystal did not have monoclinic lattice but hexagonal lattice since no peak was observed at half-integral reciprocal point. ouu-
003
1
500 -
<
^-v
§^:
B 80C 3
8 60**-^ w c
^-
S
20 -
B
0-
d
001 p
#w
007
005 1
’
r
4
1
r
COL Figure 2. Intensity of X-ray diffraction scanned along c*-axis. removed from data, corresponding to gaps.
Extremely strong peaks are
Detailed Configurationof CarbonateIons in Apatite Structure: Y. Suetsuguet al.
A-type
Ik
^
S
...
251
-type
^-
^
900
890
880
860
870
850
Frequency /cm"^ Figure 3. Angular dependenc e of polarized IR absorption by V2 (out-of-plane ) vibration of CO3. Solid line: electric vector is parallel to c-axis. Broken line: electric vector is 45 to c-axis. Dotted line: electric vector is perpendicula r to c-axis. Figure 2 shows an intensity variation along c*-axis. Diffraction peaks were apparently observed at 001, 003, 005 and 007. This extinctio n rule, the appearance of 00/ (/=2n+l) diffraction, indicates that the crystal has no screw axis. Taking a fundamenta l apatite structure into account, the space group of carbonate apatite is therefor e ascribed to P6. Cell parameter s were determine d to be a = 0.948n m and c = 0.689nm . Figure 3 shows the polarized IR absorption ascribed to V2 (out-of-plane ) vibration of carbonate ion; this figure was converte d from reflecte d spectra using Kramers-Kronig relations. Figure 4 indicates the similar IR spectra for V3 (in-plane) vibration of carbonate ion. The solid lines, broken lines and the dotted lines correspond to that the electric vector of incident polarized IR beam is parallel, 45 and perpendicula r to the c-axis of the crystal, respectively .
B-type
A-type
\j
f\
B-type
A-type
s
V
Xi
<
1700
1650
1600
1550
1500
Frequency /cm-’
1450
1400
1350
Figure 4. Angular dependenc e of polarized IR absorbance by V3 (in-plane) vibration of CO3. Solid line: electric vector is parallel to c-axis. Broken line: electric vector is 45fi to c-axis. Dotted line: electric vector is perpendicula r to c-axis.
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A peak at 882cm"^ in Fig.3 is due to V2 normal mode of A-site substituted CO3 and a peak at 873cm’^ to that of B-site substituted CO3. The V2 mode, corresponding to the out-of-plane vibration of CO3, should have a maximum absorption when the electric vector of polarized IR beam is perpendicular to the triangular plane of CO3 ion. On the other hand, peaks at 1450 and 1540cm"^ in Fig.4 are ascribed to V3 normal mode of Asite CO3 and peaks at 1410 and 1470cm’^ to that of B-site CO3. As the V3 mode is due to in-plane vibration of CO3, this mode has a maximum absorption when the electric vector is parallel to the CO3 triangular plane. The V2 mode is originally degenerate for higher symmetry field but split into two peaks in the present system. As regards the A-site substituted carbonate ion, the V2 mode has a maximum absorption when the electric vector of polarized beam is perpendicular to the c-axis and no absorption when the electric vector is parallel to the c-axis, as indicated in Fig.3. This result means that the triangular plane of CO3 is parallel to the c-axis for A-site substitution, though dichroic ratio [3] is not taken into account. TTiis conclusion is consistent to Fig.4; i.e. the V3 absorption decreases with increase of angle between electric vector and the c-axis. Such configuration of CO3 is not in accord with the previous work [4]. As shown in Figs.3 and 4, IR spectra for B-site CO3 ion show a different behaviorfi*omthat for A-site CO3 ion. Since the intensity variation of V2 mode for B-site is opposite to that for A-site, it is considered that CO3 ion in B-site is perpendicular to the c-axis. The highest symmetry in qjatite structure is P63/m (hexagonal) observed for fluorapatite. Elliott [2] reported that the space group of A-type of CAp was Pb (monoclinic) on the basis of powder X-ray difltraction study, in which the symmetry reduction from the hexagonal system to the monoclinic system was attributed to the periodicity doubled along b-axis. The present result is differentfi*omhis result; this is plausibly ascribed to the disorder of A-site substituted CO3 ion. According to the X-ray and polarized IR analyses, the present CAp, space group P6, has a six-fold axis of rotatory inversion on which the A-site substituted CO3 ion is located for its triangular plane to be parallel to the c-axis. Thus, CO3 ion substituting A-site can take three equivalent configurations around a six-fold axis of rotatory inversion. Suppose the three configurations are randomly occupied, such disordered distribution of CO3 ion causes an apparent ’six-fold axis of rotatory inversion’ by statistically averaging. Because of the existence of six-fold axis of rotatory inversion, one C-0 bond of CO3 ion should lie on a mirror plane and thus one 0-0 side of the triangular CO3 plane is parallel to the c-axis. SUMMAR Y Single crystals of carbonate apatite were grown by a flux method using CaC03. The configuration of carbonate ions in ^atite structure was elucidated by X-ray diffraction and polarized IR measurements. A carbonate ion substituted A-site for its triangular plane to be parallel to the c-axis. It was considered that one C-0 bond of CO3 lain on a mirror plane and could turn to three equivalent directions.
REFERENCES
1. Bond, G. and Montel, G., Comp, Rend.Acad ScL (Paris), 1964,258, 923-926. 2. Elliott, J.C, J, Appl Crysl, 1980,13, 618-621. 3. Fraser, R.D.B., in A Laboratory Manual of Analytical Method of Protein Chemistry, Vol. 2, Oxford, Pergamon Press, 1960,285-351. 4. Elliott, J.C, in Proceedings of International Symposium on Structural Properties of Hydroxyapatite and Related Compounds, Gaithersburg, 1968, Ch. 11.
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TISSU E CULTUR E ON AMORPHOU S CALCIU M PHOSPHAT E COATIN G Kazuya Suzuki, Yoshihiko Kageyama, Yusuke Kita, Atsuto Yoshino, Kozo Matsushita, and Tadashi Kokubo* First Department of Surgery, Hamamatsu University School of Medicine, 3600 H^da-cho, Hamamatsu, 431-31 J£q)an, Tel: 053-435-2276 , Fax : 053-435-227 2 Division of Material Chemistry, Faculty of Engineering, Kyoto University
ABSTRAC T ACP-coating technique was applied to artificial tracheae and tissue culture in our previous study because AC P was stable in handling and did not induce inflammatory reaction in soft tissue. Effect of ACP-coating on tissue culture was studied biochemicall y and pathologicall y in this study. Small fragments of blood vessel and fat tissue that were placed on ACP-coated culture plates attache d well, and produced more t-PA and PAI- 1 comparing with them on non-coatin g plates. Histocompatibilit y betwee n AC P and fibroblasts or collagen fibersseems to play a role in these results. ACP-coating is very useful because tissue culture technique is important in various medical studies. INTRODUCTIO N The adherenc e of cells to a material surface is a prerequisit e for tissue culture. Our fundamenta l study showed that soft tissues attached to amorphous calcium phosphate (ACP ) coated on polymers much better than non-coatin g or collagen-coatin g polymers [1-3]. Furthermore, viability of tissues cultured on ACP-coating were better than that on non-coating , judging from histopathologica l findings [4]. In ord^ to evaluate the biochemica l activity of cultured tissues; blood vessel and adipose tissue, comparative study was performed using ACP-coating and noncoating culture plates in this study. MATERIAL S AN D METHOD S AC P coating on the surface of culture plates was paformed according to the biomimetic process reported by Kokubo [5,6]. In this process, a biomimetic solution with higho- concentration s of calcium and phosphate ions were employed . Temperature of solution was maintained lower than 35 degree centigrade , and drying process was done quickly [4] (Fig. 1,2). Tissue fragments of fat and blood vessels (1 or 2 nmi in diameter) with whole layer were placed on the surface of culture plates, and incubated in culture medium, RPMI164 0 with 10 % FBS. PAI- 1 (plasminogen activator inhibitor, type 1) secrete d by fat tissue, and t-PA (tissue plasminogen activate^*) produced by endotheliu m of blood vessels were measured by enzyme immunoassay. For quantitative evaluation of viability, the levels of PAI- 1 and t-PA in culture medium were compared betwee n ACP-coating and non-coating groups, with or without stimulation. 253
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Fig. 1 SEM photograph of ACP-coating
Fig. 2 Photograph of a culture plate Left: ACP-coating Right non-coatin g
RESULT S The coating layer was present without exfoliatio n during incubation period. Soft tissues attached well to the ACP-coating layer just after the placement , on the other hand, soft tissue could not attach to the non-coatin g surface (Fig. 3). Scanning electro n microscopic (SEM ) observatio n showed that soft tissue including fibroblasts adhered to ACP-coating layer by collagen fibers growing around AC P (Fig. 4). Fat tissues in ACP-coating group, adhered to the surface of plates, and secrete d 51.1 – 18.5 ng/lOOmg (/day) of PAI- 1 after 2 weeks incubation. On the other hand, almost all fat tissues in non-coating group werefloatin g in culture medium, and secrete d 21.0 – 8.9 ng/lOOmg (/day) of PAI-1 . T-PA production by fragments of blood vessels were 4.85 – 2.3 ng/lOOmg (/day) in ACP-coating group, and 1.23 – 1.5 ng/lOOmg (/day) in non-coatin g group 2 weeks after incubation. By the stimulation using vitamin A and C, level of t-PA in culture medium increase d much more in ACP-coating group (Fig. 5,6). The degree of tissue degeneratio n judging from histopathologica l finding was higher in non-coatin g group than in ACP-coating.
Fig. 3 Photograph of plates during culture. Left: fat tissue attache d to ACP-coating plate. Right: fat tissue could not attache d to a non-coatin g plate.
Tissue Culture on Amorphous Calcium Phosphate Coating: K. Suzuki et al. 255
Fig. 4 SEM photogr^h of the surface of ACP-coating duringtissueculture. Left: fibroblast Right: soft tissue and collagen fibers. DISCUSSIO N AND CONCLUSIO N Soft tissue dose not attach to the surface of a non-coating culture plate in general, therefore, tissue culture is more difficult than cell culture in which only the cells attached to the surface can proliferate [4]. Our previous pathological studies revealed that ACP did not induce an inflammatory reaction, showed strong adherence in soft tissues [1-3], and soft tissue cultured on ACP-coating showedtight-attachmentand high viability [4]. In this study, PAI-I secreted by fat tissue and t-PA produced by endothelium were measured for quantitative evaluation of viability. Both level of PAI-I and t-PA in culture medium was high in ACP-coating group. These results indicated that cultured tissue attached to ACP-coating was more stable biologically than that floating in non-coating plates. Although the mechanism oftissueadhesion to ACP is unknown, fibroblast and collagenfibersseems to play a role in this reaction. ACP coating was not suitable for a long period culture because coating layer obstructed light transmission during microscopic observation. ACP-coating technique is useful for a short-period tissue culture which is important in various medical studies. Further investigation will be necessary to clarify the mechanism oftissueadhesion. (%) (%) 200
Withoutstimulatio n
600
L
With stimulatio n by vitann m A and C
I :ACP-coating
y
:ACP-coating
>:Non-coatin g
L
#:Non-coatin g
J
Vit.A&C
(mean – SEM) 1
Day 0-5
Day 5-10
Day 10-1 5
1
(meant SEM)
,1 .
Day 4-6
1
1
1
Fig. 5 Change of the t-PA, produced by cultured endothelium with or without stimulation by vitamin A and C.
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(%)
h
100
50
Day 5
DaylO
Day 1 5
Fig. 6 Change of the PAI- 1 produced by cultured fat tissue, with or without stimulation by LPS .
REFERENCES 1.
2. 3. 4. 5. 6.
Suzuki, K., Kageyama, Y., Yokoyama Y., Harada, Y. and Kokubo, T. Bioceramics Volume 7, Butterworth-Heinemann , Oxford 1994,113-11 7 Kageyama, Y., Yokoyama Y., Suzuki, K., Harada, Y. and Kokubo, T. Bioceramics Volume 7, Butterworth-Heinemann , Oxford 1994,165-17 0 Suzuki, K., Kageyama, Y., Toyoda, F., Nogimura, H., Harada, Y. and Kokubo, T. Bioceramics Volume 8, Elsevier Science, Oxford 1995,351-35 6 Suzuki, K., Kageyama, Y., Kita, Y., Ohi, S., Yoshino, A., Matsushita, K., Harada, Y. and Kokubo T. Bioceramics Volume 9, Pergamon, Oxford 1996,411-41 4 Kokubo, T., Hata, K., Nakamura, T. and Yamamoto T. Bioceramics Volume 4, Butterworth-Heinemann , Oxford 1991,113-12 0 Kokubo, T. Biomaterials, 1992,12,15 5
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BONELIK E APATIT E LAYE R FORME D ON ORGANI C POLYMER S BY BIOMIMETI C PROCESS : TEM-ED X OBSERVATIO N OF INITLU . STAG E OF APATIT E FORMATIO N H.Takadama\ F. Miyaji\ T. Kokubo^ and T. Nakamura^ ^Department of Material Chemistry, Faculty of Engineering, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan, ^Department of Orthopedic Surgery, Faculty of Medicine, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan
ABSTRAC T The mechanism of the apatite formation induced by silicate ions on a collodion film was examined by TEM-EDX. The TEM-EDX showed that the collodion film is first attached with the Si within 6 hours, then with the Ca withinfiirther6 hours andfinallywith the P within further 12 hours. The product at the final stage was identified as an apatite by electron diffraction. These results indicate that the silicate ion which is dissolved from the CaO, Si02-based glass and attached on the surface of the polymer induces the formation of the apatite through the formation of a certain calcium silicate. KEYWORDS : Apatite, Biomimetic process, TEM, EDX, Nucleation, Simulated body fluid INTRODUCTIO N It was previously shown by the present authors that a dense and uniform bonelike apatite layer can be formed on surfaces of various organic polymers at the normal temperture and pressure by a biomimetic process [1], where a polymer substrate is first placed on granular particles of CaO, Si02-based glass soaked in a simulated body fluid (SBF) [2] in order to induce apatite nucleation on the surface of the substrate, and then soaked in another solution highly supersaturated with respect to the apatite in order to grow the apatite nuclei by consuming the calcium and phosphate ions from the surrounding fluid. The mechanism how apatite nuclei are formed on the substrate in SBF has not been, however, clarified yet. In the present study, the mechanism of the apatite formation which is induced on a collodionfilmby sihcate ions dissolved from the CaO, Si02-based glass was examined using a transmission electron microscope (TEM) equipped with a energy dispersive X-ray spectrometer (EDX). MATERIAL S AND METHOD S A glass named G of the nominal composition MgO 4.6, CaO 44.7, Si02 34.0, P2O5 16.2 and CaF2 0.5 wt%, which is the parent glass of high strength bioactive glass-ceramic A-W [3], was used as the CaO-Si02-based glass for the apatite nucleation. A mixture of the reagent grade chemicals of the corresponding composition was put into a 50 ml platinum crucible and melted at 1450 C for 1 h in a MoSi2 furnace. The melt was poured onto a stainless steel plate, pressed into a plate 1 mm thick, and annealed at 600 C Thus-obtained glass was cut into rectangular 257
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specimens 10 X 10 X1 mm in size, polished wilh 3-4 iimcj) diamond paste, and washed with pure acetone in an ultrasonic cleaner. A 200-mesh Ti-grid (3.0 mm(|)) which was covered with an organic collodion film, and then deposited with carbon was used as the substrate. The Ti-grids were placed on the above glass plate taking a distance of 0.5 mm in 20 ml of SBF with ion concentrations (Na^ 142.0, K^ 5.0, Ca^^ 2.5, Mg^^ 1.5, HCOs" 4.2,Cr 148.0, HP04’ 1.0, SO/- 0.5 mM) nearly equal to human blood plasma [2] and pH 7.40 at 36.5 C for various periods until 96 hours. After removing from the SBF, the films were washed with distilled water, dried at room temperature, and observed under a TEM (JEM-2000FX III, JEOL, Co., Tokyo, Japan). The elmental analysis was also carried out for the product formed on the film by an EDX (VOYAGER III, NORAN Instruments, Inc., Middleton, U.S.A.). RESULT S AND DISCUSSIO N Figure 1 shows the EDX spectrum of an original Ti-mesh covered with collodion and carbon films. The C and O, and Ti elements correspond to the carbon, collodion film, and the Ti grid, respectively. Figure 2 shows the EDX spectrum of the Ti-mesh after soaking in SBF with glass G for 6 hours. The C, O, Ti and newly Si elements were detected. This indicates that silicate ions which were dissolved from the glass were first attached to the surface of the collodion film. Figure 3 shows the TEM image and the EDX spectrum of the Ti-mesh after soaking in SBF with glass G for 12 hours. The C, O, Ti, Si and newly Ca elements were detected. Another principal component of apatite, P, was not detected at this stage. This indicates that not both the Ca and P but only Ca is secondly combined with the silicate ions, and a certain calcium silicate is formed
o
PL^ i
o
Ti Ti
iM-
0 1 2 3 4 5 6 7 8 9 10 Energy / ke V Fig. 1 EDX spectrum of an original Ti-mesh.
0
iX. .k 1 2
i.
4 5 6 7 8 9 10 Energy / ke V
Fig. 2 EDX spectrum of a Ti-mesh after soaking in SBF with glass G for 6 hours.
Bonelike Apatite Layer Formed on Organic Polymers by BiomimeticProcess: H. Takadama et al.
i^W:
0)
0 1 2 3 4 5 6 7 8 9 Energy / ke V Element
Ca
P
0
Na
Mg
CI
Si
Atom%
1.27
0.03
92,56
0.00
0.16
0.16
5.83
Fig. 3 EDX spectrum of a Ti-mesh after soaking in SBF with glass G for 12 hours. O
Ca
Energy / keV Element
Ca
\_kXomVo 4.32
P
0
Na
Mg
CI
3.42
91.08
0.00
0.00
1.09
Fig. 4 EDX spectrum of a Ti-mesh after soaking in SBF with glass G for 24 hours.
Si
O.O9J
10
259
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0 1 2 3 4 5 6 7 8 9 Energy / ke V Element
Ca
P
O
Na
Mg
CI
Atom%
21.69
16.16
51.72
5.53
1.63
3.27
10
Si 0.00 J
Fig. 5 EDX spectrum of a Ti-mesh after soaking in SBF with glass G for 96 hours.
on the surface of the collodion film. The dark smal dots in the TEM image may correspond to such a product. Figure 4 shows the TEM image and the EDX spectrum of the Ti-mesh after soaking in SBF with glass G for 24 hours. The C, O, Ti, Si, Ca and newly P elements were detected. This indicates that P is finally combined with the calcium ions to form apatite on the surface of the collodion film. The product seen on the left side of the photograph may correspond to the apatite. Figure 5 shows the TEM image and the EDX spectrum of the Ti-mesh after soaking in SBF with glass G for 96 hours. The product at the this stage, which is seen as an aggregate of needle-like crystals in the TEM image, was identified as an apatite by electron diffraction and showed Ca/P atomic ratio of 1.34. In conclusion, the silicate ions which are dissolved from the CaO, Si02-based glass and attached on the surface of the polymer induces the formation of the apatite through the formation of a certain calcium silicate. REFERENCE S 1. Tanahashi, M., Yao, T., Kokubo, T., Minoda, M., Miyamoto, T., Nakamura, T. and Yamamuro, T., J. Am. Ceram. Soc. 1994, 77, 2805-2808 2. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T. and Yamamuro, T., J. Biomed.Mater.Res. 1990,24,721-734 3. Kokubo, T., Shigematsu, M., Nagashima, Y., Tashiro, M., Nakamura, T., Yamamuro, T. and Higashi, S., Bull. Inst. Chem.Res., Kyoto Univ. 1982, 60, 260-268
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
SINTERABILIT Y AND SECON D PHAS E FORMATIO N OF SYNTHETI C HYDROX Y APATIT E : CONTROLLIN G PARAMETER S AND EFFEC T ON BOND STRENGT H Hyuk-Joon Youn, Hyun Seung Ryu, Kug Sun Hong, Sung Soo Chung\ and Choon Ki Lee^ Colledge of Materials Science & Engineering, Seoul National University, Seoul 151-742, Korea ^Orthopaedic Surgery, SamSung Medical Center, Kangnam-ku, Seoul 135-230, Korea ^Department of Orthopaedic Surgery, Colledge of Medicine, Seoul National University, Chongno-ku, Seoul 110-744, Korea ABSTRAC T The effect of starting Ca/P ratio and aging temperature in hydroxy apatite(HA) synthesis by wet chemical processing were investigated using SEM, XRD and EPMA. The sintering and grain growth behavior were strongly dependent on starting Ca/P ratio and aging temperature. The types and amount of second phase(Ca3(P04)2(TCP) and Ca4P209(TetraCP)) formed during sintering were also correlated with starting Ca/P ratio. The presence of a second phase and its apparent inhibition to densification of HA was discussed. The necessary conditions for good bonding between bone and HA can be suggested. KEYWORD S : Hydroxy apatite, Ca/P Ratio, Aging, TCP, Second phase. Sintering, Bonding INTRODUCTIO N Hydroxy apatite(HA, Caio(P04)6(OH)2), a well known bone minerals, has been studied as a prime candidate for bone replacement applications. Various processing routes for synthetic HA ceramics have been studied, including precipitation, sol-gel, hydrothermal, etc. [1-5] In order to obtain proper HA as implant materials the essencial parameters which determine sintering behavior, strength and bio-activity of HA have been investigated. [7,8] They were reported as a functon of Ca/P ratio, microstructure, second phase formation and so on. Ca/P ratio, which is strongly dependent on processing conditions and has a large effect on HA characteristics, is one of most important parameters. [9-11] In spite of abundant studies about this parameter, complete understanding for its roles and control has not been fulfilled yet. In this study the effect of Ca/P ratio on sintering behavior and on bonding behavior as well as the dependence of Ca/P ratio on processing were examined through synthesis of HAs with a wide range of Ca/P ratios. EXPERIMENTA L PROCEDUR E Ca(N03)24H20(EP, Junsei) and (NH4)2HP04(EP, Junsei) were dissolved in deionized water to prepare 0.5 M Ca-stock solution and 0.3M P-stock solution. The pH was adjusted by adding IN NH4OH. The two solutions were mixed under vigorous stirring at appropriate ratios yielding Ca/P atomic ratios of 1.3, 1.5, 1.67, 1.83 and 2.0. The precipitates started to form just afler mixing and the slurry was aged for 24 h at one of three different temperatures : 30, 60 and 90 C. The aged precipitates were filtered and washed to eliminate unreacted species and then dried overnight at 110 C. Dried powders were calcined at 600 C for Ih. The dried and calcined powders were determined to be single phase, hydroxy apatite regardless of starting Ca/P ratios. Calcined powders were pressed into pellets, 8mm in diameter by uniaxial pressing and then sintered for 2h at various temperatures 1200 - 1350 C in air. The bulk densities, grain size, the phases and 261
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compositions of sintered specimens were analyzed by Archimedes method, by linear intercept method, by XRD(X-ray diffractometer, McSci) and by EPMA(Electron probe microanalyzer, Model Jeol), respectively. In vivo tests of bonding between sintered specimens and bone(rabbit) were carried out and after 2 - 8 weeks tensile strengths for various samples were measured. RESULT S AND DISCUSSIO N In Table 1 the type and relative amounts(relative intensities in XRD) of second phase are summarized for the hydroxy apatite heat-treated at various temperatures. While only single phase hydroxy apatite is observed for calcined powders at 600 C, additional phases such as aTCP(tricalcium phosphate) can be found, dependening on sintering temperature, starting Ca/P ratio and aging temperature. This implies that the second phases were formed during sintering process. The tendency of second phase formation can be summarized as follows : The lower Ca/P ratio, the higher concentration of second phases. Higher Ca/P ratios lowered the concentration of second phases. Figure 1 (a) - (e) shows the dependence of sintered density on aging temperature for various starting Ca/P ratios. With increment of aging temperature distinguishable behaviors can be found for different Ca/P ratios. The density/aging temperature dependence changes with Ca/P ratio. At Table 1. The type and amount(relative peak intensity) of second phases for various starting Ca/P ratio, aging temperature and sintering temperature ((a): a-TCP, (T): Tetra-CP(Ca4P209)) Ca/P = 2.0 Ca/P =1.3 Ca/P =1.67 30 C 60 C 90 C 30 C 60 C 90 C 30 C 60 C 90 C 1200 C 0.40(a) 0.12(a) 0.30(a) 0.02(a) 0.04(a) 0.03(a) 0.0 0.0 0.0 0.0 0.40(T) 0.0 1350 C 0.43(a) 0.46(a) 0.37(a) 0.05(a) 0.04(a) 0.06(a)
30 60 90 Temperature (*t)
30 60 90 TanpCTature Ot)
j-n-ml200 j-O-mlZSO hA-ranoo |-r7-ml35 0
Ca T
1.83
I
i 1 ^
30 60 90 Ten^CTature (X)
Figure 1. The bulk densities as a function aging temperatures for various starting Ca/P ratio and sintering temperature.
Sinterability and Second Phase Formation of SyntheticHydroxy Apatite:H-J. Youn et al.
3.0 h
2.8
Ca/Pl.3
(a)
-A
U26
uV^
2 6H
’
-cP -
(b)
^
- -sec -A-pc c
w
LCa/Pl.6 7
mT h-l2 I 4I i 6I
2 4bdLjLljJ_ L 2 4 6 8 10 Grain Size()ain)
I- o o
(c)
263
95.1 - I 88.7
-n-3o c
-A-po-c L Ca/P2. 0
i t X I Irl I I i I i
8 10 Grain Size((im)
-D-so’c H g2. 4 -A-Po’c
2 4 6 8 10 Grain Size(^m)
nJ 76. 0
Figure 2. The relatioship betwee n bulk density and grain size for various Ca/P ratio and aging temperature , (right Y axis: relative density %) . Table 2. The composition(Ca/ P atomic ratio) of produced hydroxy apatite for various starting Ca/P ratio and aging temperatur e (EPM A analysis) 2.0 1.67 1.83 1.3 Starting Ca/P atomic ratio 1.5 aging temperatur e ( C) 30 C 90 C 30 C 90 C 30 C 90 C 30 C 90 C 30 C 90 C 1.52 1.60 1.58 1.64 1.65 1.69 1.68 1.72 1.68 1.74 Fianl Ca/P atomic ratio low Ca/P ratios (1.3 - 1.5), density appears to increase with aging temperature ; at intermediat e Ca/P ratios (1.67 ~ 1.83), density appears constant with aging temperature ; and at Ca/P ~ 2.0,densit y appears to have parabolic dependenc e on aging temperature . Generally, it can also be found that the lower sintered density is achieve d for lower Ca/P ratios and lower aging temperatures , with higher densitie s for higher Ca/P ratios. On the basis of above results, it can be concluded that the formation of a second phase inhibits densificatio n of hydroxy apatite. Figure 2 shows the relationshi p betwee n sintered densitie s and grain sizes for various Ca/P ratios and aging temperatures . Eventhough all cases show a saturation in sintered density with increasing grain size, there is a strong dependenc e on aging temperatur e depending on Ca/P ratio. For example in the case of Ca/P ratio = 1.67, there is no dependenc e on aging temperature , however, in the case of Ca/P ratio = 1.3, there is a large dependenc e on aging temperature . This dependenc e on Ca/P ratio is very similar to the dependenc e of sintered density on aging temperature . In other words grain growth behavoir can be grouped into 3 cases, similar to sintering behavior. Thus it can be also conclude d that second phase formation has a significant effect on grain growth in hydroxy apatite. The possibility that precipitatio n of a second phase suppresse s the grain growth by grain boundary pinning can be excluded , since ultimate grain size has no dependenc e on the amount of second phase as can be seen in figure 2. For instance the case of Ca/P ratio = 1.3, which had the largest concentratio n of second phase does not show the smallest grain size. Thus it can be assumed that the evolution of second phase during sintering inhibits densification . The major second phase such as TCP(Ca3(P04)2) can be formed by the decompositio n of hydroxy apatite phase(Caio(P04)6(OH)2 ) or by the reaction betwee n adsorbed species(iuireacte d species : Ca^^, PO4", etc). Further discussion on compositio n and second phase will be followed to confirm this. Table 2 compares starting Ca/P ratios to those in hydroxy apatite. From this, a linear dependenc y of final compositio n on starting Ca/P ratio can be found, even though final Ca/P
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Table 3 The bonding strength betwee n bone and sintered hydroxy apatite with preparation condition (sintering condition of HA : 1300 C, 2h, unit of bonding strength : Kgf) Period of In-vivo test Ca/Pl.5, 90 C Ca/Pl.5, 3 0T Ca/P 1.83, 30 C Ca/P 1.67, 30 C 2 weeks non-bonding non-bonding non-bonding non-bonding 4 weeks 2.54 2.12 2.25 3.94 4.01 2.34 2.50 2.08 8 weeks ratios deviate from the Ca/P ratios of starting solution. The final Ca/P ratios also show the dependenc y on aging temperatur e : higher aging temperatur e results in higher Ca/P ratios. Consequently , hydroy apatite ceramics with Ca/P ratio of 1.5 - 1.75 can be obtained from starting Ca/P ratio of 1.3 ~ 2.0. Considering the amount and type of second phase with composition , it can be deduced that evolution of second phase such as TCP is due to decompositio n of hydroxy apatite and that this tendenc y depends on Ca/P ratio of produced hydroxy apatite. Thus it can be concluded that the formation of large amounts of second phase during sintering, associate d with low starting Ca/P ratios, causes lower densificaton . Higher densificatio n is achieve d through higher starting Ca/P ratios, and is attribute d to increase d HA stability. However the fact that in the case of Ca/P ratio =1.3, aging at 90 C, sintered densitie s remained relatively high in spite of formation of large concentration s of TCP, cannot be easily explaine d now. In Table 3 the results for in-vivo test are summarized. The bond strengths shows the highest values for starting Ca/P ratio = 1.67, 30 C aging and lowest for starting Ca/P ratio = 1.5, 30 C aging. This implies that there is an optimum concentratio n of second phases which maintains relatively high densities , as well as yields the best strength betwee n bone and synthetic hydroxy apatite. These necessar y conditions can be obtained by control of starting Ca/P ratio and aging temperatur e during powder preparation process. SUMMAR Y Hydroxy apatite with various Ca/P ratios were synthesize d by controlling starting solution and aging temperature . While single HA phase was obtained for powders dried and calcined at 600 C, additional phases such as TCP and TetraCP were formed during sintering at 1200 - 1350 C, depending on starting Ca/P ratio and aging temperature . Lower sintered density was observed for lower Ca/P ratio. This could be explaine d by the inhibition to densificatio n due to evolution of second phases. Although the highest density and lowest second phase concentratio n were obtained for higher starting Ca/P ratios, the optimum Ca/P ratio(=1.67 ) led to some second phase formation, high sintered density and best bond strength.
REFERENCES
1. P. E. Wang and T.K.Chaki, J. MaterScL Mater. Mater. Med.,4 (1993), 150-15 8 2. M.Akao, H.Aoki, K.Kato, J. Mater.Sci.,\6(1981), 809-81 2 3. T.Futagami and T. Okamoto, Yogyo Kyokai-Shi ,95 (1987), 775 4. T.Hattori, Y. Iwadate, T. Kato, J. Mater.Set.Lett., 8 (1989), 305-30 6 5. T. Hattori, Y. Iwadate, J. Am.Ceram. Soc, 73 (1990), 1803-180 5 6. A.J.Ruys, Wei, C.C.Sorrell, M.R.Dickinson, and B.K.Milthorpe, Biomaterials, 16 (1995), 406 7. R.I. Martin, and P.W. Brown, J. Mater.Sci.Mater. Med.,6 (1995), 138-14 3 8. D. De With, H.J.A. Van Dijk, N. Hattu and K.Projs, J. Mater.Sci.,16 (1981), 1592-159 8 9. S.R. Radin and P.Ducheyen , J.Biomed.Mater.Res., 27 (1993), 35-45 10. K.Ishikawa, P.Ducheyne and S. Radin, ikid,4 (1993), 165-16 8 11. K. de Groot, Biomaterials, 1 (1980), 47-50
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
POROU S SOL-GE L BIOGLASS fi FRO M NEAR-EQUILIBRIU
M DRYIN G
Jipin Zhong and David C. Greenspan U.S.Biomaterials Corporation, One Progress Blvd. #23 Alachua, FL 32615, USA
ABSTRAC T It has recently been reported that sol-gel derived Bioglassfi with large pores exhibits a rapid induction of hydroxy-carbonate apatite (HCA) and significant degradability which results in enhanced bone growth[l,2]. In this work, large pore Bioglassfi has been developed through the sol-gel route with drying under near-equilibrium(liquid/gas) conditions. The water vapor in the near-equilibrium condition creates a pressure inside the pores of the drying gel. This counteracts the capillary pressure responsible for pore deformation which results in a larger pore network. The near-equilibrium drying method to achieve large pores is less expensive, safer and easier for production than supercritical drying without the difficulties in removing residual organics associated with surface modified gels. In addition, crack-free monolithic implants from sol-gel derived Bioglassfi can also be fabricated by drying under near-equilibrium conditions. This solgel derived Bioglassfi could be an ideal candidate for bone grafting applications. KEYWORDS : Bioglassfi, Sol-gel, Bone growth, Degradability, Near-equilibrium drying. INTRODUCTIO N It has recently been reported that sol-gel Bioglassfi with large pores exhibits a rapid induction of hydroxy-carbonate apatite (HCA) and significant degradability[1,2]. Bone growth has been enhanced by the rapid induction of HCA, the mineral phase of bone, on the gel while the gel degrades[2]. It seems that large pore diameter and high surface area are responsible for the fast nucleation and crystallization[3,4] of HCA and the significant degradation rates observed. A number of methods have been documented to make large pore gels. Drying under supercritical conditions[5] and surface modification by chemical additives have been used to make areogels with large pores[6]. Other methods to make large pore gels include using hydrofluoric acid (HF) and base treatment. In this work, large pore Bioglassfi particles and monoliths have been developed through the sol-gel method with drying under near-equilibrium conditions in water. The near-equilibrium drying method to achieve large pores is less expensive, safer and easier for production than supercritical drying without the difficulties in removing residual organics associated with surface modified gels. MATERIAL S AND METHOD S The compositions of sol-gel Bioglassfi used in this work were 58S, 68S and 77S which has been reported elsewhere[7]. The gels were prepared by mixing D.I. water, HCl, TEGS 265
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(tetraethoxysilane, SiCOCzHs) ), TEP (triethylphosphate, OP(OC2H5)3) and CN (calcium nitrate, Ca(N03)2»4H20) in order. It took 2 hours to complete the mixing and the mixture was then cast into a polypropylene mold for aging at 60 C for about 60 hours. After aging, the pore liquor was removed and the gels were transferred to drying vessels which were placed to a drying chamber with water. The chamber was closed and placed into an oven, thus, the gels were dried under a near-equilibrium condition in waterfi^omroom temperature to 130 C for 3 days. The dried gels were than heated to 700 C over a 3 day period and prepared for analysis. The pore texture of the samples was evaluated using an Autosorb 6 nitrogen absorption/desorption analysis system. The formation of HCA on the gels’ surface when reacted in a simulated body fluid (SBF)[8] was analyzed using FTIR to evaluate the bioactivity. RESULT S AND DISCUSSIO N Fig. 1 shows representative samples of the sol-gel Bioglassfi materials made by the nearequilibrium processmg method. The different pore structures can be seen in Table I for the drying under near-equilibrium conditions versus ambient conditions as described above. Fig. 2 shows the FTIR reflection spectra of two samples: 77S(A) and 77S(B) with different pore sizes. Clearly, HCA formed more rapidly on the 77S(B) with large pores than 77S(A) with small pores in SBF. Generally, gels are networks of small colloid particles. The networks include voids which become pores and pore channels in the final gels. The capillary pressure due to the pores and pore channels is: P = 2YCOS0/r (where, y is liquid tension, 0 is contact angle and r is the radius of pores and pore channels). The presure will pull the network tight enabling pore collapse and gel shrinkage as liquid evaporates during drying.
Figure 1. Sol-gel Bioglassfi Samples
Porous Sol-Gel Bioglassfi from Near-EquilibriumDrying: J. Zhong and D.C. Greenspan 267
Table I. Pore Texture of Sol-gel Bioglassfi Samples Sample
Surface Area (MVg )
Ave. Pore Diameter (A)
1
ID
Ambient Dry(A)
Near Equil. Dry (B)
Ambiem Dry(A)
Near Equil. Dry (B)
58S
68
86
289
207
68S
50
65
326
305
77S
24*
40
241*
389
(* This gel begins to sinter above 600C and pore size is 30 A and SA is 43 IM^/g when heate d at 600 C) The phase diagram of water in figure 3 shows that above the critical point in supercritica l conditions, there is no differenc e betwee n liquid and gas state and the capillary pressure become s zero. Drying under this condition, the gel undergoes very minimal shrinkage resulting in large pore size and pore volume. In our designed drying method, a drying chamber and a regular oven were used and the condition inside the chamber was near the equilibrium line in Fig. 3. During drying, the water vapor inside the chamber would be allowed to evaporate out at a low rate so that the gel is dried under a high vapor pressure. It is believe d that the moisture of the near-equilibriu m drying step enhance s the reaction at the neck betwee n two particles in the network strengthenin g the neck and "back bone" of the gel structure. In addition, the establishe d vapor pressure inside the channels and pores at elevate d 4 77S(B ) unreacte d
10 77S(A) wnreacte d % R
% R
% R
5
2 1
2 1
j V_J ^ 77S(A) SBF2 4 hours
Q\
K
J V^ 77S(A)SBF2day 8 h [\
/
V
Si-O-Si
^^^ 1
100 0 Wavenumbers (cm-1 )
77S(A) Fig. 2
% R
A A
2 77S(B ) SBF2 4 hours
% R
2 77S(B ) SBF 2 Days
% R
’
Si-o-s i
2 100 0 Wavenumbers (cm-1 )
77S(B)
FTIR spectra of 77S sol-gel Bioglass from normal (A) and near-equilibriu m drying (B) reacted in SBF at various times
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Critical Point
250 ^
1 3
C/5
200 150
Liquid
Gas
100 50 0
^T
100
1
1
1
1
200
300
400
500
Temperature ( C) Fig. 3 Phase Diagram of Water
temperature of the near-equilibrium drying counteracts the capillary pressure. All of these would resist the shrinkage and collapse of the gel structure and result in large pore sizes and crack-free monoliths. CONCLUSION S The large pore sol-gel Bioglassfi particles and monoliths of 58S, 68S and 77S were developed through sol-gel route using drying under near-equilibrium conditions. The large pore structure and resulting crack-free monoliths were achieved by water vapor pressure to counteract the capillary pressure and enhance the reaction between the necks of cross-linked colloids to strengthen the back bone of the gel. The rapid mduction of HCA formation with a significant degradability[9] could make the large pore Bioglassfi ideal implants for bone grafting. REFERENC E 1. Greenspan, D.C. et al., "The in vitrodegradability of sol-gel Bioglassfi", Presented at Society for Biomaterials annual meeting, April 30-May 4, 1997, New Orleans, USA. 2. Wheeler, D.L., et al., "/« vivo evalution of sol-gel Bioglassfi, Part I: Histological Findings", Presented at Society for Biomaterials annual meeting, April 30-May 4, 1997, New Orleans, USA. 3. Pereira, M.M. et al, J. Am. Ceram. Soc, 1995, 78[9], 2463-68. 4. Greenspan, D.C. et al., Bioceramics 8, 1995, 477-82. 5. Tewari, P.H. et al., US patent #4,610,863, 1986. 6. Deshpande, R. et al., US patent #5,565,142, 1996 7. Li, R. et al., J. Appl. Biomater., 1991, 2, 231-39. 8. Kokubo, T. et al., J. Biomed. Mater. Res., 1990, 24, 721-34. 9. Greenspan, D.C. et al., "The evaluation of degradability of molten verse sol-gel derived Bioglassfi" Bioceramics 10, 1997.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CERAMI C - CERAMI C BEARIN G SYSTEM S COMPARE D ON DIFFEREN T TESTIN G CONFIGURATIONS . J. Chevalier, B. Cales, J.M. Drouin, Y. Stefani. N(»lon Desmarquest Fine C^-amics, ZI n**l, 27025 Evieux Cedex, France
ABSTRAC T Coamic - ceramic systems w»:e evaluated with different wear testsfiromthe simple Pin on Disk to the more sophisticated Disk on Disk and Hip Simulator tests. It appears that alumina sliding against zirconia gives the best results, suggesting this combination couki be an excellent candidate for hip prosthesis i^kations. KEYWORD S zirconia, alumina, wear, hip INTRODUCTIO N Ultra-high molecular weight polyethylene (UHMWPE) cups articulating against metal heads has been the choice of orthopedk bearing materials in total joint lei^acements for many years. However, it becomes nK)re and more evident that the wear (tf UHMWPE is one of the major fact(H^ that limit joint {vosthesis lifetime [1,2]. Today, their is a growing intact for ’hard / haid’ pairing such as metal-on-metal or ceranuc against ceramk:, in particular foxyoung and active patients. Metal articulating against metal was introduced in the past and seva:al conflicting results were obtained [3,4,5]. One major conc^n is the dissolution of metallic ions in the body [3]. Alumina against alumina has been used for more than tw^ty years and gives excellent clinical results [6]. Because of their improved mechanical propmies [7], yttria stabilized zirconia (Y-TZP) hip joint heads must be considered as potential candklales for ceranuc / ceramk: systems. Our goal was thertfoie to investigate cosmic matmals facing zirconia. Zirconia against zirconia and zirconia against alumina pairings woe therefore tested and compared with alumina / alumina couplings. MATERIAL S AND METHOD S Experiments were conducted on surgical grade zirconia (PROZYRfi) and alumina ceramics. All materials were Hot Isostically Pressed after sintering in order to reach full density. Mean grain size calculated by the linear intercqH method is O.S ^m and 2 ^m for zirconia and alumina respectively. Zirconia against zirconia (Z-Z), alumina against alumina (A-A) and zirconia against alumina (Z-A) combinations were tested on different testing devices: (i) Pin on Disk (POD) testing was first investigated in order to give a first evaluation of the prcqx)sed pairings. POD tests were conducted in water and in bovine serum to investigate the 271
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influence of lubricant properties on wear results. Testing conditions are summarized on Table (1). Wear was measur^ via the wear scar on the pin. (ii) Disk on Disk (DOD) tests were then conducted in a synthetic serum (Plasmionfi, Rhdne Poulenc Rogo: Bellon) to approach the real in vivo wear condition. Plamionfi contains 25 g/1 proteins which seems close to the body situation. Testing conditions are summarized on Table (2). (iii) Hip joint simulatCM- tests are still in [Hogress and only results for zirconia heads articulating against alumina acetabular cups will be presented aft^ 3 millions cycles. Tests are performed on a MTS hip simulator under physilogical loading (J. Paul loading, maximum load 2.S kN) in Plasmionfi. EXPERIMENTA L RESULTS AND DISCUSSION POD results are summarized on Table (1) in terms of wear rate and friction coefficient Tests in bovine serum w^ie stopped before that in wat^ because of bactmal degradation. It dsppeiaxs that zirconia against alumina gives the bestresults(lower wear rate) for both lubricants. Z-Z gives contradictory results in bovine serum and in wat^ : on the one hand (short tom, bovine serum lubricant) Z-Z combination gives wear rates two times high^ than the two otho* combinations, on the other hand (long term, watCT lubricant) Z-Z shows interesting behavior, bettCT than A-A. Also of impcHtance is the dS&x of lubricant on the measured friction coeffici^ts, much lower when bovine serum is used. This stresses the ^ect of lutaiflcation conditions, abeady pointed out fcH* ceramic or metal articulating against polyethylene. It was shown by several authors that protein content played a key role on wearresults,in particular on polyethylene transfm to the surfEice of its bearing antagonist In-vivo situation corresponds roughly to a protein content of about 25 to 35 g/1, in comparison with 0 g/1 and 70 g/1 for water and bovine soimi respectively. Th^efore, water should be avoided and bovine sown diluted to obtain 30 g/1 protein content Bovine serum is often used in biomaterials wear studies. Howevo*, it is v ^ sensitive to bacterial d^radation and we preferred a synthetic serum Plasmion iot the folk>wing of the study because it contents 25 g/1 proteins and is very resistant to bacterial degradation.
V^^^i 1
a) b)
Lubricant Water Bovine S^um
Sliding distance 500 m 150 km
Pin Alumina Disk Zirconia 1 Test (a) Wear rate 9 . 1 0’ Water (mniVm) Friction coeff. LI 1 Test(b) Wear rate 1 Bovine serum (mmVm) Friction coeff. 0.1
Applied toad 5N 5N
Sliding speed 1 5 cm/s 1 5cm/s _|
Zirconia Alumina 3.10*
Alumina Alumina 5.10’
Zirconia Zirconia 3.10*
1.0 7.I0*
0.6 8.10*
0.8 1.10-^
0.08
0.1
0.2
Table (1): conditions for the POD test and experimental results
1 1
1
Ceramic-CeramicBearing Systems on Different Testing Configurations:J. Chevalieret al.
1 Disk on Disk
Plasmionfi
1 Combination 1 Wear rate (mmVm) 1 Frictioncoeff
25 km
alumina - zirconia 1.10* 1
Sliding spe^ 1 5 cm/s 1
Applied pressuie 0.15 MPa
alumina - alumina 2.10’ 0,2
273
zirconia - ziiconia 1 in prosress 1 in progress 1
Table (2): DOD testing conditions and experimental results. Following these results, ZrA and A-A were tested on a DOD configuration in Plasmionfi. results are presented on table (2) also in terms of wear rate and friction coefficient Again Z-A coupling gives the lowest wear rate, wh^eas friction coefficient is high in comparison to A-A coiq)ling. For this testing arrangement, which simulates better in-vivo conditions, Z-A leads again to the best wear behavior. A set of experiments is currently conducted on a MTS hip-wear simulator. Zirconia heads again alumina cups are tested first because they may rq)resent the ^^ypropriate choice for c^amic ceramic pairing. Figure (1) rq)resents weight loss of both femoral head and acetabular cup after 3 million cycles compared to that obtained on PE cups articulating against zirconia or metal. Wear debris are expected to be much lowerwith the coamic - ceramic system. Friction confident obsCTved with Z-A pairing and rqx)rted in Table (3) is exceptionnaly low, three to six times lower than measured with PE cups against zirconia or metal respectively. Table (4) gives the roughness of both heads and cups forthe Z-A combination before and afto* wear test as well as monoclinic content to insure phase transfcmnation is not occuring during wear. No damage is obs^ved neith^ on the cup nor on the head after three million cycles. Of importance is the fact that no transformation is observed during wear of zirconia, in contradiction with previous works where severe damage was observed with POD [8] and hip-wear simulator [9] tests und^ water for Z-Z pairing. The explanation for extensive wear was attributed to tetragonal to monoclinic (t-m) phase transformation caused by an increase of temperature at the intoface between the two bearing surfaces. In a recent study, PROZYRfi zirconia heads wereshown to be poorly sensitive to t-m transformation in water [10]. This, in combination with lulHication conditions, may be the key fcxthe obs^ved good behavior of the presently studied zirconia. I 1
Combination Friction coeff.
zirconia - alumina 0.0015
zirconia - polyethylene 0.005
metal - polyethylene 1
am
1
Table (3): Values of factional torques measured fcM: different pairings on hip simulator.
1 1
alumina cup zirconiahead
roughness befcne wear test 5 nm 2nm
roughness after 3 initial monoclinic million cycles content (%) 5 nm <2% 2 nm
monoclinic 1 content after test!
<2%
1
Table (4): Values of roughness before and after wear test and monoclinic content measurements on the zirconia femoral head.
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weight loss (mg) 10
: melal - UHMWP E rZiOj-UHMWP E O : ZrOi - AI2O3
1 0 0 0 0 00
2 0 0 0 0 00
300000 0
Cycles t hip joint configuration s Figure (1): Weight loss as a function of number of cycles for differen studied with hip simulator CONCLUSIO N Wear behavior of differen t coamic - co-amic bearing systems was investigated . It appears a zirconia head bearing against an alumina cup could be an outstandin g solution for oith(^)aedi c applications.Howevo*, hip-wear simulator studies must be conducte d for longer periods, up to 10 millions cycles to conclude definitel y on the potentia l use oi this system. REFERENCE S 1 : A. Wang et al.. Wear,1995,181-183 , p. 241. 2 : P. Campbell et al.. Complications orthop., 1991,6 , p. 116. 3 : J. Black, Clinical Orthopaedics andrelated research, 1996, n^ 329S, p S244. 4 : J. Jacobs et al.. Clinical Orthopaedics andrelated research, 1996, n"" 329S, p. S256. 5 : F.W. Chan et al.. Clinical Orthopaedics and related research, 1996, ii’’333, p. 96. 6: L. Sedel et al.. Clinical Orthopaedics andrelated research, 1994, n’’298, p. 175. 7 : J.M. Drouin et al.. Journal cfBiomed. Mater, Research, 1997, 8 : Willmann et al., Biomaterials, 1996,17 , p. 2157. 9 : Oonishi et al., Bioceramics 9,1996,p.503. 10: J. Chevalier et al, Bioceramics 10,Ibid.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
DISSOLUTIO N AND MECHANICA L BEHAVIOU R OF CERAMI C COATING S FOR ORTHOPAEDI C APPLICATION S
PLASMA-SPRAYE D
N. Demonet*, P. Benaben*, J. L. Aurelle**, B. Forest* and J. Rieu* * Ecole des Mines de Saint-Etienne, Centre S.M.S., Departement de Mecanique des Biomateriaux et Traitements de Surface, 158 Cours Fauriel, 42023 St-Etienne Cedex 2, France ** S.E.R.F., 85 Chemin des Bruyeres, 69153 Decines Cedex, France ABSTRAC T Plasma-sprayed coatings of alumina (AI2O3), hydroxyapatite (HAP, Caio(P04)6(OH)2) and AI2O3/HAP duplex on Ti-6A1-4V substrates confer rather good properties to cementless prostheses: corrosion and fretting wear resistance, high superficial porosity for bone anchorage. During plasma-spraying, HAP is partly transformed into several other calcium phosphates which have a high solubility at low pH in simulated body fluids. The chemical and morphological alterations caused by the dissolution are followed by gamma spectrometry, SEM and X-rays with variable pH, stirring and temperature. Their influence is shown on the calcium dissolution kinetics. Such coatings are highly defective in their microstructure. Thanks to pull-off tests and interfacial indentation, we clear off that the resistance to damage increases with the AI2O3 content in the coating; alumina improves the inner cohesion and the mechanical adhesion with the metal below. The influence of post plasma thermal treatments is drawn too. KEYWORD S coating - ceramic - plasma-spraying projection - dissolution - radioactive tracers - adhesion INTRODUCTIO N To avoid the problems linked to the degradation of the polymeric cement of implanted prostheses, cementless ones have been developed. It is necessary to coat the surface of their stem with a high porous ceramic to prevent bone/metal contacts and to facilitate a natural bone anchorage (figure 1). After more than a 20-year use, alumina coatings are, now, well known for their chemical inertia and their good mechanical adhesion. More recently, calcium phosphate components have appeared. One of the most interesting CaP is the hydroxyapatite (HAP) which presents a chemical and crystallographic likeness to the bone structure. This property allows a natural and strong bonding between the bone and the stem. But, it was reported that plasma-spraying changes the degree of cristallinity as well as the phases composition of HAP coatings [1]. In vivo, these transformations favour the osseous regeneration. But in vitro, the coating dissolution increases at low pH. Dissolution tests are necessary to understand better the influence of the different parameters and the mechanisms. Besides the usual attempts, we have used an original method based on radioactive tracers which allows an accurate observation of the calcium dissolution kinetic in accordance with the pH, stirring and temperature in Ringer’s solution. Moreover, we were interested in the metal/ceramic interface. The adherence is the main property of control which depends strongly on the plasma-spraying parameters. 275
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Gas mixtures Voltage Current strength Projection distance Powders
Acetabular cup
Ceramic coated metallic stem Table 1. parameters
Some
Ar/H2 or Ar/He 70 V 450 A 85 or 120 mm Al203orHAPor AI2O3/HAP of
plasma-spraying
Figure 1. Total hip prosthesis This paper first reports the materials and the methods. Then, dissolution tests show the influence of experimental conditions. Finally, the mechanical tests demonstrate the improvement brought by alumina in the mechanical adherence. MATERIAL S AND METHOD S The plasma-spraying projection was realised by the S.E.R.F. on an entirely self-acting APS apparatus (atmospheric plasma system - Sulzer Metco). The parameters are given in table 1. Platelets of Ti-6A1-4V were employed as substrates. They were sandblasted with corundum at a pressure of 6 bars, single face coated, then, cut in 1 cm^ square blocks. The samples were cleaned and dried in an oven. Some of them were exposed to a thermal flux of neutrons (1.3 10^^ neutrons/cm^) in order to obtain the radioactive isotope "^^Ca (T1/2 = 4.54 days. Eg = 488.9 and 1296.8 keV) in the coating. The irradiations were done by the laboratory Pierre Sue (Centre d’Etudes Nucleaires - Saclay 46 France). The reaction Ca > 47.Ca lasts 20 minutes. At the end of variable duration of inmiersion in 30 ml of Ringer’s solution with different pH, radioactive activity of "^^Ca in the coating and in the host solution was measured by gamma spectrometry. We observed, acciwately, the calcium dissolution kinetics. A lot of mechanical tests were described in the literature [2]. But, few of them are easy to realize, to render and reproduce. Among them, we have chosen: the pull-off tests standardized by AFNOR [3] performed with cylinders (figure 2) and the interfacial indentation made with square blocks whose sides are polished to see easily the prints during the indentation. Opposcfl cyUndcr | 4 ^|
J Opposed fyiiiidcr
Tl^AM V (1)
HA P
Figure 2. Samples for mechanical adhesion tests: (1) sandblasted and cleaned cylinder, (2) cylinder coated with HAP, (3) glued coated cylinder, (4) entire assembly
Dissolution and Mechanical Behaviour of Plasma-Sprayed Ceramic Coatings: N. Demonetet al. Caldtnn qaaaatlty sttll oi tiie soM(mg) andfreefcithe Uquid (mg ki 30ml)
277
Cflldnm quantity In the host mlntlon (mg bi 30 ml)
10 8
’«3>-^, /1 / / /
-
~j
V"
’
PH 3
49 4 2
t :yo.^
Ol^ L
,
,
,
50
47 ,
Figure 3. Evolution of the ^’Ca quantity on the solid sample and in the host solution (pH 3, no stirring, room temperature)
pH6 PH 7 pH8 1
, 200
,
H
300
1
i
350
400
Ttane (hours)
Figure 4. Calcium dissolution in Ringer’s solution under different pH (no stirring, room temperature)
RESULT S AND DISCUSSIO N 1. Dissolution behaviour Firstly, we have tested the feasibility of the method. No chemical and morphological alteration caused by the irradiation has been noted by SEM and X-rays. After irradiation, dissolution tests were performed: the quantity of Ca on the solid sample decreases whereas the radioactivity in the liquid increases proportionally (figure 3). As an example, the dissolution under different pH (2 to 8) at room temperature was measured (figure 4). The group of curves show the strong influence of the pH on the kinetics. For the pH 2 and 2.5, after 2 or 3 days, the metal is completely bare with a total disappearance of coating. On the contrary, the more the pH increases, the more the dissolution slows down. At pH 8, the threshold of the spectrometer is not reached: the "^^Ca quantity is too low in the liquid. On each curve, a plateau appears: the dissolution phenomenon decreases. But a saturation doesn’t happen because any precipitation is observed on our samples. It is well known that plasma-spraying process strongly affects HAP coatings. Even though the CaP powders are rather stable, the coatings aren’t [4]. The plasma-spraying modifies the cristallinity, the specific surface and mainly the composition phases: with HAP, new CaP appear and some deshydroxylations happen. Moreover, the coating structure is favorable to a dissolution because of its lamellar aspect with many cracks, pores where some similar conditions to crevice corrosion are created (low pH, few or no fluid movement): the acid attacks are favoured. It is proved too, that some ions (for example Na^, CO3 ", in the Ringer’s solution) increase the solubility of HAP coatings [5]. The pH influence and some mechanisms (like the PO4 " ion complexations) [6] begin to be proposed for all these observations. Thermal treatment 500 C
|
Distance without 900 C (mm) 85 11.2 9.5 26.5 100 10.4 18.7 34.2 120 8.9 17.7 30.1 Table 2. Adherence (MPa) of HAP coating with the projection distance (mm) and post plasma-spraying treatment (3h30)
Eoating composition ’ HAP alone 25%Al203/75%HAP 50%Al2O3/50%HAP 75%Al203/25%HAP Alumina alone
Adherence (MPa) 11 11.6 12 12.6 15
Table 3. Adherence (MPa) according to the AI2O3 content in the coatings
278
Bioceramics Volume10 Ln<M> - (tt in HB>) 4,1
^
3,0 3,7 3,5
p^’
3,3 3,1 2.9 2,7 2.5
(
4,5
1
4.
5
1 S,5
1 6
1
0«^«Al2O3 25^*AI20J > 50<»WU2O3 O 75^WVI203 A IOOV0AI2O 3
1 1 6,5 7 L n P (P in N)
Figure 5. Inteifacial cracks length according to the AI2O3 content in the coatings 2. Adherence We tested the mechanical adhesion of AI2O3, HAP and duplex (AI2O3/HAP) coatings by the standardised AFNOR test (figure 2) and the interfacial indentation. The tables 2 and 3 present the adherence <5 {<3 = F/S) with variable plasma-spraying distances between the sample and the torch, with or without a post projection thermal treatment (table 2) and different alumina contents in the coatings (table 3). Without a post plasma-spraying treatment, the more the distance rises, the more a goes down. The maximal velocity of a melted particle is reached close to the exit of the plasma torch. When the distance between the sample and the torch decreases, the impact velocity strongly increases. The adherence is higher. In many cases, a thermal treatment improves the mechanical behaviour of coatings. The relaxation of the residual stresses, the constriction of the cracks and the reduction of the porosity explain these results [7]. In the same manner, the more AI2O3 content increases, the more a increases. We correlated such observations with interfacial indentation tests (figure 5). The crack length is lower when alumina is present in the coating. The initial powder of alumina consists of small grains (40^m) which are entirely melted by the torch. Their complete melting favours a very good degree of flattening and improves the inner cohesion and the coating/metal adhesion. SUMMAR Y A rather complete study of plasma-sprayed coatings was made by dissolution and simple mechanical tests. The strong influence of the experimental parameters on the dissolution kinetics is observed. The important role of alumina in the interfacial adhesion is shown. ACKNOWLEDGEMEN T We would like to express our thanks to S.E.R.F. compagny which provides all samples and coatings and the laboratory Pierre Sue (CEN - Saclay - France) for the irradiations. REFERENCE S 1. De Groot, K., High Tech. Ceramics,ed. by Elsevier Science Publishers, 1987, 381-386 2. Carrerot, H., Rieu, J., Aurelle, J-L., Rambert, A., Bousquet, G., Proceed.Bioceramics and theHuman Body Conf.,ed.by P. Vincenzini,Faenza, 1992, 230-235 3. Norme AFNOR S 94-072 4. Gross, K. A., Bemdt, C. C, SecondPlasma-Technik Symposium,1986, Vol. 3, 159-170 5. Driessens, F. C. M., Van Dijk, J. W. E., Borggreven, J. M. P. M., Calcif. Tissu. Res., 1978, 26, 127 6. Christofiersen, J., Christoffersen, M. R., Journalof CrystalGrowth,1981, 53, 42-54 7. Brown, S. R., Turner, I. G., Reiter, H., JournalofMaterialsScience : Materialsin Medicine,1994, 5, 756-759
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
DESIG N OF A CALCIU M PHOSPHAT E BON E CEMEN T SUITABL E FOR THE FIXATIO N OF META L ENDOPROSTHESE S FCM. Driessens, I. Khairoun, MG. Boltong and JA. Planell Department of Materials Science and Metallurgy, Universitat Politecnica de Catalunya, Avda. Diagonal 647, E-08028, Barcelona, Spain.
ABSTRAC T In order to be suitable for the fixation of metal endoprostheses a calcium phosphate bone cement CPBC must be injectable, whereas strict conditions must be fulfilled for the mixing time MT, the dough time DT, the cohesion time CT, the initial and the final setting time 1ST and FST as well as the compressive strength CS. Due to the fact that CPBC’s build up their strength in not less than 12 hours, the initial fixation of the prostheses must be done by press-fit. However, osteointegration is a matter of days so that from then on the structure is load-bearing all around. Starting with the Biocement H, a CPBC developed earlier, we now succeded in making a formulation called Biocement D which fulfills all requirements: MT < 1 min, DT « CT = 3.5 min, 1ST = 4.5 min, FST = 13 min and CS = 33 MPa. CS is thought to be sufficient for load-bearing applications because it is higher than the maximum compressive strength of human trabecular bone (30 MPa). This CPBC is osteotransductive, i.e. it is transformed into new bone after implantation in contact with living bone so that the final situation is that of the metal endoprosthesis surrounded by and incorporated in the bone structure. INTRODUCTIO N Calcium phosphate bone cements CPBC’s are known since 1983 [1]. They have the advantage over calcium phosphate bioceramics that they do not need to be delivered in prefabricated forms or as granules but that they can be molded during the operation or simply injected into the bone defect [2]. The initial setting time I and the final setting time F of CPBC’s is measured with Gilmore needles [3]. We also designed a test to measure the so-called cohesion time CT of CPBC’s, i.e. the time from when on it does not disintegrate any more when immersed in Ringer’s solution[4]. Up to now we have had some experience with clinicians about the setting behaviour of our CPBC’s. On the basis of that experience we can express the following handling requirements in minutes. 3 1 (2) F < 15 (3) As in most clinical applications the CPBC’s are applied in direct contact with human trabecular bone, it may be stated as a mechanical requirement that the strength of a CPBC for load bearing applications must be at least as high as that of human trabecular bone. In most applications the CPBC will be occluded between trabecular bone and a metal implant surface. 279
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Bioceramics Volume10
Therefore, one may presume that the compressive strength CS is most relevant. As the maximum CS of human trabecular bone is 30 MPa [5] the main mechanical requirement for CPBC’s is expected to be: CS > 30 MPa (4) In the previous study on Biocement H [6], it was presumed that there were certain areas for the liquid/powder ratio L/P and the accelerator concentration (% Na2HP04) in the cement liquid for which requirements (1) through (4) were simultaneously satisfied. To check this was the first purpose of this study. Biocement H consists of an aqueous liquid containing Na2HP04 and a powder containing a-tertiary calcium phosphate a-TCP and 2 % precipitated hydroxyapatite PHA. The second purpose of this study was to see whether additions of CaHP04 and CaCOs to the cement powder could improve the behaviour of the cement. Therefore, Biocement D was developed, of which the powder consists of 58 % a-TCP, 25 % CaHP04, 8.5 % PHA and 8.5 % CaCOs. It was especially important to see whether within the area of fiill compliance with the four requirements (1) through (4) there was a cement paste which was thin enough to be injected in the bone cavity, because this is the most appropriate way forfixationof metal endoprostheses. MATERIAL S AND METHOD S The L/P ratios of the two cements was taken to be either 0.30 or 0.32 or 0.35 or 0.40. The values chosen for the accelerator concentration were 0, 1, 2 Vi and 4 % Na2l-IP04 in water. The setting times I and F were measured as usual [3] and the cohesion time CT with the method developed previously [4]. Teflon molds were used to prepare cement cylinders with a height of 12 mm and a diameter of 6 mm and soaking was carried out during 1, 2 and 5 days in Ringer’s solution at 37 C prior to determination of CS with an Instron Universal Testing Machine Type 4507 at a compression rate of 1 mm min’\ The results were screened on their area of compliance with each of the requirements (1), (2), (3) and (4). For those L/P and Na2HP04 values for which there was full compliance it was investigated whether the paste was enough of a low-viscosity type to be injectable at time CT. RESULTS The results for Biocement H were very disappointing. It was observed that there was no combination of L/P and % Na2HP04 for which there was fiill compliance. According to Table 1 this was much improved for Biocement D. Moreover, the combination L/P = 0.40 ml/g and 4 % Na2HP04 appeared to give an injectable paste for that cement Table 1. Compliance (+) with the four requirements (1) through (4) for Biocement D as a function of the L/P ratio (ml/g) and the accelerator concentration % Na2HP04. 1
% Na2HP04 0 1
\
4
L/P
0.30 -
0.32 -
0.35 -
0.40
+ -
+ +
+ +
_
+
1
1
1
Calcium PhosphateBone Cementfor Fixation of Metal Endoprostheses:Fcm. Driessens et al.
281
DISCUSSION In this study it was found that Biocement H as such is not suitable for chnical applications. The main cause was that requirement (2) is not fulfilled. For the most promosing combinations of the L/P ratio and accelerator concentration % Na2HP04 the cohesion time CT coincided practically with the initial setting time I so that there was virtually no time period in which the clinician can apply and mold the material. According to Miyamoto et al [7] the CPBC invented by Brown and Chow [1] has the same problem, even when phosphate solutions are used as accelerator. Ishikawa et al [8] solved this problem by the addition of sodium alginate to the cement liquid. This addition did not adversely affect the biocompatibility of this material [9]. However, the authors admit that sodium alginate is not allowed yet for parenteral use. According to the present result we succeeded in avoiding any such problems by addition of CaHP04 and CaCOa to our cement powder, keeping it purely inorganic. Simultaneously, an injectable formulation was developed, which was one of the requirements for applications in fixation of metal endoprostheses. In such applications [10] the CPBC is transformed into new bone tissue so that the end situation is a completely osteointegrated metal endoprosthesis.
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REFERENCE S 1. Brown WE, and Chow LC, J.Dent.Res. 1983, 62, 672 2. Driessens FCM, Planell JA and Gil FJ, in: DL wise et al (Eds) Encyclopedic Handbook of Biomaterials and Bioengineering, Part B, Applications, Vol.2, Marcel Dekker Inc., New York, 1995, p.855-877 3. Driessens FCM, Boltong MG, Bermudez O and Planell JA, J.Mater.Sci.Mat.Med. 1993,4,503508 4. Fernandez E, Boltong MG, Ginebra MP, Driessens FCM, Bermudez O, and Planell JA, J.Mater.Sci.Letters 1996, 15, 1004-1005 5. Yaszemski MJ, Payne RG, Hayes WC, Langer R and Mikos AG, Biomaterials 1996,17, 175185 6. Driessens FCM, Boltong MG, Planell JA, Bermudez O, Ginebra MP and Fernandez E, Bioceramics 1993, 6, 469-473 7. Miyamoto Y, Ishikawa K, Fukao H, Sawada M, Nagayama M, Kon M and Asaoka K, Biomaterials 1995, 16, 855-860 8. Ishikawa K, Miyamoto Y, Kon M, Nagayama M and Asaoka K, Biomaterials 1995, 16, 527532 9. Miyamoto Y, Ishikawa K, Takechi M, Yuasa M, Kon M, Nagayama M and Asaoka K, Biomaterials 1996, 17, 1429-1435 10. Jansen JA, Wolke JGC, Hayakawa T, Planell JA and Driessens FCM, 5th World Biomaterials Congress, Toronto, 1996, p.65.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the10thInternational Symposium on Ceramics in Medicine, Paris,France,October1997) '1997 Elsevier Science Ltd
QUANTITATIV E COMPARISO N OF IN VIVO BON E GENERATIO N WIT H PARTICULAT E BIOGLASS fi AND HYDROXYAPATIT E AS A BON E GRAF T SUBSTITUTE . Yoshinobu Fujishiro*, Hironobu Oonishi**, and Larry. L. Hench* Department of Materials, Imperial College of Science, Technology and Medicine, Prince Consort Road London SW7 2BP UK. Department of Orthopaedic Surgery, Artificial Joint Section and Biomaterials Research Laboratory, Osaka-MinamiNational Hospital, Kido-Cho, Kawachinagano-Shi, Osaka, 586 JAPAN. ABSTRAC T Rates of in vivo bone generation were determined by point-counting analysis of (100-300 ycm) particulate Bioglassfi and synthetic hydroxyapatite (HA) in rabbit femora. New bony tissue was observed in 20% of the image area around Bioglassfi particles by 1 to 2 weeks, and the degree of trabecular bone growth increased with time. The interparticle space of Bioglassfi was filled by 60% bonding bone between 6 to 12 weeks. The rate constant of trabecular bone growth in the presence of Bioglassd) was calculated to be 10.9 x 10’^ day"’ at the periphery of the implantation site. HA particles led to smaller rate constants of ca. 4.6 x 10"^ day"’ at the periphery, and the HA particles developed very small amounts of bridging bone. Differences in rate constants for bone growth in the center of the defect were even larger; 7.2 x 10-3 day’’ for Bioglassfi vs 2.0 x 10-3 day’ for HA particles. Quantitative rates of bone growth associated with the particulates matched well with bioactive indices of bulk implants of the same materials. KEYWORDS : Bioactive glass, hydroxyapatite, osteoconduction, osteoproduction INTRODUCTIO N Biocompatible materials, such as calcium phosphates (a-tetracalcium phosphate (TCP), hydroxyapatite (HA) etc.) and bioactive glasses can be effective in the repair of bone defects during orthopaedic surgery. These materials exhibit varying degree of osteoconductive and osteoproductive behavior. The 45S5 bioactive glass (Bioglassfi) is an osteoproductive particulate, and has been in clinical use for filling bone cavities and replacement of lost bone. The bone response to Bioglassd) was compared by our group with a-TCP, b-TCP, Tetracalcium phosphate (TeCP), Octacalciumphosphate (OCP), and synthetic HA using a rabbit femoral defect model [1,2]. Quantitative analysis of in vivo bone growth, the amounts of implanted Bioglassd) and HA particles, regenerated bone, and soft tissue regions were determined by a point-counting analysis of SEM micrographs of sectionsfi*omimplanted samples in rabbit. The effects of materials and location of observation; e.g., the periphery of the defect where contact with intact bone and osteoconduction is possible vs the center of the defect where osteoproduction is most evident, are considered with respect to the calculated rates of bone generation. 283
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EXPERIMENTA L The samples of 45S5 Bioglassfi (45% Si02, 24.5% NajO, 24.5% CaO, 6% P2O5 in weight percent) or synthetic HA particles were prepared to have a size range of 100-300 ^m, implanted into rabbit femora for 5 days to 12 weeks. Experimental details are given in references 1 and 2. The bone-implant samples were observed via backscattered SEM analysis. All data from micrographs were measured by hand-point-counting using a 5-mm square grid sheet (24 X 30 = 720 points/image). Micrographs of different magnification were used for counting. The unit volumes of particle (Np), new bone (N^), original bone (N^^) and soft tissue (Nnb) were estimated by percentage of points, respectively [3]. Ni (%)
=
Pi / P,ota.
X
100
(1)
The ratio of bone growth, x, is calculated by eq. (2), and (3) Nspace(%)
=
100 - ( Np + N , , )
(2)
x(-)
=
N,/N,p,,,
(3)
where Nsp^^e is the percentage of interparticle space around Bioglassfi or HA particles. RESULTS The percentages of Bioglassfi and HA particles, area of soft tissue, and regenerated bone regions as a fiinction of implantation time are summarized in Table 1. The time dependence of the growth of bone associated with the two materials is plotted in Fig. 1. The rate of bone generation in the periphery of the defect filled with Bioglassfi particles increased immediately with time, and continued to increase throughout the 12 weeks. The percentages of regenerated bone in the peripheral regions were 50%, and 60% for 6 weeks and 12 weeks, respectively. The amount of Bioglassfi particles and area of soft tissue continued to decrease slowly at long implantation times. Soft tissue represented less than 6% of the section at 12 weeks. The degree of bone generation similarly increased with time at the center of the implantation site filled with Bioglassfi particulate. The amount of newly generated bone was less at the center than at the periphery for all times with the percentage reaching about 30% by 5 weeks. The percentage of area filled with Bioglassfi particles did not decrease as rapidly in the center of the defect as in the periphery. Sum of the particles and new bone increased as soft tissue was replaced with bone. As shown in Table 1 and Fig. 1, the percentage of regenerated bone at the periphery of the defect for implanted HA particles increased slowly with time, while the area of soft tissue slowly decreased. The rate of bone formation on the HA particles was significantly slower compared with that of Bioglassfi particulate at the periphery of the implantation site for early times. The amounts of new trabecular bone in the periphery were 12% and 30% for 1 week and 6 weeks, respectively. Furthermore, the rate of loss of interparticle space filled with HA particles was slow as compared with Bioglassfi. The amount of new bone was only 15 to 17%) at the center of the defect by 6 and 12 weeks for HA samples. There were no thick bone bridges between HA particles in the center of the defect, a partial thin bone layer appeared on the surface of some HA particles. The percentage of soft tissue in the center of the defect filled with HA particles remained around 30%) from 3 weeks onwards.
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Table 1. Point-counting percentage of regenerated bones, particles 45S5 Bioglassd), and area of soft tissue in SEM micrographs [1,2] by hand-point-counting technique at various implantation times. Materials
Location of Implantation site
Day/days
Percentage of point-counting/% Particles
Soft tissue
Original bone
| Regenerated Bone
Bioglassd)
Periphery
5 7 12 14 21 35 42 84
30.6–0.9 47.0–9.9 40.9–12.7 56.4 32.4–0.5 30.2–5.2 37.8–4.6 33.5
46.0–2.7. 21.8–4.5 37.0–10.1 24.6 29.3–1.1 31.7–11.8 11.5–0.5 6.4
6.6 0.0 8.7 0.0 8.9 2.0 1.4 0.0
16.8–4.9 31.2–5.5 13.4–6.2 19.0 29.4–9.6 36.1–12.7 49.3–2.5 60.1
Bioglassfi
Center
14 21 35 42
45.3–10.0 54.1–1.9 41.6–6.9 51.6–2.6
40.0–3.5 26.0–1.1 28.8–2.0 19.0–5.3
0.0 0.0 0.0 0.0
14.7–6.6 20.0–0.8 29.6–7.7 29.4–2.7
HA
Periphery
5 7 12 14 21 42 84
27.9–2.3 42.7–0.5 43.5 32.3–10.2 34.9 27.1–10.7 38.3
54.7–3.6 45.2–0.9 43.2 34.6–11.9 33.3 24.9–2.3 40.8
10.3 0.0 0.8 11.9 13.6 18.3 5.1
7.1–6.4 12.1–1.4 12.5 21.2–6.0 18.2 29.7–11.4 15.8
HA
Center
14 21 42 84
51.7 44.4 56.1–1.3 49.4
40.8 32.0 28.6–1.0 33.7
0.0 0.0 0.0 0.0
7.5 23.6 15.3–0.3 16.8
Percentag e of Regenerate d Bone
60 j
^^,«»Bioglaae «
Table 2. Relationship between the rate constants of in vivo bone generation and the bioactive indices using Bioglassfi and HA particles.
50 J
Bioactivity bonding index le/day-^
Materials
Location
Rate Constants X lOVday-^
Bioglassfi
Periphery Center Average
10.9 7.2 9.0
12.5
Periphery Center Average
4.6 2.0 3.3
3.1
40 30 20 10
r^
’
-HA
1
0
Time (Days)
Fig. 1 Percentages of regenerated bone in the periphery of a rabbit femoral defect associated with particles of Bioglassfi and HA.
HA
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KINETIC S ANALYSI S The time dependencies of ratio of bone growth, x (see eq. (3)), for Bioglassfi and HA particles were plotted and the relationships between l-(l-x)’^^ and time for the two materials were used to calculate rate constants for bone growth. Formation of bone on the surface of Bioglassfi and HA particles can be described by the surface chemical reaction controlled shrinking core model expressed by eq. (4). l-Cl-x)^’’
=
kx
(4)
where k is rate constant of bone growth [4,5]. The rate constants of bone growth produced by Bioglassfi and HA at both the periphery and the center of the femoral hole are measured by the slopes obtained from plots of eq. (4). The experimental rate constants for bone growth are listed in Table 2 along with the bioactive index (Ig) measured by experimental bone bonding reaction. The bone bonding bioactive index is given by Ig = (100 / to 5^5), where to5bb is the time for more than 50% of the surface to be bonded to bond [6]. CONCLUSION S 1. The bone growth rate constants for particulate Bioglassd) were calculated to be 10.9 x 10’^ day’ at the periphery and 7.2 x 10’^ day"’ at the center of the implantation site using a hand-point-counting technique. 2. The rate constants for HA particles under similar in vivo conditions were 4.6 x 10’^ day"’ at the periphery of implantation site and 2.0 x 10"^ day’’ at the center. 3. For Bioglassd) particulate, trabecular bone was generated by 1 week, with 60% of interparticle spaces being filled with new bone at the periphery of the implantation site in less than 6 weeks, and 30% of the center filled with new bone by 5 weeks. 4. The rate of bone growth for HA particles was relatively slow; the percentage of new bone growth at the periphery of the defect was 30% at 6 weeks and only 15% at the center of the defect. 5. The higher activity of particulate Bioglassfi in terms of rate of trabecular bone production agrees well with previously established rate constants for bonding of bulk Bioglassfi implants in cortical bone.
REFERENCES 1. Gonishi, H., Kushitani, S., Yasukawa, E., Kawakami, H., Nakata, A., Koh, S., Hench, L.L., Wilson, J. and Tsuji, E. and Sugihara, T., Bioceramics 7, 1994, 139. 2. Gonishi, H., Kushitani, S., Yasukawa, E., Iwaki, H., Hench, L.L., Wilson, J., Tsuji, E. and Sugihara, T., J. Clinical Orthopaedicsand RelatedResearch,1997, 334, 316-325. 3. Hench, L.L., Gould, R.W., Characterizationof Ceramics,Marcel Dekker, NY, 1971, 529. 4. Smith, J.M., Chemical Engineering Kinetics,3rd. Ed., McGraw-Hill Book Co, 1981, 642. 5. Habashi, F., Principles of ExtractiveMetallurgy,Vol. 1, Gordon and Breach, N. Y., 1969, 121. 6. Hench, L.L., J. Am. Ceram.Soc, 1991, 74, 1487.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TES T OF BIOACTIVIT Y IN FOUR DIFFEREN T GLASSES .
A.M. Gatti, L.L. Hench*, E. Monari, F. Gonella**, F. Caccavale** Centre di Studio dei Biomateriali, Dip.Chirurgia, Universita di Modena, Italy. Imperial College of Science, Technology and Medicine, Dpt. of Materials, London, U.K. **INFM, Dip. di Fisica, Universita di Padova, Italy.
ABSTRAC T A special preparation was carried out on granules of the glasses under investigation before their implantation in rabbit’s muscles, so as to understand how their bioactivity evolves, to know what the ions involved in the diffusion are and how the diffusion is directed. At the explantation the materials and the surrounding tissues were fixed and prepared for the scanning electron microscopic examinations and the energy dispersive analyses. The results show the difference in degradation of the glasses that induce some considerations on the mechanism of bioactivity. KEYWORD S Bioactive glasses, Dental ceramics, Implant, Biocompatibility INTRODUCTIO N The use of natural and synthetic materials in Medicine grew more and more popular during in the last 30 years. This induced the industry to study and to develop materials with physico-chemical and biological properties more and more similar to those of the natural tissues that they are called to substitute. The research led to develop interesting materials such as those bioactive which, when in contact with tissues or body fluids, react and interact with the environment and speed up some phenomena or develop favourable conditions in which cells can grow and proliferate. A good example of these materials is represented by the bioactive glasses. The first among those, called Bioglassfi, was developed by L.L. Hench at the University of Florida in the Seventies (1). Its main characteristic is that when it is inserted into bone defects, in contact with body fluids, it degrades and releases ions that are used by bone for its reconstruction. Many studies were performed in order to understand the mechanism of this bioactivity, but until now it is not yet fully understood. The present research deals with an in-vivo implantation of active glasses. MATEIUAL S AND METHOD S Nl - Bioglassfi was prepared by melting 46.1 Si02 (in mole %), 26.9 CaO, 24.4 NazO and 2.6 P2O5 at 1300. "^C. in a platinum crucible at 1300 C for 2 hours. In the same way three more glasses with the following compositions were prepared: 287
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N2 - 52.1 Si02 (in mole %) , 23.8 CaO, 21.5 NazO and 2.6 P2O5 N3 - 55.1 Si02 (in mole %) , 22.2 CaO, 22.1 NajO and 2.6 P2O5 N4 . 60.1 Si02 (in mole %) , 19.6 CaO, 17.7 Na20 and 2.6 P2O5 After they were crushed in granules, they were sieved in the range of 3(X)-5(X) |xm . These particles were treated to induce an ion exchange on the first superficial layer: Sodium against Silver. The latter is not present in the original compositio n of the glasses. Every glass was immersed in a solution of AgNOa-NaNO a with a 1% molar concentratio n of Nitrate and heated at 320 ""C for 1 hour (2). So in the first 20 |im-thic k superficial layer the Sodium of the glasses was exchange d with the Silver and now the granules appeared lightly brown. Equal quantities of the so-treate d particles were stuffed into special syringes and sterilise d in ethylen e oxide. Two rabbits were anaesthetise d and bilateral surgical pockets were created in dorsal muscles where the granules were pushed. Every pocket had no communicatio n with the others and was closed with suture. After 60 days the dorsal muscles were explanted , fixed in 4% paraformaldheyd e and dehydrate d in a series of ascending-concentratio n alcohols. The samples were embedde d in methylmethacrylat e and sections were obtained with a diamond saw (Struers, Denmark). With the same technique sections of original compositio n glasses were cut, in order to verify the original morphology and the chemical compositio n for every glass. Both sections were observed under optical and Scanning Electron Microscopy (SEM ) and analysed under an Energy Dispersive System (EDS ) by means of an X-ray microprobe to check the elementa l composition , so to verify their degradation. X-ray dot maps were made on the cross sections of the samples to identify the interface glass-biologica l environmen t and the topographic distribution of the principal elements : Sodium, Silicon, Calcium, Phosphorus and Silver. RESULT S The ED S analyses performed on the original glasses showed that a very homogeneou s superficial layer, ranging 15-23 |im, was rich in Silver and only in the core Sodium was present. Fig. la and b show the image of a part of the bioglass after the ion-exchang e treatment with EDS spectrum, detecte d on the superficial layer, for Silicon, Phosphorus, Calcium and Silver.
T00CNT
Fig. la) SEM image of Bioglassfi after ionexchange treatement .
2 . 10 K E V
10eV/ch
R
Fig. lb) EDS spectrum detecte d on the superficial layer for Si, P, Ca and Ag.
Test of Bioactivity in Four DifferentGlasses: A.M. Gatti et al.
Fig. 2 SEM microfotograp h of implanted Bioglassfi granules.
289
Fig. 3 SEM image of N.2 glass granules
All sections of the implanted materials, showed the glasses degraded and crossed by fractures. Little black debries of Silver were found around the granules, immersed in the connectiv e tissue. This showed the presence of few macrophages and giant cells, since Silver oxide is not completel y biocompatible , for its antibacteria l activity. Figs. 2 and 3 show the SEM microphotograph of implanted Bioglass* (glass N. 1) and glass N. 2 granules. It is possible to note that around a not-degrade d core, there is a grey layer, rich only in Silicon and the Silver shell. The other element s (Na, P, Ca) diffused away along with the superficial Silver, a part of which is still present. No Ca-P-rich layer was developed . The developin g of this compound was demonstrate d by different researcher s and also by the Authors in vitro and in different biological environment s (3, 4). Figs. 4 and 5 show the degradation respectivel y of glasses N. 3 and 4. The thickness of the interface s is lesser than that of Bioglass*. This is due to a lower velocity of degradation, depending on the higher amounts of silica present in the original compositio n of the glasses. The x-ray dot maps show a not degraded core surrounded by a Silicon-rich layer. Silver is still visible in some parts even if its thickness decreased .
2^0 pm
Fig. 4 SEM microphotograp h of N. 3 glass granules.
Fig. 5 SEM image of N.4 glass granules.
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Fig. 6 - a) SEM microphotograph of Bioglassfi granules and X-ray dot maps for Si, Ca and P. b) SEM image of N. 4 glass granules and X-ray dot maps for Si, Ca and P. These glasses do not show any Ca,P precipitation, present instead, though not prominently, when they are not treated \vith Silver. An example of this precipitation is shown in Fig. 6a for a Bioglass* granule. The calcium-phosphate compound is electronically dense, more dense than the layer rich in Silicon. It was demonstrated (5) that after a longer time of implantation, the granules are exhausted by difiEusion and only the not-soluble Ca/P layer remains. The behaviour of glass No. 4 is equal to that without Silver, even if it can be noted a slower velocity of degradation (Fig. 6b). Now the granules do not look electronically dense: they appear transparent also in the core. The external part of the granule is characterised by the presence of a thin layer of Silver and of waves of diffusion. CONCLUSION S Treating the granules with Silver allowed us to understand that the formation of a Ca,P rich layer is not due to the diffusion of the glass ions from the core toward the biological environment, but it is a combination of two different diffusions from and toward the glass. Simultaneously witli the diffusion there is an adsorption. Probably Silicon (of the glass) is exchanged with Calcium (of the extracellular fluid). Silver allows the glass degradation, but prevents the Ca,P precipitation. It is evident that this precipitation is not cell-mediated as some authors claim. The Authors thank Miss Simona Piaggi for her technical help. The work was granted by Ministero dell’Universita’ e della Ricerca Scientifica. REFERENCES 1. Hench L.L., Splinter RJ., Allen W. and Greenlee T.K., J. Biomed Mater.Res. Sytjip. 1971,2,117-141. 2. de Marchi G., Caccavale F., Gonella F., Mattei G., Maraldi P., Battagliani G., Quaranta k.,Appl.PhysA.1996, 63, 403-407, 3. Andersson O.H.and Karlsson K.H., J. Non-Cryst.Solids1991,149, 145-151. 4. Gatti A.M. and Zaffe D., Biomaterials,1991,12, 345-350. 5. Gatti A.M., Hench L.L, Yamamuro T., Andersson O.H., Cells andMat.1993, 3, 283-291.
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
COMPARISON OF BONE-IMPLANT ATTACHMENT STRENGTH BETWEEN THE IMPLANTS WITH HYDROXYAPATITE-COATING AND TRICALCIUMPHOSPHATE -COATING ON TITANIUM ARC SPRAYED TITANIUM K. Hayashi,T. Hara,T. Imamura, Y. Iwamoto Department of Orthopaedics,Faculty of Medicine,Kyushu University, 3-1-1 Maidashi,Higashi-ku, Fukuoka 812,Japan ABSTRACT We compared the bone-implant attachment strength between arc sprayed titanium without ceramic coating(Proarc), hydroxyapatite(HA)-coated Proarc(Proarc HA), and tricalciumphosphate(TCP)-coated Proarc(Proarc TCP) by push-out test in dog femurs.Push-out strength of Proarc-HA was significantly greater than that of Proarc(p<0.05) 4 weeks postimplantation,There was no significant difference between Proarc and Proarc TCP.There was no significant difference between Proarc HA and Proarc TCP.Combination with significant difference between three materials was not found. INTRODUCTION We developed hydroxyapatite(HA)-coated titanium arc sprayed titanium implants to solve the problems of the possible separation of HA from the metal substrate under shear loading[l-3]. It is sometimes pointed out that ceramic coated on titanium should be disappeared as early as possible not to produce wear particles under long-term loading.lt is believed that tricalciumphosphate(TCP) can be resorbed in vivo earlier than hydroxyaptite(HA).Therefore,we fabricated tricalciumphosphate(TCP)-coated arc sprayed titanium implants.This time we compared the bone-implant attachment strength between the HA-coated and TCP-coated arc sprayed titanium implants. 291
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MATERIALS AND METHODS Titanium arc sprayed titanium cylinders(diameter:6 mm, length:13 mm) without ceramic coating(Proarc),HA-coated Proarc were inserted into the diaphyseal region of the right femur of dogs,and a—TCP-coated Proarc into the left femurs.Mechanical push-out test was performed 4 and 12 weeks after implantation. Figure 1 shows each material.
iiliiiiiTiliii i Figure 1. Each material(Proarciarc sprayed titanium without ceramic coating,Proarc-HA:HA-coated Proarc, Proarc TCP:TCP-coated Proarc)
Comparison of Bone-ImplantAttachmentStrength:K. Hayashi et al.
293
RESULTS AND DISCUSSION 1. 4 weeks postimplantation : Push-out strength of Proarc ,HA-Proarc,and TCP-Proarc were 5.816.8,13.9+3.1, and 12.7+5.7 MPa, respectively(N=8). Strength of Proarc HA was significantly greater than that of Proarc(P<0.05).There was no significant difference between Proarc and Proarc TCP.There was no significant difference between Proarc HA and Proarc TCP. 2. 12 weeks iCombination with significant difference between three materials was not found(N=6) (Figure 2).Although significant difference between Proarc HA and Proarc TCP was not found 4 weeks,datas of Proarc HA had a little variation ,compared with that of Proarc TCP.It seemed that Proarc HA had the more stable fixation than Proarc TCP at the early stage after implantation.
N.S
(MPa) 20 r
N.S -|
1
] Proarc mm Proarc HA ^ ^ Proarc TCP
P<0.05 N.S I
II
1
10
I 4 weeks (N=8)
1 2 weeks (N=6)
Figure. 2 Bone-implan t attachmen t strengt h (mean–S.E. ) (Proarc : arc spraye d titanium, Proar c HA: HA coate d Proarc , Proar c TCP: TCP coate d Proarc )
ACKOWLEDGMENT This work was supported by a Grant-Aid for Scientific Research(C)(05807133),The Ministry of Education, Science,Sports,and Culture, Japan.
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REFERENCES l.Hayashi K.,Inadome T.,Mashima T.,and Sugioka Y.,J. Biomed. Mater., Res.,1993,27,557-563 2. Inadome T. ,Hayashi K. ,Nakashiina Y. ,Tsumura H. ,and Sugioka Y. ,J. Biomed. Mater. Res.,1995,29,19-24 3. Nakashima Y.,Hayashi K., Inadome T.,Uenoyama Y. ,Hara T.,Kanemaru T.,and Sugioka Y.,J. Biomed. Mater. Res.,1997(in press)
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EFFECT OF HYDROXYAPATITE COATING ON BONY INGROWTH INTO GROOVED TITANIUM IMPLANTS K. Hayashi, T.Mashima, K. Uenoyama,T.Hara,and Y.Iwamoto Department of Orthopaedic Surgery,Faculty of Medicine, Kyushu University, 3-1-1 Maidashi,Higashi-ku, Fukuoka 812,Japan ABSTRACT Various grooved titanium implants were inserted into the cancellous bone of the intercondylar region of the distal femur of the dog. After sacrifice of the dogs at intervals postimplantation,the implants were evaluated mechanically by means of a push-out test.Our results demonstrated that when grooved titanium implants are used, the addition of hydroxyapatite(HA) coating significantly improved the biologic fixation. In addition, a groove depth of 1 mm was found to give significantly better fixation than 2 mm.When compared to implants with traditional, beads-coated porous surfaces,HA-coated grooved titanium implants were found to show better fixation at 4 weeks after implantation, but, signifacantly inferior fixation at 12 weeks after implantation. INTRODUCTION HA-coated implants have the problem of the possible separation of HA from the metal substrate under shear loading[l,2]. Therefore,we developed the new titanium implants with rough surface by shield arc spray in which HA coating does not cause obstruction of the rough surface[3] . However, HA-coated, grooved, (macrotexured) titanium has been considered useful as an alternative composite.This time,we compared the bone-implant attachment strength of various HAcoated grooved titanium with that of bead-coated titanium. 295
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MATERIALS AND METHODS <1> Uncoated(Ti-Groove) and HA-coated titanium cylinders(HA -Groove)with circular annular grooves(width:2.0min, depth:0.75inm, interval between grooves:2. 0mm) were inserted into the intercondylar region of the dog femurs(Ti-Groove:right side,HA-Groove:left side). Mechanical push-out test was performed to compare the bone-implant attachment strength at 4 and 12 weeks after implantation in the same way of the previous study[1]. <2> HA-Groove and commercially available bead-coated porous titanium(Beads)cylinders were compared in the same way of 1. <3> HA-coated titanium with circular and longitudinal grooves(HA-Lattice Groove)(width:1.0mm, depth:0. 75mm ,interval between grooves:1. 0mm) and Beads were compared in the same way of 1.Figure 1 shows each implant.
iiii A
IB
C
D
Figure.1 A.Ti-Groove(uncoated Ti-6A1-4V with circular annular grooves;B.HA-Groove(HA-coated Ti-6A1-4V with circular annular grooves);C.HA-Lattice Groove(HA-coated Ti-6A1-4V with circular and longitudinal grooves);D.Beads(Beads-coated Ti-6A1-4V)
Effect of HA Coating on Bony Ingrowth into Grooved TitaniumImplants: K. Hayashi et al.
297
Results <1> The strength of Ti-Groove and HA-Groove were 14.8±7.1,61.2± 40.2 kgf, respectively at 4 weeks(N=6)(p<0. 05) and 58.2±32.9 and 55.5±16.4 kgf at 12 weeks(N-6)(N.S.) (Figure 2).
<2> The strength
of Beads and HA-Groove were 50.8±17.1,71.0±21. 5 kgf at 4 weeks (N=12)(p<0.01) and 87.7±26.3,57.7±15.7 kgf at 12weeks(N=10)(p<0.05 )(Figure 3).
<3> The strength of Beads and HA-Lattice Groove were
64.5±30.5,84.6±46.6 kgf at 4 weeks(N=6)(p<0. 05) and 83.4±40.5, 61.7±37.8kgf at 12 weeks(N-6) (p<0, 05)(Figure 4), When compared to implants with traditional beads-implants,both HA-coated grooved implants were found to show better fixation at 4 weeks ,but, significantly inferior fixation at 12 weeks.
(kg-f) 10 0
N.S
Ti-Groove
P<0.0 5 P<0.0 5
N.S
4 weeks (N=6)
12 weeks (N=6)
50
O^-
Figure. 2 Bone-implan t attachmen t strengt h (mean–S.E. ) (Ti-Groove : Uncoate d Ti-6AI-4V with circula r annular grooves . HA-Groove : HA-coate d Ti-6AI-4V with circula r annular grooves. )
HA-Groove
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(kg-f) 100 r
P<0.0 1 P<0.0 1
P<0.0 5
[jjgM Bead s ^^M HA-Groove
50
1 2 weeks (N=10)
4 weeks (N=12)
Figure. 3 Bone-implan t attachmen t strengt h (mean–S.E. ) (HA-Groove : HA-coate d Ti-6AI-4V with circula r annular grooves . : Beads-coate d porous Ti-6AI-4V ) Bead s
I
N.S
P<0.0 5
(kg-f) lOOr
P<0.0 5
P<0.0 5
[jjjijg:!:::; ) Bead s HA-Lattice Groov e
4 weeks (N=6)
12 weeks (N=6)
Figure. 4 Bone-implant attachmen t strengt h (mean–S.E. ) (Beads : Beads-coate d porous Ti-6AI-4V. HA-Lattice Groove: HA-coated Ti-6AI-4V with circula r and longitudinal grooves. )
Effect of HA Coating on Bony Ingrowth into Grooved TitaniumImplants: K. Hayashi et al.
299
DISCUSSION This study indicated that the osteoconductive effects of HA-coating on the grooved titanium were only important to the boneimplant attachment strength that 4 weks and the effect of the morphology of the groove mainly responsible for attachment strength at 12 weeks.More research needs to be addressed to enhance the longterm bony ingrowth of the HA-coated implants. ACKNOWLEDGEMENT This work was supported by a Grant-in-Aid for Scientific Research(C)(05807133),The Ministry of Education, Science,Sports,and Culture,Japan. REFERENCES 1. Hayashi K.,Inadome T. ,Mashima T.,and Sugioka Y.,J. Biomed. Mater. Res.,1993,27,557-563 2. Inadome T.,Hayashi K., Nakashima Y.,Tsumura H.,and Sugioka Y.,J. Biomed.Mater. Res.,1995,29,19-24 3. Nakashima Y.,Hayashi K., Inadome T. ,Uenoyama Y. ,Hara T. ,Kanemaru T. ,and Sugioka Y.,J. Biomed. Mater. Res.,1997(in press)
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Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MECHANIS M OF TH E INFLAMMATOR Y CONVENTIONA L CALCIU M PHOSPHAT E CEMENT.
REACTIO N
OF
Kunio IshikawaV Youji Miyamoto^, Masani Nagayama^ and Kazuomi Suzuki^ ^Department of Dental Materials, Okayama University Dental School, 2-5-1 Shikata, Okayama 770 J^an, and ^First Department and Oral and Maxillofacial Surgery, School of Dentistry, Tokushima University, 3-18-15 Kuramoto, Tokushima 770 Japan
ABSTRAC T To understand the cause of an inflammatCM^ response to conventional calcium phosphate cement (c-CPC) when its paste, not the set mass, was implanted subcutaneously in a rat several calcium phosphate powders were implanted subcutaneously in the rat. Although the set c-CPC showed excellent tissue response, all calcium phosphate powder caused an inflammatory response even though there was a difference in the degree of the inflammatory response. The inflammatory response was severer in the order of dicalcium phosphate anhydrous (DCPA) > tetracalcium phosphate (TTCP) > mixture of TTCP and DCPA, i.e.,powder phase of CPC > crushed set c-CPC. It should be noted that crushed c-CPC also showed an inflammatory response even though it is an apatitic mineral. We concluded, therefore, that CPC shows excellent tissue response only when it is set to form apatitic mass. Thus, CPC should be used so that its setting reaction can be assured. INTRODUCTIO N Calcium phosphate cement invented by Drs. Brown and Chow consists of an equimolar mixture of tetracalcium phosphate (TTCP: Ca4(P04)20) and dicalcium phos› phate dihydrate (DCPA: CaHP04) [1-3]. When mixed with an aque› ous solution, it sets to form hydroxy apatite (HAP: Cal0(PO4)6(OH)2), the putative mineral of tooth and bone. The set mass shows excellent tissue re› sponse towards hard and soft tis› sues. However, conventional CPC (c-CPC) caused a severe inflamma› tory reaction when the paste, not the set mass, was implanted
Figure 1. 301
A4)pearance of rat abdomen 1 week after implantation.
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subcutaneously in a rat immediately after mixing (Figure 1). In contrast fast-setting calcium phosphate cement (FSCPC) [4,5], which set within approximately 5 min, as opposed to 30 60 min - the setting time of c-CPC, and its anti-washout type (aw-FSCPC) [6,7], the paste would not be washed-out but set within approximately S min even if the paste was immersed in serum immediately after mixing, showed excellent tissue response. c-CPC was found to be completely crumbled whereas FSCPC and aw-FSCPC kept the same shape as at implantation. In addition, unreacted DCPA was found only in the case of c-CPC even 24 hours after implantation. One of the reasons for the inflammatory response observed in c-CPC may be the crumbling property of c-CPC. When the paste was arushed into powder, TTCP and DCPA could hydrolyze to form HAP, not in a symmetric way. It also forms HAP powder instead of the set HAP mass. In this investigation, several calcium phosphate powders, 1) TTCP, 2) DCPA, 3) mixture of TTCP and DCPA, and 4) crushed set CPC, were unplanted subcutaneously in rats and tissue response to each powder was examined to shed some light on the cause of the inflammatory reaction of c-CPC. MATERIAL S AN D METHOD S Specimens preparalion TTCP was made by heating the mixture of DCPA and CaCQ3 at 1500 C for 12 hours and crushed into powder as descr9)ed previously. DCPA obtained commercially was grounded in 90% ethanol to reduce the size to 0.9 (Jim in diameter. The powder phase of CPC was made by mixing an equimolar amount of TTCP and DCPA. The CPC powder was mixed with distilled water, at a powder to liquid (P/L) mixing ratio of 3.5 and kept in an incubator at 37 *C and 100 % relative humidity for 24 hours to get set mass. Some of the set mass was crushed to get a crushed set CPC. Animals and implantation procedure Ten-week-old male rats of Wistar strain obtained conmiercially (Charles River, Yokohama, Japan) and given standard pellets and water ad libitumywere used for the unplantation study. All powders were tested by implantation in all of the 20 rats. The rats were anaesthetized by Lp. injection of sodium pentobarbital (Nembutalfi, Abbott Co., Chicago, IL). For the implantation of CPC, the abdomen of the rat was shaved, washed and disinfected with iodine. Three longitudinal incisions of about 1 cm were made through the full thickness of the skin. Subsequently, lateral incisions to the subcutaneous pockets were created by blunt dissection with scissors. Each experimental material (0.3g) was unplanted using a cylindrical mold made by cutting thefrontportion of a 1 cm^ plastic syringe (Terumo, Tokyo, Japan). Set CPC was also implanted subcutaneously in rat as control materials. Finally, the wounds were carefully closed. Histological preparations At the end of the implantation period the rats were killed with an excess dose of Nembutal. After soft x-ray photographs were obtained to record the behaviour of calcium phosphate powers in each rat, the implant materials, including all surrounding tissues, were removed, fixed in 10% neutral buffered formalin and onbedded in methylmethacrylate (HistoDuifi, Leica Co., Nussloch, Germany). After polymerization, thin serial sections were cut using a rotary microtome. The sections were stained with hematoxylin-eosin and investigated by light microscopy. RESULT S
Mechanism of InflammatoryReaction of ConventionalCalcium Phosphate Cement:K. Ishikawa et al.
Figure 2 shows the typical appear› ance of the rat*s abdomen 1 week af› ter surgery. As shown, the most severe swelling with fluctuatio n by palpation was apparent around the DCPA . TTCP , the other compo› nent of the CPC , also showed an inflammatory response . The size of swelling was smaller in the case of the mixture of TTC P and DCPA , i.e., the powder phase of CPC . Crushed set CPC also induced an inflammatory response but its swelling was smallest in size. No gross evidenc e of inflammatory^ re› sponse was observe d when set CPC was implanted subcutaneousl y in rat (data not shown). The size of swelling formed 1 week after im› plantation was in the order of the following (Table 1)
303
^h^^’t^ Crushe d se t CPC Mixture of TTCP and DCPA
..-^ , -Figure 2. Appearance of rat abdomen 1 week after implantation.
DCPA > TTCP > mixture of TTCP and DCPA > crushe d se t c-CP C When the skin covering the calcium phosphate powder was cut with scissors, a copious inflammatory effusion compose d of serous, slightly viscous, yellowish, transparent fluid, was observed. In contrast, no effusion was observe d where set CPC had been implanted, and the set mass was covered only by a thinfibrouscapsule. Histological evaluatio n reveale d that a large vesicle containing abundant inflammatory effusion was formed around the calcium phosphate powder one week after implantation . The wall of the vesicle was compose d of thick vascular granulation tissue which containe d many foreign-bod y giant cells and moderate infiltration of inflammatory cells - consisting of lymphocyte s and plasma cells. In the cytoplasm of the foreign-body giant cells, calcium phosphate powder stained with hematoxyli n was frequently observed. Table 1 Implanted calcium phosphate and size of swelling, the amount of inflanunatory effusion formed subcutaneousl y in rat 1 week after implantation Particle
Powder Size^) dim)
DCP A TTC P Mixture of TTC P and DCP A Qushed set CPC
0.9 11.0
-
n.d.
Size of the Swelling^) Inflammatory Effusion) Length x Width x Height
28 21 18 14
X X X X
a) Average diamete r of calcium phosphate b) In mm effusion
16 X 8 x 11 X 9 x
8 4 2 1
Very Large Large Medium Small
c) Relative amount of inflammatory
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DISCUSSIO N This investigation clearly demonstrated that TTCP and DCPA caused an inflammatory response even though they are the components of CPC. Also, crushed set CPC caused an inflanmiatory response even though it is the apatitic mineral. These resuhs are consistent with the results that c-CPC caused an mflammatory response when implanted immediately after mixing. It is confirmed that the important factor in obtaining satisfactory clinical results is to use CPC so that it can not be crumbled in the implanted area, CPC set to form HAP as shown in eq 1. However the reaction ]»roceeds in a symmetric way only when both TTCP and DCPA stay close. When TTCP aoid DCPA are apart for some reason, for example crumbling property, TTCP and DCPA hydrolyze to produce Ca(0H)2 and H3P04, respectively. The pH of the surrounding area will increase and decrease due to Ca(0H)2 and H3PO4 formation, respectively. 2Ca4(P04)20 + 2CaHP04 3Ca4(P04)20 + 3H20 10CaHPO4 + 2H2O
• •
•
Caio(P04)6(OH)2
Caio(P04)6(OH)2 + 2Ca(OH)2 Caio(P04)6(OH)2 + 4H3P04
(1) (2) (3)
It should be noted that crushed set CPC also caused an inflammatory response even though the size of the swelling was smallest within the experimental group. These results indicate that the size of the implant materials is an important factor to decide the tissue response. The largest size of swelling and the largest amount of inflanunatory effusion observed in DCPA may be the resuU of two factors. First, hydrolysis of DCPA produces H3P04 as by product and reduces the pH of the surrounding tissue. Second, the particle size is smallest among the calcium phosphate powders examined in this present study. These factors are thought to be oweed, in part, to the inflammatory response of c-CPC since unreacted DCPA was found in the case of cCPC. Further evaluation of tissue response to calcium phosphate powder with controlled particle size is awaited to understand the factors to determine tissue response. SUMMAR Y All calcium phosphate mineral relating to the component of CPC mcluding apatite showed inflammatory response when they are unplanted in powder form. Therefore, CPC should be used so that its setting reaction can be assured. ACKNOWLEDGMEN T This investigation was supported in part by a Grant-in-Aid for Scientiflc Research from the Ministry of Education, Science, Sports and Culture, Japan, and in part by a Grant-m-Aid for Scientiflc ResearchfromUehara Memorial Foundation.
REFERENCES 1. 2. 3. 4. 5. 6. 7.
Brown, W.E. and Chow, L.C. US Patent No. 4,612,053 1986. Brown, W.E. and Chow, L.C. In: Cements Research Progress, American Ceramic Society, Westerville 1986, 351-379. Chow, L.C. and Takagi, S. In: Specialty Cements with Advanced Properties, Materials Research Society, Pittsburgh 1989, 3-24. Ishikawa, K., Takagi, S., Chow, L.C. and Ishikawa Y. J Mater Sci: Mater Med 1995, 6, 528-533. Miyamoto, Y., Ishikawa, K., Fukao, H., Sawada, M., Nagayama, M., Kon, M. and Asaoka, K. Biomaterials1995,16, 855-860. Ishikawa, K., Miyamoto, Y., Kon, M., Nagayama, M. and Asaoka K. Biomaterials1995, 16, 527-532. Miyamoto, Y., Ishikawa, K., Takechi, M., Yuasa, M., Kon, M., Nagayama, M. and Asaoka, K. Biomaterials1996,17, 1429-1435.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
FRACTUR E OF ALUMIN A CERAMI C HEA D IN TOTA L HI P ARTHROPLAST Y -REPOR T OF TW O CASE S WIT H HISTOLOGICA L EXAMINATIO N AND PARTICL E CHARACTERISATION Y. Kadoyal, A. Kobayashi^ P. A. Revell 2, H. Ghashi 1 ,Y. Yamano^, G. Scott ^ and M. A. R. Freeman^. 1 Dept. of Orthopaedic Surgery, Osaka City University Medical School. Osaka. Japan. 2 Osteoarticular Research Group, Dept. of Histopathology. Royal Free Hospital School of Medicine. Pond Street, London NW3 2QG. UK. 3 Bone & Joint Res. Unit. Royal London Hospital. Ashfield Street, London El IAD. UK.
ABSTRAC T Two fractures of the ceramic femoral head are reported. Detailed histological examination and SEM characterisation of the extracted particles were performed. Fragmented ceramic acted as a third body which caused severe metal and polyethylene wear, so that urgent revision procedure should be indicated. SEM and histological study showed that ceramic particles themselves were small enough to elicit foreign body reaction. It was also suggested that hydroxyapatite coating might prevent particle migration and subsequent osteolysis. KE Y WORDS : ceramic fracture, hydroxyapatite coating
total hip arthroplasty,
metallosis, ceramic particles,
INTRODUCTIO N The use of a ceramic femoral head in total hip arthroplasty (THA) has been popular because it produces much less polyethylene wear compared with a conventional metal head[1.2]. Ceramic has been preferentially utilised in younger and active patients where the reduction of wear is of particular importance to prevent osteolysis in the long term. However, this material is extremely hard and brittle thus susceptible to fracture. Although several fracture cases of ceramic head has been reported [3-9], the exact nature of the ceramic particles and consequent histological reaction to this particle has not been well documented. In this paper, two failures of the ceramic femoral head were investigated with detailed histological examination of the periprosthetic tissues. Furthermore, particles were extracted by a tissue digestion method [10,11], and characterised with scanning electron microscopy (SEM) . CAS E 1: A 35 year old man who had ankylosing spondylitis underwent left THA in 1988. The component was a Freeman hydroxyapatite (HA) coated stem (Corin Medical Ltd. Cirencester, UK) with 26 mm alumina ceramic head (Vitox_, Morgan Matroc Ltd. Surrey, UK). The acetabular component was HA coated superolateral fin (SLF) design. Thirty-one months after the operation, the patient developed sudden left hip pain when lifting a heavy load. A radiographic examination after 1 month demonstrated a fracture of femoral head (Fig. 1). At revision, the ceramic head was shattered into multiple small fragments. Marked metallosis was noted and the surface of the Morse taper was severely abraded and roughened (Fig. 2). The polyethylene liner had been deeply scored by the trunnion of the femoral component (Fig. 3). The interfaces of both components were sound and well fixed. Histologically, there was good bone growth onto the HA coating over much of the stem. In spite of a heavy metal and ceramic particle deposition in the bone marrow (Fig.4,5), there was no clear evidence of heavy infiltrate immediately next to the HA coating(Fig.4). 305
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SEM examinatio n on the extracte d particles showed that the majority were submicron metal particles. However, ceramic particles around 5 ^m in diamete r were also observed (Fig.6). CAS E 2: A 57 year old woman who had osteoarthriti s was managed with right THA in 1991. The component configuratio n was the same as the first case. In 1995, the patient felt discomfort in her hip and radiography demonstrate dfi-actureof the femoral head. There was no evident history of trauma to the hip.
Figure 1. Fractured ceramic femoral head (white arrow).
Figure 2. Fragmented ceramic Figure 3. Deeply scored head and damaged taper. polyethylene cup (arrowheads).
Figure 4. Metal particles (*) in the bone marrow . HA-bone interface ( arrowheads ) remains intact. x50
Figure 6 SEM photograph showing small ceramic particles (arrows), bar =lpm
Figure 5. Ceramic particles (arrows) exist between the band of metal particles (*). x200
Figure 7. Ceramic particles in the tissue (arrows), Smaller particles are also present. x2(X)
Fracture of Alumina Ceramic Headin TotalHipArthroplasty: Y. Kadoya et al. 307
The intraoperative findings at revision surgery were identical to the first case including diffuse metallosis and damaged femoral taper and polyethylene. The acetabular component was loose. Histologically, numerous small particles presumed to be ceramic particles (2-5|Lim) were observed (Fig.7) with severe metallosis. Polyethylene particles were occasionally seen and these were generally very large (50-100)im) and predominately provoked a giant cell reaction. DISCUSSIO N The use of a ceramic femoral head has been advocated in the young and active patient because of its improved wear characteristics when articulated with polyethylene. However, due to its extreme hardness and brittleness, several cases of fracture have been reported[3-9]. One common findinjg among these fracture cases is the existence of severe and diffuse metallosis. Following ceramic head fracture, small ceramic fragments embedded in the plastic acetabular component are potentially one of the most abrasive materials[3]. The severe damage on the trunnion in our cases has confirmed that fragmented ceramic could act as a third body which causes severe metal wear necessitating urgent revision surgery . In the previous literature on fracture of the ceramic head, few papers mentioned the existence of ceramic particles probably because of severe metallosis. Consequently, no attempt was made to extract and characterise the particles in the tissue. We demonstrated histologically that there were abundant ceramic particles in the infiltrating granulomatous tissue. Furthermore, the extraction and SEM study confirmed that these particles were around 5 ^m which is in accord with the histological observation that they were small enough to elicit a foreign body reaction. Although the duration of particle challenge was relatively short, it was suggested that HA coating might prevent particle migration to the bone-implant interface and subsequent osteolysis. SUMMAR Y This paper highlighted the role played by the fragmented ceramic head as the third body which accelerate the wear of metal and polyethylene. The exact size of the ceramic particles was determined (^ S^im) and shown to be small enough to elicit foreign body reaction. HA coating acted as a seal against particle migration at least during the observed period. REFFERENCES 1. Oonishi, H., Takayama, Y., Clarke I.C, and Jung, H. J. Long-Term Effect of Medical Implants. 1992 2, 37-47. 2. Davidson, J. CUn. Orthop. 1993, 294,3 61-378. 3. Kempf, I and Semlitsch M. Arch. Orthop. Trauma Surg. 1990, 109, 284-287. 4. Hummer, CD., Rothman, R.H., Hozack W.J. J. Arthroplasty. 1995, 10, 848-850. 5. Higuchi, F., Shiba, N., Inoue, A. and Wakebe, I. J. Arthroplasty. 1995, 10, 851-854. 6. Callaway, G.H., Flynn W., Ranawat, C.S. and Sculuco T.P. J. Arthroplasty. 1995, 10, 855859. 7. Krikler, S., Schatzker, J. J. Arthroplasty. 1995, 10, 860-862. 8. Michaud R.J., Rashad, S.Y. J. Arthroplasty. 1995, 10, 863-867. 9. PuUiam I.T. and Trousdale R. T. J. Bone and Joint Surg. 79-A, 1997, 118-121. lO.Campbell P.,. Ma, S., Yeom, B., McKellop, H., Schmalzried, T,P. and Amstutz, H.C. J Biomed. Mater. Res. 1995,29,127-31. ll.Kobayashi, A., Bonfield, W., Kadoya, Y., Yamac, T., Freeman, M.A.R., Scott, G., Revell, P. A. Proc Instn Mech Engrs, Part H in press.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EXPERIMENTA
L COMPARATIV E S T U D Y BETWEE N
ROUGH-BLASTE D AN D HYDROXYAPATIT Young Ho Kim\ Tae Soo Park^
Jong Seok Park^
II Yong Choi\
E COATE D IMPLANT S Myung Ryool Park\
Department of orthopedic surger>^ Hanyang University, Kuri Hospital, 249-1, Kyomoon-Dong, Kuri, Kyunggi-Do, 471-020, Korea Department of orthopedic surgery, Soonchunhyang University, Chonan Hospital, 23-20, Bongmyung-Dong, Chonan, Chungchungnam-Do, 330-100, Korea ABSTRAC T We performed radiographic, biomechanical and histologic comparative assessment between rough-blasted (KB) and HA coated implants to identify the efficacy of HA coating on roughblasted titanium compared to KB surface on titanium in dogs. The results were as foUowings. Radiographically, HA coated implants had earlier and more abundant incorporation and proliferation of bone. Biomechanically, push out failure load increased gradually until 1 year after implantation in both groups and was significantly higher in HA coated implants compare to RB implants since 3 months after implantation. Histologically, more than 90% of surface coverage by bone was achieved since 3 months after implantation in HA coated implants but such same level of surface coverage was achieved as late as 1 year after implantation in RB implants, and earlier and more profuse production of osteoid and earlier maturation of bone with less intervening fibrous tissue around implants were found in HA coated implants compared to RB implants. Conclusively, more active osteoconduction with profuse surrounded marrow tissue maintained until 1 year after implantation in HA coated implants compared to RB implants. INTRODUCTIO N In experimental study, HA coated implants show better bone ongrowth than plain titanium press-fit or porous coated implants[2, 6] and HA can help to achieve such ingrowth even under condition of micromotion[4, 5]. But there were some problems such as the possibility of osteolysin as a reaction to loose HA particls and delamination[l]. We performed radiographic, biomechanical and histologic comparative assessment between rough-blasted(RB) and HA coated implants to identify the efficacy of HA coating on roughblasted titanium compared to RB surface on titanium in dogs. MATERIAL S AN D METHOD S Cylindrical rod of Ti-6Al-7Nb titanium alloy were prepared, measuring 4.5mm in diameter and 6mm in length. Two types of coating were applied on the rod, which were HA coating with 5 //m of thickness using plasma spray technique and rough blasted coating. The rods were inserted into predrilled holes in the lateral cortex of 309
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adult canine femur using sterile surgical techniques. The holes were drilled slightly oversized (4.7mm), which allow the rods to be implanted without undue laxity. A total 42 rods were inserted into the femur of 7 dogs. 21 RB rods were implanted on one side of femur and 21 HA coated rods were implanted on the other side of femur without surgical complication. Dogs were sacrificed 2 at 6 weeks, 2 at 3 months, 2 at 6 months and 1 at 1 year after implantation. 42 bone segments containing the plugs were obtained after cutting with air saw. Radiographs were taken for all 42 segments taken each time to identify the radiographic difference for the osseointegration of 2 kinds of implants with time. The 34 bone segments containing the plugs were positioned in a testing jig to allow accurate alignment of the loading axis to the long axis of the plugs. The plugs were pushed out from the surrounding bone using an Instron machine 8501 with a crosshead speed of 1 mm/minute to get push out failure load each time. The 8 bone segments which were selected at random each time were prepared with the section of 100 lim thick for light microscopic examination. Villanueva stain was used for each prepared section. Percent surface coverage was estimated by the use of a transparent square grid [7]. RESULTS Radiographically, HA coated implants had earlier and more abundant incorporation and proliferation of bone. Biomechanically, push out failure load increased gradually until 1 year after implantation in both groups and was significantly higher in HA coated implants compare to RB implants since 3 months after implantation(Table 1). Histologically, more than 90% of surface coverage by bone was achieved since 3 months after implantation in HA coated implants but such same level of surface coverage was achieved as late as 1 year after implantation in RB implants(Table 2), and earlier and more profuse production of osteoid and earlier maturation of bone with less intervening fibrous tissue around implants were found in HA coated implants compared to RB implants. But histologic findings at 1 year after implantation were similar in both groups, showing well surrounded mature bone, except more profuse appearance of marrow tissue around HA coated implants(Fig. 1, 2, 3, 4). Table 1. Pushout failure load Postop time 6Wks 3Mons 6Mons lYr
Failure load(N) No.
RB
5 423.60–84.83 5 613.13–151.28 4 652.61 –133.07 2 982.47 –115.84
P-value HA
489.32–83.73 791.67–126.89 832.18–59.84 1178.67 –125.56
0.055 0.012 0.018 0.022
ComparativeStudy BetweenRough-Blasted and Hydroxyapatite Coated Implants: Y.H. Kim et al.
311
Table 2. Percent surface coverage RB(?’6)
HA(%)
6 Wks 3 Mons 6 Mons 1 Yr
30 45 60 95
70 90 95 100
DISCUSSION AND CONCLUSION In our biomechanical study, push out failure load was significantly higher in HA coated implants comared with RB implants, which finding was similar to other reports[3, 7]. Our results represented exceptionally high osseointegrative properties in HA coated implants until at least 1 year after implantation, especially since 3 months after implantation. The findings that more than 90% of surface coverage by bone was achieved as early as since 3 months after implantation in HA coated implant may make a significant contribution to the mechanical strength of the interface between bone and HA. In our experiments, dogs were allowed to move freely immediately after implantation, which means they were under condition of micromotion. Even under such condition of micromotion, earlier and more profuse production of osteoid and earlier maturation of bone with less intervening fibrous tissue were achieved in HA coated implants compared with RB implants, which findings were supported by the other reports[l, 4, 5]. Even if two kinds of
Fig 1. Postoperative 6 weeks histologic section (Villanueva stain ; original magnification x 1) of rough-blasted plugs implant in bone shows scanty distributed osteoid around the implant.
Fig 2. Postoperative 6 weeks histologic section (Villanueva stain ; original magnification x 1) of hydroxyapatite coated plugs implant in bone shows abundant distributed osteoid around the implant.
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Fig 3. Postoperative 1 year histologic section (Villanueva stain original magnification X 40) of RB plugs implant in bone shows well maturated bone in close contact with the implant associated with relatively poor surrounding marrow component.
Fig 4. Postoperative 1 year histologic section (Villnueva stain original magnification X 40) of HA coated plugs implant in bone shows well maturated bone in closer contact with the implant associated with abundant surrounding marrow component, compared with the findings of RB implant.
implants had similar histologic findings at 1 year after implantation, which were well surrounded mature bone around the implants, normal morrow tissue existed more abundantly around HA coated implants compared with RB implants. The findings may suggest osteoconductive activity of HA coating was still maintained and remained until 1 year after implantation and efficacy of HA coating for the capacity of osseointegration after implantation. Conclusively, More active osteoconduction with profuse surrounded marrow tissue maintained until 1 year after implantation in HA coated implants compared to RB implants. REFERENCE 1. Bleobaum R.D. and Dupont J.A., / Arthroplasty, 1993, 8, 195-202. 2. Geesink R.G,T. and Hoefnagels N.H.M., / Bone Joint Surg, 1995, 7 7 - B , 534-547. 3. S(5balle K., Acta Orthop Scand suppl 1993, 64, 1-58. 4. S0hdl\e K., Brockstedt-Rasmussen H. and Hansen E.S., Acta Orthop Scand, 1992, 63, 128-140. 5. S(Pballe K., Brockstedt-Rasmussen H., Hansen E,S., / Bone Joint Surg, 1993, 75-B, 270-278. 6. SOdlXe K, Hansen E.S., Brockstedt-Rasmussen H. and Pedersen CM., Acta Orthop Scand, 1990, 61, 299-306. 7. Wong M., Eulenberger J Schenk R. and Hunziker E., / Biomed Mat Res, 1995, 29, 1567-1575.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MECHANICA L AN D BIOLOGICA L PROPERTIE S OF ALUMIN A BEA D COMPOSIT E M. Kobayashi^ ^ T. Nakamura^ T. Kikutam^ Y.Okada^ N. Ikeda^ S. Shinzato^ aiKiT.Kokubo^ ^Dq)artment of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, Kawahaia-cho 54, Shogoin, Sakyo-ku, Kyoto 606-01, J^an, ^Orttiopaedic Surgery, Otsu Red Cross Hospital, Otsu, Japan, ^Nippon Electric Glass Co. Ltd., Otsu, J^an, and ’*Dq)artment of Material Chemistry, Faculty of Engineering, Kyoto University, Kyoto, Japan. ABSTRAC T We have developed a new composite (ABC) consisting of alumina bead powder and bisphenol-^glyddyl methacrylate(Bis-GMA)4)asedresin, which has both high mechanical strength andexcellentosteoconductivity. Aluminabeadof99.7%puretooksphericalforms // minaverage size andcontainedamorphous,and 5-and r-crystalphases. Anothercon5)osite(SGC)filledwith amorphous silica was used as referential material. The proportion offilleradded to each composite was 70% w/w. Mechanical testing of ABC indicated that it would be strong enough for use under weight-bearing conditions. Histological examination using rat tibiae forup to 26 weeksrevealedthat ABC had excellent osteocondactivity, which was eqaivalent to that of a con5)osite containing AW-GC rq)orted previously. And at 26 weeks, no marked biodegradation had occurred. Whereas,in SGC-implanted tibiae, there was poor cJrect bone formation even at 26 weeks. ABC may have a potent promise as a both mechanically strong and highly biocon5)atible material. KEYWORDS ; Alumina, Bis-GMA, Composite, Osteoconduction, Biological Property INTRODUCTIO N Alumina ceramics have good biocompatibility, high mechanical strength, high resistance to fatigue, and excellent lubrication properties. [1] However, alumina ceramics arebioinert andhaveno bone-bonding ability. Thus, attenq)ts havebeen madeto in5)rove the bone-bonding strength, mainly by surface modification. We have succeeded in developing a new conq)osite (ABC), consisting of Bis-GMA-basedresin as an organic matrix and alumina bead powder, produced by fusing a -alumina powder and subsequently (|iencfaing it as an inorganic filler, and in demonstrating that ABC has excellent osteooonductivity.[2] hi order to develop a material which has high bioactivity, higji mechanical strength, and less potential for fatigue, osteooonductivity in rat tibiae of ABC for iq) to 26 weeks after in^lantation were assessed. MATERIAL S AN D METHOD S 1) Prq)aration of powder Bulk a -alumina (99.5%pure AI2O3) was prq)aredin an electrical melting fmnacefrom calcined alumina powder,producedbyBayer’sprocess. The bulk a -aluminawasthenpulverizedandparticles under 10 fi m in diameter were coUected. The coUectedpowderwas fused and quenched subsequeiitily to produce alumina bead powder (AL-P). Powder XRD andFT-IRRS oftheAL-P showed that it 313
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contained amorphous and 6 -crystal phases of alumina in its main crystal structure . Its average paitide size was 3.0 fi m, the density was 3.6 g/cni, and specific surface area was 0.7 t i i / g. Spherical particles of amorphous silica glass powder (SG-P) were also prq)aredby the fusingcpendiing method firomhighly purified quartz (>99.7% pure). It had the density of 2.2 g/cni, average particle size of 3.0 /x m, and specific surface area of 12 mV g. Each powder was silane4reate d with r -methacryloxypropyltrimethoxysilane , and benzoyl peroxidB at 0.4% per unit weight of the treate d powder was added Bis-GM A and triethyleneglyoo l dimethacrylat e (TEGDMA ) weremixed in equal weight. N,N-dimethyl^-toluidine , at 1.0%perunit weight ofthemixtur e ofBis-GM A and TEGDMA , was dissolved.[3-6 ] The weight ratio of the filler powder mixed into the composite was 70%. The composite was prepared by mixing the ^propriate powder into the mixture of Bis-GM A and TEGDM A and stirring it for 1 minute. It was polymerize d within 3-4 minutes. The two types of composite , containing either AL- P or SG-P as afiller,were designate d AB C and SGC , respectively . The ultimate con5)ressiv e strength, bending strength, dastic modulus of bendng (Young’s modulus), tensile strength and fiacturetoughness of ABC and SGC , measured after soaking in simulatedbocf y fluid (SBF) at 37t: for 1 day, were 196–4 and237– 14 (MPa), 151 – 10 and 157 –10(MPa),7.2–0.2and8.6–0.2(GPa),58–3and59–5(MPa),andl.44–0.05andl.69–0.1 1 (MP a m^ ^), respectively . [2] 2) Aiumal e>q)erimen t Ten-week-ol d male Wistar rats were operate d on under general anesthesi a (Nembutal: 40 mg/kg hodyweight). Cortical bone defect s measuring 2 X 5 nun were aeated on the medial aspect of the proximal metaphysis of both tibiae, and the bone marrow was curette d The intramedullary canals firom both bone defect s in each individual animal were packed with the same kind of composite , and allowed to cure in situ.Twdve rats receive d ABC and twelve SGC ^ with three rats in each groi?) being killed at 2, 4, 8, and 26 weeks after the operation.[2-6 ] 3) Histological examinatio n All tibial segments containing composite samples were excise d and dehydrate d in serial dilutions ofethanol , then embeddedi n polyesterresin . Thin sections (500 // m thick) were cut with a band saw (EXAC T BS-3000, Nonderstedt , Gemiany), papendicula r to the axis of the tibia Two sectionsfiromeach tibia in the ABC and SGC (i.e. 12 specimen s in total for each subgroup) were ground to a thickness of 100 /x m using a diamond 1^ disk (Maruto Ltd., Tokyo, J^an)for contact microradiogr^hy. Several sectionsfiromthe six subgroups were ground to a thickness of 100 ix m for Giemsa surface staining. Several 500 /x m sections takenfiromthe AB C and S GC gjcovps at every time interval studied were polished with damond p^er. These sections were used to study the bone-compositeinterface,usin g a SEM (Hitachi S-800, Tokyo, J^an) connectedt o anEDX(Horib a EMAX-3000 , Tokyo, Japan).[2-6 ] RESULT S Histological examinatio n by contact microradiogr^hy and Giemsa surface staining reveale d that new bone had formed direcdy on the ABC surface, without an intervenin g soft tissue layer, by 2 weeks post-inq)lantation , and at 4 and 8 weeks, newly formed bone almost completel y surrounded the con9)osit e surface within the tibiaeimplanted with ABC . Furthermore,this was maintained at 26 weeks after the operation . However, in SGC-implanted tibiae,poor direct bone formation was observed on the SG C surface throughout the e5q)erimenta l period (Figure 1). Examination by SEM-ED X clearly demonstrate d direct bone formation on the AB C surface (Fig. 2a). ED X profiles of the bone-ABC interface reveale d slightiy inaeased intensity for calcium and
Mechanical and Biological Propertiesof Alumina Bead Composite:M. Kobayashi et al
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.* .
^^) (b) Figure 1. Contact inicroradiogram of ABC and SGC in rat tibiaeat 2 weeks after implantation (a) AB C and (b) SGC . C; composite , B; bone.
(a) (b) , (a) Back-scattere d electro n image, (b) Figure 2. SEM-ED X of AB C at 8 weeks after implantation EDXprofiles. C, composite ; B, bone; Ca, caldum; P, phosphorus; Al, alununum.
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phosphorus (Fig. 2b). In SGC implanted tibiae, no such a layer was evident until 26 weeks. DISCUSSIO N The results of the present stu^ indicate that ABC has excellent osteooonductive ability. Although few studies have been conducted on amorphous alumina as a biometerial, amorphous alumina is thought to have excellent biocompatibility.[7] In the previous study, although apatite was not induced on the ABC surface after soaking in SBF for 28 days, the ABC surface had made direct contact with bone via a l^er containing calcium, phosphorus, and alumina powder. [2] Li et al. rq)orted that alumina gd did not induce ^atite formation when inunersed in SBF for 21 d^s, whereas both pure silica gel and gd-derived titania were hydroxy^atite inducers. [8] However, once a material is implanted in the body, it elicits several responsesfromliving tissues. They include protein adsorption and cell attachment and adhesionas well as ionic exchange. We deduced that ABC has the ability to bond directly withbone, which was induced not by a simple diemical reaction but by some surface property of the AL-P whidi encouraged calcification or apatite formation due to the actions of proteins and cdls in vivo. The predse medianism of direct bone formation on the ABC surface is as yet undear. However, we consider ABCto showpromiseasabasisfordevelopingahighly osteoconductive and mechanically strong biomaterial. We are now plarming to evaluate its bone4)ondng ability, and mechanical properties after long-term in5)lantation. REFERENCES 1. Z. Li, T. Kitsugi, T. Yamamuro,Y. S. Chang, Y. Senaha,H. Takagi,T. Nakamura,arKi M. Oka, "Bone-bonding behavior under load-bearing conditions of an alumina ceramic implant incorporating beads coated with glass-ceramic containing apatite and woUastonite," J. BiomedMater.Res., 29, 1081-1088 (1995). 2. M. Kobayashi, T. Kikutani, T. Kokubo, and T. Nakamura, "Direct bone formation on alumina bead con5)osite," J. BiomedMater.Res., in press. 3. K. Kawanabe, J. Tamura,T. Yamamuro, T. Nakamura,T. Kokubo, andS. Yoshihara, "A new bioactivebone cement consisting of BIS-GMA resin andbioactive glass powder," J. Appl.Biomater.,4, 135-141 (1993). 4. J. Tamura,K. Kawanabe,M. Kobayashi,T.Nakamura,T. Kokubo, S. Yoshihara,andT. Shibuya. "Mechanical and biological properties of two types of bioactivebone cements containing MgO-CaO-Si02-P205-CaF2 glass and glass-ceramic powder," J. Biomed. Mater.Res., 3 0, 85-94 (1996) 5. M. Kobayashi, T. Nakamura, J. Tamura, H. lida, H. Fujita, T. Kokubo, andT. Kikutani, ’^Mechanical and biological properties of bioactive bone cement containing silica glass powder," J. Biomed.Mater.Res., in press. 6. M. Kobayashi, T. Nakamura, J. Tamura,T. Kokubo, andT. Kikutani, "Bioactivebone cement: con5)arison of AW-GC filler with hydroxy^atite and 13 -TCPfillerson mechanical and biological properties," J. Biomed.Mater.Res., in press. 7. A. Naji and M.F. Harmand, "Cytocompatibility of two coating materials, amorphous alumina and silicone carbide, using human differentiated cell cultures," Biomaterials,12, 690-694 (1991). 8. P. Li, C. Ohtsuki, T. Kokubo, K. Nakanishi,N. Soga, and K. De Groot, "The role of hydrated silica, titania, and aliunina in inducing apatite on implants," J. Biomed.Mater. Res., 2S, 1-15(1994).
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
REMODELIN G STEM S
OF BON E AROUN D HYDROXYLAPATITE-COATE
D
FEMORA L
A. A. Edidin and M. T. Manley R & D Corporate, Osteonics Corp., Allendale, NJ, USA
ABSTRAC T Stress shielding of the proximal femur following hip arthroplasty has been well documented around both cemented and cementless femoral stems. Attempts to limit the degree of stress shielding fall into two primary classes: Reduction of the structural stiffness of the stem, usually by reduction of its modulus of elasticity, and the addition of proximal interface enhancements designed to transmit axial forces as proximally as possible to the surrounding bone. We report on a series of patients who underwent THA using femoral stems at opposite ends of the design space. Specifically, one part of the cohort received an extensively porous-coated CoCr prosthesis while the other part received a proximal HA-coated Ti6A14V prosthesis. Using Dual Energy X-Ray Absorptiometry (DEXA) Bone Mineral Density (BMD) quantification techniques, we were able to detect an up to 30% greater retention of proximal bone density in the Ti6A14V cohort as compared to the CoCr cohort. INTRODUCTIO N Stresses in the femur arise from the axial and bending components of the load across the femoral head. While the ratio of axial to bending load is patient and gait pattern specific, the bending component dominates by about 3:1. Thus attempts to transfer more of the bending load to the bone must be focused on reducing the structural stiffness (EI) of the implant. While the moment of inertia is more quickly reduced than the modulus by simply narrowing the implant’s cross-section, physiology using cemendess implants dictates that an implant must be large enough to contact the endosteum if biological integration following mechanical stability is to occur. Thus reduction of the modulus.of elasticity (E) is the most effective means of reducing the implant’s structural stiffness. Attempts to transfer the axial load component proximally are generally limited to encouraging biological integration in the metaphyseal region using coatings or ingrowth regions. This report compares two successful femoral arthroplasty stems at opposite ends of the mechanical spectrum. The first stem is made of cobalt-chromium alloy with a modulus of 220 MPa, and is extensively coated using porous sintered beads. The second stem is made of Ti6A14V alloy with a modulus of 110 MPa and is proximally HA-coated. While both stems have a long (>10 yr.) and successful clinical record, the latter stem would be expected to stress shield the femur to a lesser degree. We used the DEXA technique to determine if in fact this expectation was met. 317
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MATERIAL S AND METHOD S Patients requiring unilateral primary arthroplasty of the hip with simple osteoarthritis as a diagnosis were eligible for inclusion. Inflammatory arthritis, trauma, femoral dysplasia, and trochanteric osteotomy were grounds for exclusion. Twelve patients received an extensively porous-coated CoCr hip stem (AML, Depuy, Warsaw, IN) and 18 patients received a proximally HA-coated Ti6A14V hip stem (Omnifit-HA, Allendale, NJ). The mean age of the patients in the first group was 54 years; patients in the second group had a mean age of 50 years. There were 5 males and 7 females in the CoCr cohort with 8 males and 10 females in the Ti6A14V cohort. DEXA films were obtained preoperatively, at five days post-operatively, and at six weeks, six months, and one and two years post-operatively. BMD was measured both in the Gruen Zones about the femur and alternatively in 2 cm intervals about the femur. The former measurements provided a stem-proportionalized breakdown of BMD changes, while the latter provided a stemlength independent assessment of BMD changes. RESULT S Radiographs of patients with each of the implants are shown in Figures la and lb. Both patients show good bone quality in keeping with their relatively young age and simple OA diagnosis.
Figure la: Lateral radiographof a patientwith the CoCr stem.
Figure lb: Lateral radiographof a patient with theTi6Al4V stem.
Bone mineral density changes for the anterior-posterior view Gruen zones are shown in Figure 2a and 2b for the HA-coated Ti6A14V and porous-coated CoCr stems respectively. In the proximal zones there was 25% greater bone resorption measured in patients receiving the CoCr stem as opposed to the Ti6A14V stem at two years follow-up. The bone resorption was also seen to move further down the stem into zones 2 and 6 in patients implanted with the CoCr stem.
Remodelingof Bone Around Hydroxyapatite-CoatedFemoral Stems: M.T. Manley et al.
Zone 1
Zone 7
Zone 2
Zone 6
Zone 3
Zone 5
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Zone 4
Post op S ix Weeks 0 6 Months H i Year [ 2 Year s
Figure 2a: AveragedBMD Changes by Gruen Zone as measuredin the patientcohort receivingtheHA-coated prosthesis.
Zone 1 0 -5 10 15 20 -25 30 -35 -40
Zone 7
Zone 2
Zone 6
Zone 3
Zone 5
Zone 4
^ I I I’ I I I’ 111^ IM ’ HP’ LP I T 11 i^ i r I
91B
rj |^[r n
I
Post op S ix Weeks De Months H i Year 2 Year s
Figure 2b: AveragedBMD Changes by Gruen Zone as measuredin thepatient cohort receivingtheporous-coatedprosthesis.
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DISCUSSIO N Maintenance of proximal bone density at preoperative levels is not expected nor possible using existing technology. In order to preserve as much bone stock as possible, the procedure should still be designed to minimize proximal stress shielding. This study showed that by using a stem with a lower structural stiffness in conjunction with a biocompatible proximal coating, the degree of stress shielding could be substantially reduced. Because the femoral stems investigated in this study bound the extremes of the design space available today, the results presented herein may be considered to bracket the extremes of expected bone loss after hip arthroplasty. Thus the use of a lower-stiffness HA-coated stem may reduce the magnitude of proximal stress shielding by up to 25% at two years. In addition the region of most pronounced resorption is limited to the proximal two zones, as opposed to the mid-stem bone resorption seen using the CoCr porous-coated stem. ACKNOWLEDGMENT S The authors gratefully acknowledge the participation of William Jaffe, M.D., Fredrick Jaffe, M.D., and David Scott, M.D.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PROCESSIN G AN D CHARACTERISATIO N OF BIOLOGICA L HYDROXYAPATIT E DERIVE D FRO M CATTLE , SHEE P AN D DEE R BON E Michael R. Mucalol, Glenn S. Johnson 1 and Michel A. Lorier^ iChemistry Department, University of Waikato, Private Bag 3105, Hamilton, New Zealand, ^Meat Industry Research Institute of New Zealand, P.O. Box 617, Hamilton, New Zealand
ABSTRAC T In New Zealand, the routine slaughter of cattle, sheep and deer produces a large amount of waste bone which is normally converted to low-cost blood-and-bone fertiliser. The porous architecture of bone makes it a valuable material for use in biomedical implants and thus this poster will describe a study into the processing and characterisation of waste bone from animal species for conversion to materials for clinical purposes. Bone samples from these species were defatted using novel methods such as microwaving which was found to be a highly rapid fat removal method. Defatted bone cubes were then bleached using hypochlorite reagents. Infrared spectroscopy monitored bulk fat removal and have demonstrated that the hypochlorite treatment decollagenated bone cubes while leaving carbonate intact. Solid state NMR showed there was still some tenaciously held organic matter in the samples even after hypochlorite treatment. The work demonstrates that waste animal bone can be efficiently processed to produce modifiable materials for clinical use. KEYWORD S bone, hydroxyapatite, implants, defatting, deproteination, FTIR/NMR INTRODUCTIO N New Zealand as a major producer of meat foodstuffs often has to deal with the large amounts of by-products that result from the routine slaughter of livestock. Often these by-products have no use other than as materials for low value products or else are disposed of, which can potentially lead to environmental problems. A collaborative project between the University of Waikato and the Meat Industry Research Institute of New Zealand was established with the aim of converting one of these waste products, animal bone, to high value-added products. The unique porous architecture of bone makes it an ideal material for use in non-loading bioactive implants where tissue ingrowth is an important requirement. In addition, such bone is an obvious source of hydroxyapatite which may be used in the production of synthetic biomaterials or as a matrix for Drug Delivery Systems. An additional advantage is that the restrictions which necessitate the extraction of bone materials only from controlled herds (e.g. Kiel bone in Germany) in the European Community do not apply in New Zealand. This project was initially inspired by an actual clinical application which required that the bone material be aesthetically presentable and cuttable to a desired form. As-received bone after defatting is extremely hard and therefore cannot be cut or preformed into a convenient size as can synthetically produced porous hydroxyapatite. Further treatment of the bone was, therefore, 321
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necessary to soften the structure to allow for shape modification of the bone material. MATERIAL S AN D METHOD S All solvents, solutions and reagents employed in this study were used as received without further purification. Pre-frozen bone specimens from cattle and deer were cut from the condyles of mature beasts using a sharp band saw. Defatting of the bone involved the use of several procedures. In one procedure, the bone samples were effectively autoclaved by boiling for several days in a domestic pressure cooker at ca.l5 psi. Autoclaved bone samples were subsequently immersed in refluxing isopropanol solvent at 82 C in a wide-necked round-bottomed flask for 90 minutes after which excess solvent was removed and the specimens air-dried at ambient temperatures. In another novel procedure, raw bone cubes were placed in ca. 50 mL of water in a beaker and microwaved for 5 minutes in an 800 W domestic microwave oven. The water which contained a significant level of fat was then discarded and the bone cubes subsequently immersed in refluxing isopropanol (82’’C) to extract the remaining fat. Extraction using supercritical CO2 was also trialled. In this procedure, 5 bone cube specimens which had been previously autoclaved as described earlier were placed in a steel reaction vessel and subjected to supercritical CO2 solvent treatment. After 90 minutes of this treatment, the residue collection trap was inspected for fat drawoff and then every 15 minutes after this. After a total two hour treatment time, fat drawoff was not detected and the extraction was judged to be complete. As a comparison, 10 bone cubes not subjected to the previous autoclaving treatment were also subjected to the supercritical CO2 extraction. Fat drawoff from these specimens, in contrast, was still being observed after ca. 3 hours and 15 minutes of supercritical CO2 treatment. Defatted bone specimens destined for clinical use were subsequently subjected to a bleaching process which served to both improve the aesthetic appearance of the bone as well as to decollagenate and thus soften the bone materials in order that they could become cuttable. As with the defatting procedures, several bleaching processes were employed all of which made use of hypochlorite. In the "Cadivar" method of bleaching [1], bone specimens were heated in 1 L of a cloudy solution containing 150 g of Na2C03, 100 g of Ca(0Cl)2 and 150 g of NaOH. Variations on this method involved using lower strength solutions. Other bleaching methods involved the use of commercial strength (1% and 3% NaOCl w/v) bleach solutions. Hydroxyapatite powders were produced by autoclaving partially granulated pieces of bone material and subsequent enzymatic treatment using food grade enzymes such as lipase, nutralase and alcalase to break down extraneous organic matter adhering to the the exterior of the specimens as well as bone marrow. A 10-day dissolution in 5% (v/v) HCl solution ensued. Filtration of scum and undissolved bone material gave a solution which was subsequently treated by stirring in with saturated Ca(0H)2 solution in order to reprecipitate the calcium phosphate. The crude phosphate powder was further purified by redigestion in acid, filtration and subsequent reprecipitation. All bone materials were characterised after the processing steps in the form of ground powders. IR spectra were recorded as KBr disks on a Digilab FTS-40 FTIR spectrometer. ^IP and 13C Solid State NMR spectra were recorded on a Bruker AC200 NMR spectrometer equipped with a solid state probe and MAS using KH2PO4 (for 3lp) and adamantane (for l^c with crosspolarisation) as secondary references. Fat and protein levels in bleached bone specimens were determined using gravimetric petroleum ether soxhlet extraction and Kjedahl analytical procedures. RESULT S AN D DISCUSSIO N Cuttingand Defatting The cutting history of the bone cubes from the condyles was found to be a critical factor
Processing and Characterisationof HA Derivedfrom Cattle,Sheep and Deer Bone: M.R. Mucalo et al.
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in achieving efficient defatting. The continued use of a blunt blade led to significant heating of the cutting blade and tended to produce cubes with "sealed" surfaces caused by the conversion of the collagen to gelatin upon contact with the heated metal surface. This effectively blocked pores on the cube thereby preventing efficient defatting. The literature shows that a variety of defatting and deproteination procedures have been used on natural bone [2-6]. Cutting the bones in the frozen state appears to be of importance in past studies [5,6] and was adopted in the present study. However, the use of hydrazine as a deproteination reagent as used by previous workers [2,3] was not deemed suitable for bone destined for eventual clinical use due to concern about hydrazine residues and thus solutions containing hypochlorite were used as a clinically more acceptable alternative. In general, autoclaving of the raw bone samples in water followed by immersion in refluxing isopropanol was judged to be the best method for defatting. Specimens prepared using this procedure and subjected to later deproteination gave the most aesthetically presentable specimens for clinical use. Microwaving of the raw bone specimens was trialled as a novel defatting method due to the fact that microwave treatment will selectively heat the bone matrix immersed in solvent. When initially microwaved in water, bone cubes were observed to heat up in the water which caused streams of liquefied fat to pour out of the bone and collect on the the top of the solution. Removal of water and the subsequent treatment in refluxing isopropanol removed the bulk of the remainder of the fat. Overall, a 58% weight loss was observed in bone cubes treated by this method. Isopropanol was the most effective solvent for fat removal due to its lower polarity compared to solvents such as methanol or ethanol. Microwaving is a highly rapid way of removing bulk fat but an undesired side-effect is excessive heating of the bone in its raw state which can lead to collagen being transformed to gelatin inside the bone thus hindering the penetrability of solvents into the porous structure and hence the defatting and deproteination efficiency. Supercritical CO2 extraction has been used on sheep bone before [6] since CO2 is known to be a good solvent for lipids. Also penetrability of supercritical fluids into porous materials is less of a problem due to elimination of surface tension. The supercritical CO2 extraction studies in the present study demonstrated that some prior treatment of the bones such as autoclaving is essential to cut down on defatting time. The specimens previously subjected to autoclaving were relatively clean in appearance after the supercritical treatment although the "raw" specimens tended to be bloodstreaked and less well-defatted relative to the previously treated samples. Deproteinationand SpectroscopicCharacterisation The cadivar method proved to be an efficient method for efficient bleaching (both inside and outside the cube) and deproteination, but tended to soften the bone cubes excessively to such an extent that they were not clinically acceptable. A problem with bleaching is that deproteination is not uniform so that the outer porous network of bone is weakened more relative to the interior structure. Use of 1% commercial bleach solution was found to be the most acceptable clinically in terms of the strength and aesthetic characteristics. In such samples, the fat and protein content was found to be less than ca. 0.1 g %. Fig. 1(a) is an FTIR spectrum of bovine bone after the microwave(H20)/refluxing isopropanol treatment. The spectrum is typical of bone IR spectra showing features due to collagen, hydroxycarbonate apatite and entrained water, however fat-associated hydrocarbon peaks are extremely weak. Fig. 1(b), in contrast demonstrates the change caused by hypochlorite treatment (using commercial NaOCl solutions). Collagen features (and associated water peaks) are generally absent although the carbonate in the natural bone structure is left intact. In this deviation from the natural bone state, the bone becomes softer and easier to shape to a desired form.
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2000
1
r
1800 1600 1400 1200 Uavenumbers
1 400
Figure 1. FTIR spectra of KBr disks of (a) microwave(H20)/isopropanol-refluxe d bone and (b) microwave(H20)/isopropanol-refluxe d bone subjecte d to NaOC l treatment . In the FTIR spectra of the reprecipitate d hydroxyapatit e from bone, there is predictably an absence of collagen feature s and a much reduced carbonate peak at 1427 cm-1. Although fatassociate d peaks are weak in the reprecipitate d hydroxyapatit e specimens , there is still believe d to be organic matter tenaciousl y held to the surface of the reprecipitate d powders. This was demonstrate d by solid state ^^c NM R spectroscop y which is not often used to characteris e bonederived hydroxyapatit e powders. It was found that in bone or reprecipitate d hydroxyapatit e subjecte d even to hypochlorite treatment , a relatively well-define d complex peak at -30 ppm was observed in 13C NM R spectra which could be due to residual organic component s (e.g. fat and/or collagen/gelatin ) adsorbed strongly to the solid particles at levels not detectabl e by FTIR . Occasionally, weak peaks at 168-17 0 ppm and 185 ppm due to carbonyl groups in collagen and fat respectivel y were also observed . Since the l^c NM R spectrum of this residual organic matter differs from that of pure solid collagen and pure fat derived from bone, this will require further investigatio n by solid state NM R and X-ray photoelectro n spectroscop y to clarify the interpretation . ACKNOWLEDGEMENT S We wish to acknowledg e the New Zealand Foundation for Research, Science and Technology for funding support for this project. We are also grateful to E)r Roger Mederr of Forestry Research Institute of Rotorua for recording of solid state NM R spectra.
REFERENCES 1. 2. 3. 4. 5. 6.
University of Otago Medical School, Dunedin, New Zealand, Private Conmiunication Walters, M.A., Leung, N.C., Blumenthal, R.Z., LeGeros, R.Z., Konsker, K.A., J. Inorg.Biochem,, 1990, 39, 193-200 . Rehman, I., Smith, R., Hench, L.L., Bonfield, W., J. Biomed. Mater,Res.,1995, 29, 1287-1294 . Akazawa, T., Kodaira, K., Phosphorus Res.Bull.,1991,1 , 215-220 . Walsh, W.R., Ohno, M., Guzelsu, N., /. Mater. Sci.Mater. AfeJ.J994 , 5, 72-79. Frayssinet, P., Asimus, E., Autefage, A., Fages, J., J. Mater. Sci.Mater.Med.,1995, 6, 473-478 .
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CATASTROPHI C WEA R O F META L BAL L O F BIPOLA R HI P PROSTHESI S AFTE R FRACTUR E O F ALUMIN A CERAMI C SCREW S USE D FOR ACETABULA R BON E GRAF T H. Ohashi, Y. Yutani, A. Kobayashi, Y. Kadoya and Y. Yamano Department of Orthopaedic Surgery, Osaka City University Medical School, Asahimachi, Abenoku, Osaka 545, Japan ABSTRAC T Adverse effect of alumina ceramic was investigated in a patient with a rapid progress of prosthetic loosening after catastrophic wear of metal ball following a fracture of alumina screws used for acetabular bone graft. Alumina ceramic fragments were detected on the bearing surface of polyethylene insert. The fragments were considered to rasp the surface of metal ball producing abundant sub-micron metal particles. The levels of II-la, II-1P, 11-6 and TNF-a in the joint fluid were high. The progress of prosthetic loosening was supposed to be accelerated by production of the cytokines. Since alumina ceramics, especially these of screw shaped, are brittle, we concluded that the use of alumina ceramic screws in prosthetic replacement was contraindication. KEYWORDS : alumina ceramic, metallosis, total hip replacement, wear, metal particles, loosening INTRODUCTIO N Alumina ceramic fragments exsisted in articulating interface of joint prostheses can cause typical third-body abrasive wear. Several cases of severe metallosis have been reported after fracture of alumina ball [1] and alumina screws [23] In these cases, prosthetic loosening was observed, however little has been discussed about the mechanism of loosening. Recently, small polyethylene (PE) particulates are reported to play a great role to induce periprosthetic osteolysis by activating macrophages [4,5]. In case of metallosis, the mechanism of loosening is not well recognized. We experienced a rapid progress of loosening after catastrophic wear of metal ball of a bipolar hip prosthesis with a fracture of alumina screws. The aim of this study was to reveal the mechanism of the catastrophic wear and to investigate the relationship between a fracture of the alumina screws and prosthetic loosening. This adverse effect of alumina ceramics may warn its clinical applications in certain conditions. MATERIAL S AN D METHOD S A 47-year-old woman imderwent a right bipolar hip hemiarthroplasty (Bateman UPF, Co-Cr alloy) due to bilateral coxarthrosis. Acetabular dysplasia was supplemented with a bone graft fixed by two alumina ceramic screws (Sapphire Screw, monocrystal alumina ceramic) (Fig. 1). Her postoperative course was uneventful, howener proximal migration of the outer head was observed on serial radiographs. Nine years postoperatively, radiographs revealed a fracture of alumina screws due to impingement with the migrated outer head. Five months later, she felt severe coxalgia (Fig. 2) and a revision arthroplasty was performed. 325
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F i g . 1 Radiograph taken after the first operation.
F i g . 2 Radiograph taken before revision surgery.
At surgery, black joint fluid was collected, and the capsule and the periprosthetic tissue were diffusely stained with a black material. The metal ball wore out like a rugby ball (Fig. 3) and the femoral component was loosened. There was no evidence of infection. Geometry of metal ball and bearing insert was analysed by a coordinate measuring machine (BHN305, Mitutoyo Co.). Surfaces of retrieved metal ball and outer head were examined using scanning electron microscopy (SEM). Debris sticked to the bearing surface of PE insert was analysed by energy-dispersive analysis of X-rays (EDAX). Metal particles were extracted from the joint fluid by tissue digestion [6]. The retrieved metal particles in SEM photographs were measured by a computerized image analyzer and the size was estabhshed using the equivalent circle diameter (ECD). The levels of I l - l a , II-ip, 11-6, TNF-a in the joint fluid were measured by radioimmuniassay. RESULT S Volmetric wear of the metal ball was 729mm3. and that of the PE insert was 236mm3. SEM of the metal ball represemted many sharp scratches (Fig. 4). SEM of the bearing surface of PE insert represented roughened surface with two kinds of debris. EDAX revealed that they were
F i g . 3 Severely worn metal ball.
F i g . 4 SEM of the surface of metal ball.
Catastrophic Wear of Metal Ball of Bipolar Hip Prosthesis:H. Ohashi et al. 327 >K4 .
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metal and alumina (Fig. 5, 6). Mean BCD of the metal particles was 0.99–0.86 ;^m, and the size distribution was shown in Fig. 7. Histologically, metal debris as well as PE debris were diffusely observed in the periprosthetic tissue (Fig. 8). The levels of I l - l a , II-ip, 11-6 and TNF-a in the joint fluid were 42.5 ng/ml, 90.5 ng/ml, 3040.3 pg/ml and 73.0 pg/ml, respectively. DISCUSSIO N AN D CONCLUSION S Alumina ceramic has good biocompatibility, while alumina tends to break to small fragments due to their brittleness. Hardness of alumina ceramic fragments is greater than that of Co-Cr alloy, thus alumina ceramic fragments can rasp the surface of metal ball. From the observation of retrieved prosthesis, the mechanism of the catastrophic wear in this case is considered to be an abrasive wear due to alumina fragments sticked to the bearing surface of PE insert. Recently, sub-micron particulates are considered to play an important role in periprosthetic osteolysis [5,7], and bone-implant interface around failed joint prostheses was reported to contain high level of cytokines especially in regions with radiographic evidence of bone loss [8]. Experimentally, Co-Cr particles were reported to induce proliferation of macrophages [9], and stimulate cytokine production [10]. In this case, most metal particles were sub-micron in size and high levels of Il-la, II-1 (3,11-6 and TNF-a, that are associated with bone resorption, were detected in the joint fluid. These results indicated that the fracture of alumina ceramic screws brought about catastrophic wear of metal ball. The abundant sub-micron metal particles stimulated the production of cytokines that supposedly accelerated the progress of prosthetic loosening. From this point of view, we concluded that the use of alumina ceramic screws was contraindication in prosthetic replacement.
REFERENCES 1. Kempf, I. and Semlitsh, M., Arch. Orthop.TraumaSurg., 1990, 109, 284-287. 2. Matsuda, Y , Yamanuro, T., Kasai, R., Matsusue, Y. and Okumura, H., J. Arthroplasty,1992, 7S, 439-445. 3. Watanabe, M., Okumura, H., Kihara, Y. and Shibata, T., Arch. Orthop.TraumaSurg., 1993, 112,296-298. 4. Chiba, J., Rubash, H., Kim, K.J. and Iwaki, Y., Clin. Orthop.,1994,300,304-312. 5. Shanbhag, A.S., Jacobs, J.J., Giant, T.T., Gilbert, J.L., Black, J. and Galante, J.O., J. Bone Joint Surg., 1994,76-B, 60-67. 6. Yamac, T., Kobayashi, A., Bonfield, W., Kadoya, Y., Freeman, M A R . , Scott, G. and Revell, RA., Transactionsof Fifth WorldBiomaterialsCongress,1996, 861. 7. McKellop, H.A., Campbell, P., Park, S.H., Schmalzried, T.P., Grigoris, P., Amstutz, H.C. and Sarmiento, A., Clin. Orthop.,1995,311, 3-20. 8. Chiba, J., Schwendeman, L.J., Booth, RE., Crossett, L.S. and Rubash, HE., Clin. Orthop., 1994,299,114-124. 9. Howie, D.W. and Vernon-Roberts, B., Clin. Orthop.,1988,232, 244-254. 10. Dowd, J.E., Schwendeman, L.J., Macaulay, W., Doyle, J.S., Shanbhag, A S . , Wilson, S., Herdon, J.H., Rubash, HE., Clin. Orthop.,1995,319, 106-121.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
ANTIBACTERIA L PROPERT Y OF Ag-DOPE D CALCIU M PHOSPHAT E COMPOUND-CELLULOS E COMPOSITE S K. Okada\ Y. Yokogawa^ T. Kameyama^ K. Kato^ Y. Kawamoto^ K. Nishizawa^ F. Nagata^ and M. Okuyama^ ^R & D Center, NGK Spark Plug Co., Ltd., 2808 Iwasaki, Komaki, Aichi 485 Japan ^Bioceramic Laboratory, National Industrial Research Institute of Nagoya, 1-1 Hirate-cho, Kita-ku, Nagoya 462 Japan ^Laboratory of Bioorganic Chemistry, National Industrial Research Institute of Nagoya, 1-1 Hirate-cho, Kita-ku, Nagoya 462 Japan ABSTRAC T Ag-doped calcium phosphate compound (CP)-cellulose fiber composites were prepared by immersing CP-cellulose fiber composites into AgNGs aqueous solution after depositing CP on the cellulose fibers by soaking Ca(0H)2-treated phosphorylated cellulose into 1.5xSBF(Stimulated Body Fluid). The composite obtained became dark-brown from white gradually as Ag"*^ ion concentration in the AgNOa aqueous solution increased. The decrease of Ag* ion concentration in the AgNOa aqueous solution was also observed by ICP analysis and the broad XRD peak due to CP was shifted slightly for Ag-doped CP-cellulose fiber composites. It appeared from these that Ag was doped into the CP lattice. The antibacterial property of the composites was investigated by using Bacillus subtilis. As the amount of doped-Ag increased, the growth of Bacillus subtilis was inhibited. INTRODUCTIO N Hydroxyapatite (Caio(P04)6(OH)2); HAp) has been used as bone substitutes and dental implants because of its similar structure as the mineral phase in bone and teeth and has a high affinity for living bone. HAp has been also used as an adsorbent for high performance liquid chromatography column because of its adsorptive properties for virus, bacteria, and protein[l]. On the other hand, metals such as silver, copper, and zinc were well-known to have antibacterial property and antibacterial products have been being fabricating by mixing these metals-doped powders such as zeolite and HAp with fibers or resins. In the previous paper, calcium phosphate compound (CP) growth on cellulose fibers phosphorylated in 1.5xSBF was studied[2]. CP-cellulose fiber composites doped with these metals in the CP layer are expected to be useful as a virus and bacteria adsorptive filter with antibacterial property. In this present study, Ag-doping was attempted into CP layer of CP-cellulose fiber composites, and their characteristic and antibacterial property were evaluated. 329
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EXPERIMENTA L Deposition of Calcium Phosphate Compound Phosphorylation of cellulose fibers (Cotton Ciegalfi;100 % cotton) was carried out following the method described in an earlier report[2]. 16 pieces of cellulose fibers (ca. 5.5 g) were placed into a round-bottomed flask with 40 g of urea and 250 mL of DMF, and heated up to 120*0 with mechanical stirring under N2 atmosphere, and then approximately 32.0 g of 85 % H3PO4 in 100 mL of DMF was added into the solution. The temperature of the reaction mixture was increased to 155*0 and refluxed for 1 h. The reaction mixture was then left to cool under N2 gas flow, and the cellulose fibers were filtrated and washed thoroughly in water to rinse out excess phosphoric acid. The phosphorylated cellulose fibers were soaked without stirring in a saturated solution of ca. 500mL of Ca(0H)2 (pH = 12.5) in 1000 mL closed plastic bottle for 8 days. The Ca(0H)2 solution was renewed every 4 days. After soaking, the fibers were washed with pure water, and dried at 60*C under vacuum. 0.1 g of Ca(0H)2-treated phosphorylated cellulose fibers were The 1.5xSBF was prepared in immersed into the 200 mL of 1.5xSBF for 8 days at 3 6 . 5 0 . the same manner of an earlier publication[3]. The 1.5xSBF solution was basically prepared by dissolution of NaCl, KCl, CaCh, MgCh, NaHCOa, NaS04, and K2HPO4 in pure water with buffering agents "TRIS" ((CH20H)3CNH2) and HCl to keep the solution pH of 7.2 - 7.3 during soaking experiments. After soaking, the cellulose fibers were washed with pure water and dried at 6OO under vacuum. Ag-doping and Characterizatio n Ag-doping were carried out with two methods as follows. For method A, the CPcellulose fiber composites prepared were soaked into AgNOa aqueous solutions at ambient temperature for 1 day, where the Ag amounts were 0.5, 1, and 3 mol% against to Ca amount in the CP deposited. After soaking, the cellulose fiber composites were washed with pure water and then dried. For method B, the Ca(0H)2-treated phosphorylated cellulose fibers were soaked into 1.5xSBF containing Ag"^ ion at 3 6 . 5 0 for 8 days, where the amounts of Ag"*^ ion were 1, 5, 10, and 30mol% against to Ca^"*^ ion in 1.5xSBF prepared by using nitrate salts instead of chloride salts. The microstructure, the amount of Ag doped and crystalline phase were characterized with scanning electron microscopy (SEM), energy dispersive X-ray (EDX) and X-ray diffraction (XRD) analyses. The Ag concentration change in AgNOa solutions were studied by ICP analyzer. Evaluation of Antibacterial Property Antibacterial property was evaluated by following method. Yeast extract, polypepton, MgS04 and agar were added into O.IM phosphate buffer solution and dissolved by boiling. After sterilizing the solution with an autoclave, each 10ml of the solution were pipetted into sterilized dishes. The Ag-doped CP-cellulose fiber composites were put on the culture medium which 10^ pieces of Bacillus subtilis was set on, and then
AntibacterialProperty of Ag-Doped CP Compound-CelluloseComposites:K. Okada et al.
331
cultivated at 37 C for 24 h. After cultivation, the growth of Batillus subtilis around the sample and under the sample was evaluated. RESUL T AN D DISCUSSIO N The composites prepared by method A changed from white to dark-brown gradually as Ag^ ion concentration in the AgNOa aqueous solution increased. However, Ag amount in/on the composites was too small to be detected by EDX. The decrease of Ag"*" ion concentration in the AgNOa aqueous solution was observed by ICP analysis, and also the broad peak due to CP was shifted slightly for Ag-doped CP-cellulose fiber composites. It appeared from these that Ag was doped into CP lattice. Major difference was not observed in the microstructure such as morphology and size of primary grain. On the other hand, the weight increase of the composites prepared by method B were observed, but the amount of the compound deposited on the fiber decreased as the amount of Ag"*^ ion in the solution increased. Especially no deposition of CP was observed when the solution including 30mol% Ag"*^ ion was used. The color change to dark-brown and the shift of XRD peak due to CP were not observed differently from method A. ICP analysis didn’t show any changes in Ag concentration in the solution. It was considered to be difficult to prepare Ag-doped CP-cellulose composite by using l.SxSBF including Ag^ ion. Antibacterial properties were evaluated for Ag-doped CP-cellulose fiber composites prepared by method A, CP-cellulose fiber composite, and cellulose itself. In Figure 1, the inhibition of the growth of Bacillus subtilis was indicated for Agdoped CP-cellulose fiber composite, while it was not observed for cellulose and CPcellulose fiber composite. Also as the amount of Ag doped into CP increased, the inhibition of the growth of Bacillus subtilis was remarkable around the sample as well as under the sample.
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Tinhibited area l Figure 1 Photographs of growth of Bacillus subtilis around the samples (a) Cellulose, (b) Calcium Phosphate Compound-Cellulose Composite, (c) Agdoped Calcium Phosphate CompoundCellulose Composite SUMMAR Y Ag could be doped into calcium phosphate compound (CP) deposited on cellulose fibers by immersing CP-cellulose fiber composites into AgNOa solution and Ag-doped CPcellulose fiber composites had antibacterial property for Bacillus subtilis qualitatively. It is believed that these composites can be applied to the filter which have antibacterial property.
REFERENCES l.Tsuru S., Shinomiya N., Katsura Y., Uwabe Y., Noritake M. and Rokutanda M., BioMedical Materials and Engineering, 1991, 1, 143-147 2.Mucalo M.R., Yokogawa Y., Toriyama M., Suzuki., Kawamoto Y. and Nishizawa K., Journal Material Science, Material in Medicine, 1995, 6, 658-669 3.Li P., Otsuki C , Kokubo T., Nakanishi K., Soga N., Nakamura K. and Yamamuro T., Journal Material Science, Material in Medicine, 1993, 4, 127-131
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
WEA R BEHAVIOU R OF POLYETHYLEN E CUP AGAINS T 28mm ALUMIN A BAL L IN TOTA L HI P PROSTHESE S H. Oonishi. N. Murata. S. Kushitani, 8. Wakitani. K. Imoto. Y. Iwaki. and N. Kin. Department of Orthopaedic Surgery. Artificial Joint Section and Biomaterial Research Laboratory. Osaka-Minami National Hospital. 2 - 1 . Kidohigashimachi. Kawachinagano-Shi. Osaka. 586. JAPAN
ABSTRACT The wear rate of the U. H. M. W. polyethelene cups in combination with 28 mm alumina femoral head was measured on the radiographs whithout any complications and on the retrieved cups due to slight loosening of the p r o s › theses or due to late infections between bone and components. In both cases, the thicker the polyethylene cups, the lower the wear rate. The average wear rate of the cups of 7 and 8 mm thickness was about twice of that of 10 and 11 mm. From these results, the thickness of the polyethylene cups have to be used more than 11 mm. KEYWORD S wear of polyethylene cup. polyethylene cup thickness, total hip p r o s t h e › sis, alumina head INTRODUCTIO N We reported previously that the wear rate of the cup on 28 mm metal head(T-28. stainless steel head ball) was 2.5 times of that on 28 mm alumina head (Bioceram) [1]. The objective of our study was to find the relationship between the wear and the polyethylene cup thickness of cemented alumina ball total hip prostheses from the radiographs and the retrieved cups. We reproted previously the effect of the polethylene cup thickness of 7 to 9 mm to wear on the radiographs [1]. In this sutdy the polyehtylene cup thickness of 7 to 11 mm was compared on the radiographs and on the retrieved cups. CU P WEA R ON THE RADIOGRAPH S 1) Materials 111 joints in 102 cases were considered suitable for inclusion in this study. 14 joints in 13 cases were in male and 97 joints in 89 cases were in 333
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female. All cases were secondary osteoarthritis due to dysplasic acetabulum. 93 cases were unilatereal and 9 cases were bilatereal. 2) Methods The X - r a y were taken of A - P views of both hips in standing position (weightbearing). We used the decrease in thickness of the UHMWPE cup as our estimate of wear. Observation periods were from 1 month after s u r › gery and the longest year after surgery. For measurement on the radiographs, we used backlit-type digitizers with 20urn resolution, 5 x magnification view› ing loupes, and specially designed angle scales (0. 2 mm graduation). Data collation and analysis wear via computer. In our previous studies of the measurement of the wear of the polyethylene cups on the radiographs, we found the tendency that the thicker the cups, the lower the wear of the cups. The similar results were reported in hip simulator tests by Saikko [2]. T h e r e › fore, in this study the relationships between the cup thickness and the wear were investigated. Volumetric wear rate was calculated from linear wear rate using a system developed by J. Michael Cabo et al [3]. 3) Results The linear wear rate and the volumetric wear rate of each cases were shown in the figure 1 and 2 . The wear rate of each cases were scattered very widely. The average wear rate of the cups of 7 and 8 mm thickness was about twice of that of 10 and 11 mm. On the whole, the thicker the cup, the lower the wear rate. r
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Wear Behaviour of PE Cup Against 28 mmAlumina Ball In Total Hip Prostheses:H. Oonishi et al. 335
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CUP WEAR ON THE RETRIEVED CUPS 1) Materials T h e d e c r e m e n t of t h e cup t h i c k n e s s , including w e a r and c r e e p d e f o r m i t y , of Bioceram p r o s t h e s e s w a s m e a s u r e d on t h e r e t r i e v e d prostheses. The retrieved prostheses were due to slight loosening of the stem, cup or both, or due to late infection between bone and components. Prostheses damaged by other than wear by femoral head were excluded from this study. 2) Methods Generally, the inner surface of the retrieved cup has two spherical s u r › faces. The distance between two centers was defined as t h e length of t h e femoral head movement. In this case, as the initial wear, which was extremely higher than the steady steady state wear, could not be excluded, the wear rates included the initial wear in the steady state wear rate. 3) Results Relationships between linear and volumetric wear rates and cup t h i c k › ness were shown on the figure 3. The wear rates of the cups of 7 mm and 8 mm thickness were almost the same. The wear rates of the 7 and 8 mm cup thickness was twice of that of the 11 mm cup thickness. The thicker the cup. the lower the wear rate. DISCUSSIO N In the case on radiographs, if one case of the 44 mm in thickness and two cases of 50 mm in thickness, which showed extremely higher wear rates, were excluded from the population, the average line on the on the graph will get nearer to slow sloping or straight line. In the case on retrieved cups, the
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8 9 10 SOCKET THICKNESS(mm)
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Fig. 3 ; Linear and volumetric wear rate of polyethylene cups on the retrieved prostheses. Relationships between wear rates and cup thickness. average line on the graphs showed a slow sloping curve. In both studies on wear measurement on the radiographs and on retrieved prostheses, similar t e n d › ency and reasonable results were obtained in relationships between cup t h i c k › ness and wear rate of the cups. As a result, on the whole, the thicker the polyethylene cups, the lower the wear rate. The average wear rate of the cups of 7 and 8 mm thickness was about twice of that of 10 and 11 mm. The same tendency was reported in hip simulator test using water lubricant by Saikko [2]. However, the wear rate on the retrieved cups was higher by 50% than that on the radiographs. Because, in the retrieved cases, as the prostheses were not removed from the patients without any complications, but removed from the patients suffering from loosening of the components or late infections, higher wear rate was supposed to be found. Moreover, in the case on the r a › diographs, the initial wear, which was extremely higher than the steady state wear, was excluded However, in the case on the retrieved cups, as the initial wear was included, the wear rate in the case on retrieved cups became higher than that in the case on the radiographs. From these results, the thickness of the cups must be used more than 11 mm.
REFERENCES 1) OONISHI, K , TAKAYAMA, Y.. CLARKE I. C. and JUNG H.; J. of LongTerm Effects of Medical Implants. 1992. 2(1). 3 7 - 4 9 . 2) SAIKKO. v.. ; Acta Orthop. Scand. 1995. 66. 501-506. 3) CABO. J. M, ; J. of Bone and Joint Surgery. 1993, 7 5 - B ( 2 ) , 2 5 4 - 2 6
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
IN VITR O CEL L BEHAVIO R OF OSTEOBLAST S ON PYROS T BON E SUBSTITUT E J.S. Sun^ , F.H. L i n ’ , Y.H. Tsuang \ Y.S. Hang \ C.Y. Hong \ and H.C. Liu’ ^ Department of Orthopedic Surgery, National Taiwan University Hospital, and ’ Center of Biomedical engineering. College of Medicine, National Taiwan University, Taipei, Taiwan.
ABSTRACT
We have, elucidated the in vitro cell behavior of osteoblasts on Pyrost bone substitute. Using primary culture of rat osteoblasts, the changes in cell morphology on the surface of Pyrost bone substitute were studied. At 1 hour, 3 hours and days 1, 3, 7 after layering, the cell behavior was observed with SEM. The processes of trypsinized osteoblasts adhesion and spreading on Pyrost bone substitute consisted of: 1). contact of rounded osteoblasts with the Pyrost substrate; 2). attachment of osteoblasts at point of contact; 3). centriftigal growth of filopodia; 4). flattening and spreading of the osteoblasts on the Pyrost substrate; 5). division and growth of osteoblasts; 6). suspension of the osteoblasts across the pores by their processes. This result demonstrated that Pyrost can form a physico-chemical bond with osteoblasts. The Pyrosts bone substitute can support both attachment and proliferation of osteoblasts. KEYWORDS : pyrost bone, osteoblasts, adhesion, spreading.
INTRODUCTION
Approaches to bone regeneration for the treatment of various clinical conditions, such as fi-actures with bone loss, bone infections or bone tumors, involve the use of autogenous grafts or allografts [1]. Autogenous cancellous bone is the most effective bone graft material to date, but it also has drawbacks, including donor site morbidity and limited availability, especially in children [2]. The advantage of allografts over autografts lies in better availability of supply and their ability to be used for reconstructing large bone defects [1]. The major disadvantages of allogeneic bone include disease transmission and the graft’s tendency to elicit an immune response that can lead to high failure rates [3]. Pyrost bone substitute has been shown to be a promising orthopedic biomaterial. When used as a bone graft substitute, bony ingrowth into the implants without any adverse reaction can be demonstrated [4]. When implanted into bone, Pyrost can form a physico-chemical bond with bone tissue. However, little is known about the mechanisms responsible for the osteogenesis that occurs between bone and Pyrost bone substitute. The in vivo as well as in vitro bone formation are closely associated with the behavior of the cells. The formation and deposition of bone directly on to the implant require a surface that is not only non-toxic but also allows or favors the cell behavior [5]. 337
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We therefore initiated studies of the morphology of osteoblasts to test the in vitro cell behavior on the Pyrost bone substitute. MATERIALS AND METHODS
Sequential digestion of newborn Wistar-rat calavria was performed by using a modification of the methods described by Wong and Cohn [6]. Briefly, the dissected calvaria were sequentially digested with collagenase (180 U/ml, Sigma Co.) in solution A \yith EDTA. The cells released after last two treatments were harvested by centrifugation and resuspended in culture medium. The Pyrostfi (Oscal AG, Swiss) bones (10 x 5.0 x 5.0 mm) were placed in sterile dishes. Confluent rat osteoblast cultures were seeded into each well on top of the implant materials at a density of 3 x 10’* cells/cm^ The culture media used was DMEM supplemented with 10% fetal calf serum (Gibco, UK) and antibiotics (penicillin G sodium 100 units/ml -streptomycin 100 mg/ml, Gibco, UK). The dishes were incubated at 37 C in an atmosphere supplemented with 5% CO2 and fed with complete changes of medium twice a week. The day of plating was considered as the zero day of culture. The test samples were removed from wells at 1 hour, 3 hours and days 1,3,7, fixed in 3% formaldehyde in 0.1 M PBS buffer (pH 7.4). For electron microscopic examination, the Pyrost bone substitute blocks were fixed, dehydrated and critical dried. Specimens were sputter-coated with gold and examined by scanning electron microscope. RESULTS AND DISCUSSION
After layering on the Pyrost, the scattered round-shaped osteoblasts settled on the substratum with the proteinaceous sheets within 3 hours. One day after layering, flattening of some osteoblasts were visible. After 3 days in culture, the cells exhibited close contact with each other via filopodial processes. The surface of Pyrost was coated by an almost complete layer of osteoblasts by the day 7. The cells initially repopulate the Pyrost bone substitute by settling out of suspension, attach to the available surfaces provided, and then give rise to the final populations by mitotic expansion. In vivo, this condition is mandatory for osteogenesis to occur in an implanted material without interposition of fibrous or granulation tissue [5]. In this series, after trypsinization, the cells appeared spherical to ovoid in shape. The population consisted of cells with smooth, rounded surface, and surface that possessed numerous ’bleb’-like vesicular protuberances within 1 hours after layering (Fig. 1). Some cells showed smooth surface without microvilli or blebs. The numerous foldings and blebs on the surface of harvested cells are to accommodate the excess surface membrane as the cells round up from the flattened state in response to trypsin treatment [7]. The major events in the process of adhesion and spreading of these cells seem to be attachment of the cell to the substratum, radial growth of filopodia, cytoplasmic webbing and the resultant flattening of the cell. When cells layered on Pyrost bone substitute after 3 hours, the cells adhered very firmly to the surface. This was affected by microvilli-like cell processes. Their growth occurred only at the point of contact with the Pyrost (Fig. 2). It is likely that the microvilli-like projections were formed all over the cell surface but were later withdrawn except at point of contact with the substratum. The first step in cell spreading constitutes the centrifugal growth of microvilli-like processes that elongated into filopodia (Fig. 3). It appears that these spherical tips of filopodia may play a direct role as specialized structure of attachment to the substratum [7].
In Vitro Cell Behavior of Osteoblastson Pyrost Bone Substitute:J-S. Sun et al.
339
I’ig. 1 S E M exaJTiination of trypsniharvested osteoblasts fixed vvilbin 1 hour after layering. (Bar: 2,6 |Lim)Fig. 2 SHM examination oftrypsijiharveslcd osteoblasts fixed at 3 hours after layering. Spherical cells with inconspicuous microvilli-]ike projections (Bar: 1.9 ^irn). Fig. 3 SUM examination of trypsinharvestcd osteoblasts fixed at 3 hours after layering. The contact area with long filopods- (Bar: 2.S ^ m ) .
Figs. 4-7 SEM examination of trypsin-harvested osteoblasts. Fig. 4 One day after layering: Flattened cell with cytoplasmic webbing (Bar: 4.4 |Lim). Fig. 5 Three day after layering. Cells with the long filopods (Bar: 5.0 fim). Fig. 6 Three day after layering. The division and growth of osteoblasts (Bar: 3.9 |Ltm) Fig. 7 Seven day after layering. Cells lay densely on Pyrost; and cells spanned the pores apparently by first expanding explorative filopods across the macropore. (Bar: 13.0 jam).
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Close contacts made by the advancing filopodia is followed by extensive spreading as a thin web between the filopodia (Fig. 4). By the first day after layering, cells have flattened considerably. It is reasonable to conclude that the strength of adhesion of a cell settling on the Pyrost bone substitute would increase progressively as the area of its contact increases (Fig. 5). Some of the blebs still persisted in the surface of the flat cell. The growth of filopodia followed immediately by cytoplasmicwebbing seems to be the pattern of origin of the ruffling membrane. The origin of ruffling membranes may indicate the final stage of spreading and the beginning of the cell movement [7]. Following the cell attachment, division and growth of cells occurred at 3 days after layering (Fig. 6). During mitosis a reversal of these events occurs by cytoplasmic de-webbing; the cell becomes spherical. After mitosis, the two daughter cells become flat again by cytoplasmic webbing. At the 7ih day, proliferating cells lay flat and in close contact with the Pyrost surface (Fig. 7). On the other hand, cells appeared suspended across the pores by their processes (Fig. 7). As a result, the cells were not in close contact with the underlying substratum. The formation and deposition of osteoblasts directly on to the Pyrost surface imply that it is not only non-toxic but also allows or favors the osteoblasts behavior. SUMMARY
Surface reactivity may be a key factor in determining the morphological and functional responses observed during the osteoblast-substrate interactions. Pyrosts bone substitute is considered osteo-compatible [4]. This experiment has defined that Pyrosts bone substitute is not only support osteoblasts attachment but it also allows proliferation of the cells. It should be emphasized that these different stages are not discretely separable but are different phases of a contiguous process. This is not a synchronous cell population so variation in the duration of these phases exists and the degree of overlapping of these events is observed. Complete interpretations of these events, however, require further investigation of both morphological and functional responses of osteoblasts to Pyrost bone substitute, in particular with human osteoblasts. ACKNOWLEDGEMENT
S
The authors sincerely appreciate the National Science Council (ROC) for their financial support to accomplish the research. REFERENCES 1. Mankin, H.J., Gebhardt, M.C., and Tomford, W.W., Orthop. Clin. North Am., 1987, 18, 275289. 2. Begley, C.T., Doherty, M.J., Hankey, D.P., and Wilson, D.J., Bone, 1993, 14, 661-666. 3. Bos, G.D., Goldberg, V.M., Zika, J.M., Heiple, K.G., and Powell, A.E., J. Bone Joint Surg., 1983, 65A, 239-246. 4. Katthagen, B,D. 1986 Bone regeneration with bone substitutes: An animal study. Springer, Berlin Heidelberg New York, 29-50. 5. Bagambisa, F.B., and Joos, U., Biomaterials, 1990, 11, 50-56. 6. Boonekamp, P.M., Kekkelman, J.W., Hamilton, J.W., Cohn, D.V., Jilka, R.L., Proc. Kon. Acad. Wet. B., 1984, 87,: 371-384. 7. Rajaraman, R., Rounds, D.E., Yen, S.P.S., Rembaum, A., Exptl. Cell Res., 1974, 88, 327-339.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TH E EFFICAC Y OF HYDROXYAPATITE-TRICALCIU M PHOSPHAT E FILLE R FO R BON E DEFECT S ASSOCIATE D WIT H HUMERA L PSEUDOARTHROSI S : COMPARISO N WIT H AUTOGENOU S ILIA C BON E GRAFT S Katsuji Suzuki and Mitsuko Yamada. Department of Orthopaedic Surgery, Fujita Health University School of Medicine. 1-98, Dengakugakubo, Kutsukake, Toyoake, Aichi, 470-11, Japan. ABSTRAC T The efficacy of a hydroxyapatite-tricalcium phosphate ( HAP-TCP ) AIICT for bone defects associated with humeral pseudoarthrosis was studied by comparing 7 patients treated with HAP-TCP and 7 who received iliac autografts. There WCTC no significant differences in age, additional injuries, and the non-union period between the HAP-TCP group and the iliac autograft groups. Tha-e were also no significant differences in the postoperative bone union time and range of motion ( ROM ) recovery time. KE Y WORD S hydroxyapatite-tricalcium phosphate humeral pseudoarthrosis iliac autograft bone union range of motion INTRODUCTIO N In patients with humeral pseudoarthrosis, it is necessary to resect and freshen the sclerotic bone and fibrous tissue, as well as filling bone defects. Iliac autografts are often used, but present problems due to the limited amount of bone that can be harvested and the occurrence of symptoms at the donor site. We compared the effectiveness of hydroxyapatite-tricalcium phosphate ( HAP-TCP ) and iliac autografts for achieving bone union in patients with humeral pseudoarthrosis. SUBJECT S Fourteen patients with hum^al pseudoarthrosis and bone defects requiring filling underwent surgery at this department between 1987 and 1995 ( 9 males and 5 females ; average age : 35.1 years; range : 7 - 75 years ). HAP-TCP group HAP-TCP was used in 7 patients ( 5 males and2 females; average age : 35.4 years ; range : 7-69 years ). The avaage period between injury and surgay for psedoarthrosis was 70.3 + 26.6 ( M – S E ) months and the range was 12 to 240 months. Two of them had shaft pseudoarthrosis, 2 had lateral condylar non-union, 1 had comminuted concfylar non-union, 1 had neck non-union, and 1 had supracondylar non-union. In addition, there was tardy ulnar nerve palsy in 1 patient, radial nerve palsy in 1, axillary nerve palsy in 1, and osteoporosis in 1. 341
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Iliac autograft group Iliac autografts were used in 7 patients (4 males and 3 females ; average age :34.9 years; range 11 - 75 years). The average period until surgery for non-union was 108.0–59.5 ( M – S E ) months and the range was 12 to 456 months. Three of these patients had lateral condylar non-union, 2 had shaft non-union, 1 had medial condylar non-union, and 1 had supracondylar non-union. In addition, there was tardy ulnar nerve palsy in 2 patients and osteoporosis in 1. METHOD S In both groups, we measured the time until bone union was apparent on X-ray (bone union time) and the time required for the range of elbow joint motion to recover to greater than 80% of that on the healthy side ( ROM recovery time ).
Figure 1. Case 1. A 30-year-old man from the HAP-TCP group. 1-A: Pseudoarthrosis of the shaft of the right humerus 34 months after injury. Ender nailing was done twice, but pseudoarthrosis of the humeral shaft was developed 1-B: Four weeks after surgery. The bone defect at the site of pseudoarthrosis was filled with an HAP-TCP block and granules. 1-C: One year after surgery, bone union is good.
Efficacy of Hydroxyapatite-TricalciumPhosphateFiller for Bone Defects: K. Suzuki and M. Yamada 343
RESULT S The bone union time was 23.4–4.6 ( M – S E ) weeks in the HAP-TCP group and 19.4–3.6 weeks in the iliac autograft group (P=N.S.; Wilcoxon test). The ROM recovery time was 17.9–4.3 ( M – SE ) weeks in the HAP-TCP group and 13.9 – 1.4 weeks in the iliac autograft group (P=N.S.; Wilcoxon test). Postoperative infection did not occur in either group. Pain at the site of bone removal occurred in 3 patients from the iliac autograft group, while th^e were no complications in the HAPTCP group. CAS E REPORT S Case 1 was a 30-year-old man ( HAP-TCP group) with pseudoarthrosis of the right humeral shaft. He suffered a right humeral shaft fracture in a traffic accident and was operated on twice with Ender nails at another hospital. However, bone union was not obtained afto" 34 months. At our department, the Ender nails were removed and the sclerotic bone and fibrous tissue at the pseudoarthrosis were resected. Then an HAP-TCP block and granules were used to fill the bone defect and firm internal fixation was achieved with a titanium plate and screws. The bone union time was 24 weeks and the ROM recovery time was 6 weeks.
Figure 2. Case 2. A 34-year-old man from the iliac autograft group. 2-A: Pseudoarthrosis of the shaft of the right humerus 30 months afta* injury. 2-B: Immediately aftCT surgery. The bone defect at the pseudoarthrosis site was filled with an iliac bone block and cancellous bone chips. The bone defect at the site of intCTlocking nail removal was filled with HAP-TCP granules. 2-C: Twelve weeks after surgery, bone union is good.
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Case 2 was a 34-year-old man (iliac autograft group ) with pseudoarthrosis of the right humeral shaft. He suffered a right humeral shaft fracture after falling from a height, and was treated with an interlocking nail at another hospital. However, bone union was not obtained after 30 months. At our department, the pseudoarthrosis was resected and freshened, and the interlocking nail and a screw were removed. An iliac bone block and cancellous bone chips woe used to fill the bone defect, and firm internal fixation was achieved with a titanium plate and screws. The bone union time was 12 weeks and the ROM recovery time was 8 weeks. DISCUSSIO N HAP-TCP filler is a composite of hydroxy apatite (HAP) and tricalcium phosphate (TCP) which shows excellent biocompatibility and bone conductivity. TCP also fuses strongly with the surrounding bone because it acts as a bone substitute. Suzuki et al.,[l] packed traumatic bone defects of tibia with HAP-TCP fill^ and compared its efficacy with that of autogenous bone grafts. The weight-bearing recovay time was significantly shorter in the HAP-TCP group than in the autogenous bone group. Suzuki etal.,[2] also filled traumatic bone defects of the distal radius with the HAP-TCP filler and compared its efficacy with that of iliac autografts. Although osteoporosis was significantly more conmion in the HAP-TCP group, there was no significant dififoience in the ROM recovery time and grip powCT recovery time. In addition, the 1-year postop^ative rado-ulnar distance was significantly greater in the HAP-TCP group than in the iliac autograft group. HAP-TCPfillCTfiises directly with the surrounding bonefix)man early stage. TCP also forms a strong union with the surrounding bone by gradual progressive subsUtuUon, and HAP gradually increases in strength after grafting because of its excellentbone conductivity [3]. Thus, the final strength of HAP-TCP filler approaches that of normal cancellous bone. When autogenous bone is used, the graft is gradually substituted by new bone and its mechanical strength remainsreducedduring this process[4]. In the present stucfy, tha^ were no significant differences in age, complications, and non-union period between the HAP-TCP and iliac autograft groups, and there were also no significant difli^Tences in the bone union time or ROM recovery time. CONCLUSIO N There waie no significant differences in bone union time and ROM recov^y time between the HAP-TCP and iliac autograft groups. It was concluded that HAP-TCP was an effective filler for bone defects associated with humeral pseudoarthrosis.
REFERENCES 1. Suzuki, K. and Kurabayashi, H. In: BioceramicsVolume 7, Butterworth-Heinemann, Oxford 1994, 435-440. 2. Suzuki, K., Yamada, M., Yamamoto, K. and Muramatsu, K. In: BioceramicsVolume 8, Pergamon, Oxford 1995, 225-229. 3. Hon, M., Munemiya, M., Takahashi, S., Sawai, K., Niwa, S., Tagai, H., Kobayashi, M., Ono, M. andTakeuchi, K. Cent.Jpn, J. Orthop.Traumat.1984, 27, 2133-2135. 4. Nakamura,S. KitazatoIgaku 1998, 18, 406-419.
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EXPERIMENTA
L STUD Y OF APATIT E CEMEN T INCLUDIN G CISPLATI N
Y. Tahara, Y. Ishii, S. Sasaki, I. Takano, and K. Ohzeki Department of Orthopaedic Surgery, Kyorin University School of Medicine. 6-20-2 Shinkawa, Mitaka-shi, T(*yo 181, Japan
ABSTRAC T An implant consisting of calcium phosphate cement and cisplatin (CDDP) in order to apply a concept of drug delivery system to the therapy for malignant bone tumor and to maintain an anticancer drug at higher local concentrations with simultaneous complementation of structural disadvantages. In investigation the slow-realing activities in vitro and the influences to other organs and tissues in an implant group and a CDDP systemic dose group in vivo(with Japanese white male rabbits), we concluded that the implant containing 10% CDDP was ideal. KE Y WORD S : Cis-Diamminedichloroplatinum, Ceramic, I>rug Delivery System OBJECTIVE S There are not a few adverse reactions or invasion into oth^ organs and tissues after the existing therapeutic methods for malignant bone tumor. Therefore, we prepared a ceramic implant containing an anticancer drug in order to maintain the anticancer drug at a higher local concentration and simultaneously supplement local structural disadvantages after ^plication of a concept of drug delivery system. Slow-releasing actvities and influences on other organs and tissues were investigated in vitro and in vivo using the implant thus prepared. METHOD S An implant, ISOmg in weight, Smm in diameter and 4mm in height, was prepared with calcium phosphate cement, cicplatin powder and a consolidating solution. Cisplatin was contained in terms weight ratios of 0,5,10 and 20%. For the systemic administration, Randa Injectable manufactured by NIPPON KAYAKU Co., Ltd. containing CDDP by 0.5mg/ml. Japanese white male rabbits weighing about 3kg were used as experimental animals. In vitro slow-releasing experiments, platinum(Pt) in CDDP was determined under the conditions to allow it stand in a thermostat at 37A6 in 100ml of phosphate buffer at pH7.4 at each concentration. In in vivo experiments, each implant was embedded in the distal epiphysis of rabbit femur and changes in body weights and Pt concentrations in the bone marrow surrounding the implant, the bone marrow 1cm distant from the implant, the kidney and the liver were determined. Changes in body weights and Pt concentrations in the bone marrow, the kidney and the liver were determined with a single standard dose of 3mg/kg in humans in a systemic dose group. 345
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RESULT S In vitro slow-releasin g experiments , with CDD P content s of less than 10%, cumulative slowreleasing rate after 4weeks was about 30% while slow-releasin g velocity was 0.03mg/da y while with a CDD P content of 20% it was 0.1mg/day(Fig . 1 and 2). In vivo the body weight remarkably decrease d in the systemic dose group. With a CDD P content of 20% in which the slow-releasin g velocity and the cumulative slow-releasin g rate were both favorable in vitro, Pt concentration s in the local bone marrow were much higher in the bone marrow and high in other visceral organs and tissues. With a CDD P content of 10%, Pt concentration s in the local bone marrow were higher than the systemic dose group while Pt concentration s in other visceral organs and tissues were lower(Fig. 3,4 and 5).
70 n 60
2 on
I
4)
1
i
50 H
40 30 H
1.0
2.0
3.0
Time(week)
Figurel.Cumulative slow-releasin g rate from the implant.
Figure2. Slow-releasin g velocity from the implant
9000-
—0—
10%Inplant
^
i
— • — 20%Inplant
80007000-
e
1
60005000-
7
4000-
B e u
3000-
a.
200010002
Time(week)
Figure 3. Pt concentratio n in the tissue in a CDD P systemic dose group.
3
4
Time(week)
Figure 4. Pt concentratio n in the bone marrow in an implant -embedding group.
ExperimentalStudy of Apatite CementIncluding Cisplatin: Y. Tahara et al.
347
(fig/tissue-g ) 10-
3
4
Time(week)
Figure 5. Pt concentration in the tissue in an implant-embedding group. DISCUSSIO N The body weight decreased in a wider range in the systemic dose group than in the implantenbedding group so that it might give influences on the gastrointestinal tract. In comparison of Pt concentrations between at the local bone marrow and in the bone marrow 1cm distant, CDDP was considered to be released in a range of less than 1cm. Pt concentrations in the local bone marrow with an implant containing CDDP by more than 5% were more than 100 times higher than those in other visceral organs and tissues so that higher local antitumor effects could be expected under the conditions of few effects on other visceral organs. It is an ideal implant if Pt concentrations are higher locally but lower in other visceral organs and tissues. From the present experiments, an implant containing 10% CDDP is considered as ideal. REFERENCES 1. Uchida A., Shinto Y., Araki N., Ono K., Jpn. J. Cancer Chemother,. 1989,16,3231-3235. 2. Shinto Y., Uchida A., Araki N., Ono K., Jpn. J. Cancer Chemother,.1991,18,221-226. 3. Kitamoto K., Hamanishi C, Yoshii T., Tanaka S., J. Jpn. Orthop. Assoc,. 1994,68,S 1602.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) (D1997 Elsevier Science Ltd
IN VIVOEVALUATIO N OF SOL-GE L BIOGLASS* . BIOMECHANICA
L FINDINGS .
Wheeler DL, Hoellrich RG, McLoughlin SW, Chamberland DL, Stokes KE Oregon Health Sciences University, Orthopaedic Research Laboratory 3181 SW Sam Jackson Park Road, L477, Portland, OR 97201
ABSTRAC T Bioglass* (BG) has proven to be an effective bone graft material due to the apatite layer which forms on the surface of the glass, promoting bone formation. Sol-gel Bioglass*, which has greater porosity and surface area, accelerates apatite layer formation and degradability. The objective was to biomechanically evaluate bone formed within distal femoral cancellous bone defects filled with Bioglass*particulates (BG) and two compositions of sol-gel Bioglass* (SGI and SG2) compared to normal cancellous bone (NORM) using a rabbit model. Compressive modulus for the BG group was significantly greater than SGI at 4 and 12 weeks (p<0.05). However, at 8 weeks all moduli were equivalent. In general, the modulus of the regenerated bone within all graft materials was equivalent to normal cancellous bone. The BG modulus was higher at 4 and 12 weeks, which may be attributable to the hardness/density of the BG particles. KEYWORDS : Bioglass*, Sol-gel, in vivo, cancellous bone defect, biomechanical testing INTRODUCTIO N Some osseous defects and fiactures may require supplimentation to suRWrt and promote healing. Bioglass* (BG) has proven to be an effective graft material for oral and maxillofacial bone augmentation’-^. The success of BG is, in part, due to the apatite layer which forms on the surface of the glass, promoting bone formation^. Therefore, Sol-gel Bioglass* was developed which has greater porosity and surface area, accelerating the formation of the apatite layer and increasing degradability\ Sol-gel Bioglasses* are produced through a new drying method under humid conditions^ (compositions outlined in Table 1). The objective of this in vivo investigation was to mechanically evaluate the bone formed within a 349
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distal femur defect filled with Bioglass* particulates (BG) and two compositions of Sol-gel Bioglass* (SGI and SG2). Graft materials were compared to normal rabbit cancellous bone of the distal femur. MATERIAL S AND METHOD S Twenty-four NZW rabbits were randomly assigned to 3 groups based on the bone graft material used (BG, SGI, and SG2). Defects were surgically created consisting of 6 mm holes drilled through the femoral condyle from lateral to medial cortical shell using a low-speed, saline-cooled drill. The defects were filled with the assigned graft material. Distal femoral condyles without defects were used to assess normal bone compressive modulus. Animals were sacrificed after either 4, 8, or 12 weeks of healing. The distal femurs were harvested and bisected into medial and lateral condyle specimens. The lateral condyles were processed for histological analyses. The medial condyles were frozen at -20 C for biomechanical analysis. Specimens were thawed to room temperature prior to testing, the central bisected surface was mounted on a 1.5 mm thick plastic slide, and cut to a thickness of 8 mm. The defect area was identified using contact radiographs and macro-inspection. Compression strength of the cancellous graft was assessed using a servohydraulic testing machine (MiniBionix, MTS Corp., Minneapolis, MN). The specimens were placed on a flat stainless steel plate attached to the MTS load cell. A 3.15 mm diameter stainless steel pin was used to compress the center of the 6 mm diameter graft area to a depth of 6.5 mm at a rate of 0.1 mm/sec. Data was statistically analyzed using unbalanced block ANOVA at a significance level of p=0.05. RESULT S A compilation of the mechanical results are presented in Figure 1. Compressive modulus for BG was significantly greater than SGI at 4 weeks (p<0.05). At 8 weeks all moduli were equivalent. However, at 12 weeks the modulus for BG was again significantly greater than SGI (p<0.05). The histological results are presented in Figure 2. At 4 and 8 weeks the defects filled with BG had significantly greater bone ingrowth than those filled with SGI or SG2 (p<0.05). However, at 12 weeks the percent of the defects area filled with bone was significantly greater for the SGI- and SG2-augmented defects than the BG-filled defects (p<0.05).
In Vivo Evaluation of Sol-Gel Bioglass^ -Biomechanical Findings: D.L. Wheeleret al.
351
DISCUSSIO N The modulus of the regenerated bone within all graft materials was equivalent to normal cancellous bone after 4 weeks of healing and maintained mechanical competency through 12 weeks. The compressive modulus of BG-filled defects was higher than SGI -filled defects at 4 and 12 weeks. Although not statistically significant, a trend toward higher modulus for BGfilled defects versus SG2 and NORM was observed at 12 weeks. The higher modulus of BGaugmentated bone may be due to the relative hardness and density of the BG particles and less resorption of BG particles was observed compared to SGI and SG2 particles. By 12 weeks the SGI and SG2 particles had resorbed 36% and 45% (based on major axis), respectively; compared to less than 10% resorption of the BG particles. The ingrowth of bone within the cancellous defect was greatest for the BG-filled defects, fiirther enhancing the density of the defect area compared to SGI and SG2. However, the histological data at 12 weeks is counterintuitive. The amount of bone within the BG-augmented defect was significantly lower than that of SGI and SG2 as well as reduced from levels observed at 4 and 8 weeks. This data is disparate from other experiments using the same model^ and conducted in the same laboratory^. The research confirms that Bioglass* and Sol-gel Bioglass* are able to promote osseous filling of critical-sized cancellous defects without adverse tissue responses in a rabbit model. The newly formed bone matrix acheives and maintains mechanical competency equivalent to that of normal cancellous bone of the distal rabbit femur by 4 weeks postimplantation. Investigations are currently underway to confirm the long-term histological and mechanical competency of these cancellous graft materials using a rabbit model and a large animal model.
REFERENCES (1) J. Wilson et al., Biomaterials7:223-228,1987; (2) H. Oguntebi et al.,y. Dent.Res. 72:484-489, 1993; (3) L. Hench, J. Am.Ceram.Soc. 74:1487-1510,1991; (4) M. Pereira et al., J. Biomed.Mater.Res. 28:693-698, 1994; (5) H. Oonishi et &U Bioceramics8:137-144,1995; (6) D. Wheeler et al., J. Biomed. Mater.Res. In review ACKNOWLEDGMENT S The authors would like to thank US Biomaterials for their support of this project. Table 1. Composition of Sol-gel Bioglasses*
LJD 11
Surface area (IVf/g) |
Composition
Bulk Densify (g/cm^)
Pore diameter (D)
1
77S
1.23
40
389
SG 2 11
58S
0.99
86
207
4585
2.65
N/A
0.02
SG I
BG
1
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350 0 -r-
KEY a: BG > SG1 (p<0.05)
Treatment by Time
Figure 1. Modulus of Regenerated Bone within Defect
KEY
a: BG>SG1.SG 2 b: SG1,SG2>B G (p<0.05 )
BG
SGI
4 weeks
SG2
BG
SGI
SG2
8 weeks Treatment by Time
Figure 2. Percent Bone within Defect
BG
SGI
12 weeks
SG2
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
FIXATIO N OF HI P PROSTHESE S BY HYDROXYAPATIT
E COATIN G
G. Willmann CeramTec, Medical Product Division, P.O.B. 1149, D-73201 Plochingen, Germany
ABSTRAC T Hydroxyapatite (HA) is used as bulk material for bone grafting and as coating on any kind of metal implants. There are clinical results that prove that HA-coating is enhancing the bony ingrowth and fixation of implants. Aftitureproblem may be the higher costs of coated implants. HYDROXYAPTIT E Hydroxyapatite (HA) is a one of the various calcium phopahtes known. HA is bioactive (= osseoconductive). HA is available as a granule or bulk material for bone grafting and as coating on all kinds of implants. HA-coating is oflFering the opportunity to enhance the osseointegration of implants. MATERIAL S AND METHOD S For about 10 years HA-coating has been investigated for bone implantfixation.There are lots of papers about basic research work, animal studies, and clinical reports [1,2,3,4]. Often the results are very controversial. They rangefi-om"HA has a positive effect on bone-implant fixation" to the statement that HA-coating has no benefits at all. Today bioactive HA-coating is successfiilly used for dental implants (see figure 1) and for components in total hip replacement and total knee raplacement (e. g.figure2). In Central Europe far more than 100,000 stems and cups for hip arthroplaysty have been HA-coated. The following statements are based on more than ten years of experience with HA-coating of numerous dental implants and components for hip and knee arthroplasty. Table 1.
Recommendation for HA-coating on some standard biomaterials.
stainless steel
cp-titanium or
yes
yes
cobalt chrome titanium alloy yes
PE-UHMW alloys no
THE STAT E OF THE ART 353
carbon CFRP
no
no
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The mostly used process is the air plasma spraying process (APS). There are no technica l problems to coat whateve r metal biomaterial is used for implants, see table 1. It was proposed to coat polyethylen e (PE-UHMW ) for cementles s fixation of acetabula r cups made of polyethylenen . HA coating of polyethylen e (PE-UHMW ) and carbon fibrereinforce d plastic(-CFRP) is possible, but it is not recommended . It is assumed that the materials propertie s expeciall y biocompatibil y change when polymers are HA-coated. Due the coating process the substrates be heated up and the structure of the polymeres will degradate . Any kind of metal surface may be coated. It does not matter if it has a structure or not. One of the prereqmsue s is that the surface of the substrate has to be roughed before coating. The surface of an implant may be just roughened (see figuresI) or structurized , e.g. Liibecker spongiosa porous coating (beads), or with a mesh on it, see figures2. Good HA-coatings have a chemical composition , a content of mineral phases, and crystalinity that is according to the American standard AST M F 1185 [6]. There is still a discussion going on about standardization , e. g. [7,8] The thickness of the HA-coating may be in the range froittSO ymto 250 yaa[4]. HA-coating is always porous with interconnectin g pores [5]. The lower the thickness, the higher the strength of the HA-coating. Typical values for pull-ofl^ strength are > 35 MP a for a coating with a thickness of 200 \aa.There are no problems with delamination or fatigue which was proven by in vitro testing [9].
Figure 1.
Dental implant made of titanium, partially coated, diamete r 3.5 mm, thickness of coating 50 ym(Tiolox screw according to Hotz [10])
Fixation of Hip Prosthesesby Hydroxyapatite Coating: G. Willmann 355
Figure 2.
HA-coated tibia component (just to show that there is no limitation to coat structurized surfaces)
The resorption rate of HA-coating depends on the purity of HA. The HA-coating should befreeof tricalcium phosphate (TCP) and calcia (CaO). The resorption rate is extremly influenced by the Pn-value of the environment. Resorption is accalerated i case of an inflammation which causes a shift of the Pn-value and enhances the solubility of HA Animal tests show that HA-coating has the capability to enhance bony in-growth [3], i. e. HA is enhancing thefixationof any kind of implant [1,2,3]. Even more important: Clinical reports about long-term experiences in orthopedics became available [1,2]. They prove that an HA-coating is improving the osseointegration of coated implants. FUTUR E DEVELOPMENT S HA-coating may be used a s a carrier for BMP (= bone morpohogenetic protein) [11]. BMP is an osseoinductive material that induces bone transformation and is accelerating the osseointegration of an implant by far more than any bioactive material alone. BMP in combination with HA as a carrier is a very attractive aproach to solve the problem of osseointergation of implants. Despite all the good clinical results there is a discussion going on if the advantages of HA-coating discussed and hsted above justify higher prices for the coated componetsin total hip arthroplasty. The cost pressure in orthopedics may jeopardize the concept of HA-coated implants.
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SUMMARY HA-coating can be applied to all metallic implants known. Animal tests and clinical results prove that HA-coating is enhancing the bony ingrowth and thefixationof an implant. The cost pressure in orthopedics may jeopardize the concept of HA-coated implants. REFERENCES 1. Geesink, R. G. T., M. T. Manley (eds.) HydroxyapatiteCoating in Surgery^ Raven Press (1993) Orthopedic 2. Epinette, J.A., R. G. T. Geesink (eds.) HydroxyapatiteCoatedHip and Knee Arthroplasty^ Expansion Sientifique Francaise, Paris (1995) 3. Soballe, K. Acta Orthop. Scand. Suppl. No. 225 (1993) 4. Willmann, G. Brit. Ceram. Trans. 1996, 95. 212 - 216 5. Willmann, G. Interceram 1993, 42. 206 - 208 6. ISO / DI S 13779 Implants for surgery - Ceramic materials basedon hydroxyapatite (1996) 7. Horowitz, E., J. E. Parr (eds.) Characterization andPerformance ofCalcium PhosphateCoatingsfor Implants,STP 1196ASTM, Philadelphia (1994) 9. Willmann, G., H. Richter, and M. Wimmer, Biomed. Technik 1993 38. 14-16 10. Hotz, W. Zahnarztliche Praxis 1991 42. 254 - 256 11. Hartwig, C.-H., W. Kusswetter, and G. Willmann Abstract Nr. 32 XIX. Munchener Symposium f. experim. Orthopadie March 8, 1997 Munich
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
ACETABULA R RECONSTRUCTIO
N WIT H AN ARTIFICIA L BON E BLOC K
S. Yoshii, *M. Oka, **T. Yamamuro, **H. Iida,**Y. Kakutani, K. Ikeda,H. Murakami and **T. Nakamura The Department of Orthopaedic Surgery, Kansai Denryoku Hospital, Hukushima 2-1-7, Hukushima-ku, Oaska533 Japan. Phone: 6-458-5821, Fax: 6-458-6994 *The Department of Artificial Organ, The Research Center for Bioengineering, Kyoto University, Kyoto 606 Japan **The Department of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, Kyoto, 606 Japan KEYWORDS : glass-ceramics, acetabular dysplasia, reconstruction, artificial bone. ABSTRAC T Operations on the pelvis are useful in congenital dysplasia of the hip to increase the coverage d" the femoral head. These operations require relatively long time spica cast immobilization and avoiding weight bearing until the graft has become stable. It is a troublesome technical problem in the bony part of the operation. We have developed a block of glass-ceramic to solve these problems. A forty-one-year-old woman with bilateral acetabular dysplasia had severe pain in both hip joints. The implant was fixed on the lateral surface of the ilium just above the joint capsule. The patient had no immobilization and returned to her daily life without a cane 4 weeks after the operation. She had no pain of the operated hip joint one year after the operation. The hip joint had full range of motion. The Harris hip score was 51 points preoperatively and 93 points at one year postoperatively. Our procedure is unique in that we used a bioactive glass-ceramic block to restore the acetabular dysplasia and the result is excellent INTRODUCTIO N Persistent acetabular dysplasia may lead to progressive subluxation of the femoral head. It is widely accepted that inadequately supported subluxation of the hip often progress into a painful and disabling condition in the young adult. The diminished weight-bearing area and increased shear stress cause attrition of the hyaline cartilage, and ultimately degenerative arthritis and pain. Operations on the pelvis are useful in congenital dysplasia of the hip to increase the coverage d" the femoral head. They give adequate support to the subluxated femoral head and may provide lasting symptomatic relief [1-3]. Koenig 1891; Chiari 1955; Salter 1961). The shelf procedure is useful for acetabular dysplasia in which no other osteotomy will establish a congruous joint. In a classic shelf operation, the acetabular roof is extended laterally by a graft or the lateral cortex of the ilium. However, these operations require relatively long time spica cast immobilization and avoiding weight bearing until the graft has become stable . A thick augmentation requires an abundant bone graft. It is a troublesome technical problem in the bony part of the operation. We have developed a block of glass-ceramics to solve these problems and to make the shelf operation a fool-proof procedure. We made a block of the glass-ceramic to use it as an artificial bone. We 357
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Studied for ten years with rabbits and dogs, which proved that it has excellent bone-bonding capability [4, 5] and the artificial bone was excellent to reconstruct acetabular defect, MATERIAL S AN D METHOD S The patient A forty-one-year-old woman was initially seen three months after an gradual onset of pain in both hip joints. She was diagnosed to have bilateral acetabular dysplasia after physical and roentgenological exstmination. She had a conservative treatment with rest, cane and medication foe ten years. The pain of right hip joint aggravated gradually. She could walk two kilometers with severe pain and a limp. She decidedto have an operative treatment. She wanted to have implantation of an artificial bone block because she wanted to return to her daily life as soon as possible. Characteristics of Implant The chemical composition of the glass-ceramics is MgO 4.6, CaO 44.9, Si02 34.2, P205 16.3, CaF2 0.5 in terms of weight ratio. We made the glass-ceramics by compacting glass powder of the above composition under a pressure of 400 kgw/cm2 and heating it to 1050 C for sintering and crystallization. The resulting glass-ceramics had high bending and compressive strengths of 196 and 1076 MPa, respectively The Young’s modulus was 117600 MPa, and the porosity 0.7% [6]. We made a block of the glass-ceramics measuring 30 x 27.5 x 20 mm (Fig 1). The block had two holes, one 6.5 mm diameter and one 2.8 mm diameter. Operation The operation was performed using Hardinge’s lateral approach. 3 x 3 cm lateral iliac surface was exposed subperiosteally just above the hip joint. After the trial, the lateral surface of the ilium was planed with a 30 mm diameter ilium plane. The block of artificial bone was placed on the
Figure 1. A block of the glass-ceramics measuring 30 x 27.5 x 20 mm. It has two holes, one 6.5 mm diameter and one 2.8 mm diameter.
Figure 2. Right: preoperative radiograph, left: one year after the operation.
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lateral surface of the ilium and fixed with two titanium screws. The intraoperative hemorrhage was 167 gm and the operating time was 1 hour 42 minutes. No complication occurred postoperatively. Active and passive range-of-motion exercises of the hip were started the day after the operation. Crutch walking was permitted with partial weight bearing on the affected side three days after the operation. She was allowed to walk with single cane two weeks after the operation. The cane was discarded4 weeks after the operation and she returned to her daily life. RESULT S AN D DISCUSSIO N She had no pain of right hip joint and had slight limp because of pain of left hip joint one year after the operation. Her operated right hip joint had full range of motion and manual muscle test of all muscles scored excellent She could walk without any support for two hours without pain of right hip joint. We performed the Trendelenburg test at three months after operation and it was negative. Electrophysiological assessment of gluteus medius and tensor fascia lata was undertaken by a neurologist. He obtained a normal electromyogram pattern. Radiographic Study The radiographs showed that the acetabular angle of Sharp had decreased by 11 degrees (preop. 42 degrees and postop. 31 degrees), the center-edge angle had increased by 31 degrees (preop. 3 degrees and postop. 44 degrees) and the coverage of the femoral head had increased by 40% (preop. 60% and postop. 100%) as compared with the preoperative measurements (Fig 2). The weight-bearing and minimum joint spaces had increased by 1.2 mm and 1.2 mm, respectively, between the preoperative and one year follow-up measurements. Hip Scores The Harris hip score was determined at one year postoperatively. The hip score was 51 points preoperatively and 93 points at one year postoperatively. The difference between these scores was 42 points. The pain score improved from 10 points preoperatively to 40 points one year postoperatively. The gait score improved from 20 points preoperatively to 30 points one year postoperatively. She has pain of left hip joint and wants to have the implantation of glass-ceramic block on the left hip joint. Our procedure is unique, to our knowledge, in that we used a bioactive glass-ceramic block to restore the acetabular dysplasia The result of our procedure s^peared to be satisfactory. It was easy to know the correct position of the implant. The fixation of the implant was fod-proof. Our patient had not to have cast immobilization. The patient walked without a cane 4 weeks after the operation. The screws kept the block on the surface of ilium initially. Upon a previous study with dogs, the bonding between bone and the ceramic implant is thought to be chemical and biological and considered to advance between 2 weeks and 1 month after the implantation [5]. This procedure is an operation that places the femoral head beneath a surface of joint capsule with the capacity for fibrocartilagenous regeneration. The implant becomes a shelf and the capsule is interposed between it and the femoral head. The capsule under the shelf undergoes metaplasia to fibrocartilage [7]. SUMMAR Y Augmentation of the deficient acetabulum is effectively achieved with the glass-ceramic
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block. This procedure is simple and easy. It takes short operation time and little loss of blood. Patients have not to have cast immobilization. They can walk without a cane 4 weeks after the operation. We recommend the operation for patients in the third and later decades who have symptomatic early osteoarthrosis of the hip with acetabular dysplasia
REFERENCES 1. Konig F. Osteoplastishe Behandlung der Kongenitalen Hueftgelenksluxation (mit Demonstration eines Pareparates). Verh, Dtsch, Ges, Chir, 1891, 20, 75-80. 2. Chiari K.Ergebnisse mit derBeckenosteotomie als Pfannendachplastik. Z. Orthop,1955, 87, 14-26. 3. Salter RB. Innominate osteotomy in the treatment of congenital dislocation and subluxation of the hip. 7. Bone. Joint,Surg. [Br] 1961, 43B, 518-39. 4. Yoshii S, Yamamuro T, NakamuraT, Kitsugi T, Oka M, Kokubo T, Takagi H. Strength d bonding between A-W glass-ceramic and the surface of bone cortex. J. Biomed.Mater.Res. 1988,22,327-338. 5. Yoshii S, YamamuroT, NakamuraT, Oka M,Takagi H, Kotani S. J. Appl. Biomat. 1992, 3, 245-9 6. Kokubo T, Shigematsu Y, Nagashima Y, Tashiro M, NakamuraT, YamamuroT, Higashi S. Bull. Inst. Chem. Res. Kyoto Univ. 1982, 60, 260-8. 7. Moll EK Jr. J. Pediatr.Orthop.1982, 2, 573-6.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PARTICIPATIO N OF CALCIU M PHOSPHAT E CEMENT S FOR HEALIN G OF ALVEOLA R BONE M. Yoshikawa\ H. Oonishi^ Y. Mandai^ K. Minamigawa^ and T. Toda^ Department of Endodontics, Osaka Dental University, 5-31, Otemae 1-chome, Chuo-ku, Osaka 540, JapanV Department of Orthopedic Surgery, Artificial Joint Section and Biomaterial Research Laboratory, Osaka-Minami National Hospital, 2-1, Kido-higashi, Kawachi-nagano 586, Japan^. Nitta Gelatin Inc., Research Laboratory, 2-22, Futamata, Yao 581, Japan^. ABSTRAC T Experimentally developed calcium phosphate cements, a tetracalcium phosphate and dicalcium phosphate dihydrate (TeDCPD) cement and a a-tricalcium phosphate (a-TCP) cement, were applied in rat alveolar bone through mandibular first molar canals and osseous healing was evaluated. The liquid phase of the cements consisted of citric acid, tannic acid and distilled water. The pomt of a reamer was extruded from the apical foramen to mjure alveolar bone. The cements and commercially available zinc oxide eugenol sealer (ZOE) were respectively introduced into alveolar bone via the root canal. In the control group, no material was applied. TeDCPD cement caused no or slight inflammation and was displaced by new bone. a-TCP cement caused a moderate inflammatory reaction, and was displaced by hard tissue. ZOE was not displaced by hard tissue and mild inflammatory responses continued. In the control group, inflammatory reactions contmued throughout the experimental periods. It was concluded that osseous healing accompanied absorption of the cement. KEYWORD S Tetracalcium phosphate cement, a-tricalcium cement. Alveolar bone, Root canal sealer INTRODUCTIO N Sealers used in endodontic clinics should be applied to a root canal so as not to extrude from apical foramen into alveolar bone. Sealers, however often extrude into alveolar bone because of condensation of root canal filling materials. Commercially available sealers can irritate tissue causing inflanmiatory reactions in periapical tissue and alveolar bone [1]. In contrast, calcium phosphate cements have high biocompatibility [2] and mduce formation of hard tissue [3]. Such cements may prove suitable as sealers. In this study, the responses of alveolar bone to calcium phosphate cements were examined. MATERIAL S AND METHOD S Cement preparation Tetracalcium phosphate (TeCP) and a-tricalcium phosphate (a-TCP) were obtained by the dry synthetic method of sintering the stoichiometric mixture of calcium phosphate dihydrate and calcium carbonate (mol ratio of Ca/P : TeCP=2.0, a-TCP=1.5) at 1480 C for 8 hours, in the laboratory of Nitta Gelatin Inc.. The particles were 0.6 to 44.0 ^m, average of 12 to 13 ^m, in size. Dicalcium phosphate dihydrate (DCPD) was purchased from Wako Pure Chemical Inc.(Osaka, Japan). The calcium phosphates were passed through a 32 ^un-sieve. The powder 361
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phase of TeDCPD (an equimolar mixture of TeCP and DCPD) cement consisted of 70% TeDCPD and 30% barium sulfate, while that of a-TCP cement consisted of 70% a-TCP and 30% barium sulfate. The liquid phase of both cements was composed of 37 % citric acid, 5 % tannic acid and 58 % distilled water. The powder and liquid ratio of each cement was 1.2 g/ml. Stability and disintegratio n of the cement s The cements were kneaded for one minute and stored for hardening in an incubator for 1 hour at 37–1 C and 100% relative humidity. The hardened cements were immersed mto distilled water in a glass vessel and stored in the incubator for 23 hours. The cements were then incubated at 100 C to dryness. Once weight changes had stabilized to 0.5 mg or less, percentages of disintegration were calculated from the difference in the weight of the cements before immersion in distilled water and after drying. Three measurements were carried out. Differences were evaluated statistically by Students T-test (p<0.01). Cement pH during the setting process The pH of the cements kneaded for 1 minute was measured three times for 60 minutes at 5 or 10 mmute intervals. Significant differences were also evaluated statistically by Student’s Ttest(p<0.01). Histopathologica l examinatio n Access cavities of mandibular first molars of 48 rats were opened using a #1/2 round bar under general anesthesia. The root canals were prepared by usmg No. 15 to 25 reamers. The point of the reamer was extruded approximately 2 mm from the apical foramen into periapical tissue in order to injure the alveolar bone. All instruments used in this study were sterilized by ethylene oxide gas. The animals were divided into four groups. TeDCPD cement (Group 1), a-TCP cement (Group 2) or commercially available zmc oxide eugenol sealer (ZOE, Group 3) were introduced into the alveolar bone of anunals using a lentulo spiral. No material was applied in Group 4. The access cavities were sealed using a light curing resm. Mandibles were removed for histopathological examination at 1, 3 and 5 weeks after preparation. Serial 6 |xm thick paraffin sections were made and stamed with hematoxylm and eosin (H.E. staining). RESULT S AND DISCUSSIO N Stability and disintegratio n of the cement s TeDCPD cement differed significantly from a-TCP cement. The disintegration of TeDCPD and a-TCP cement was 8.6–1.2 and 4.5–0.3, respectively. TeCP may have been converted to hydroxyapatite (HAp), while DCPD is not related to the reaction and likely remamed on the cement surface. Accordingly, the stability of TeDCPD cement was considered inferior to that of a-TCP cement. Cement pH during the setting process There were no significant differences between the cements. The pH was approximately 4 during the setting process. Though TeDCPD has buffer capacity as a solid phase [4], this was exceeded by the acidity of the liquid phase. Citric acid was added to set the cement and to give viscosity. However, strong acidic material may damage tissue. Histopathologica l examinatio n In all Groups, at 1 week after preparation, severe defects of alveolar bone were observed.
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Figure 1. Five weeks after placement of TeDCPD cement in the periapical area Alveolar bone is processing in the reconstruction. A small amount of the cement is present. F: Apical foramen n: Newly formed hard tissue a: Apex I: Incisor canal (H.E. stain, Orig. mag. xl7.6)
Figure 2. Three weeks after placement of a-TCP cement in the periapical area Newly formed hard tissue (n) is present in the damaged area of alveolar bone. A: Apex I: Incisor canal (H.E. stain, Orig. mag. x20.0)
Figure 3. Five weeks after application of commercially available sealer ZOE are present in the large area of alveolar bone. T: ZOE sealer I: Incisive canal a: Apex (H.E. stam, Orig. mag. xlO.O)
Figure 4. Five weeks after injury to alveolar bone induced through the root canal Large area of periapical scar is seen. Fibrous connective tissue infiltrates in the bone marrow. F: Apical foramen I: Incisor canal S: Periapical scar C: Fibrous tissue proliferated in the bone marrow (H.E. stain, Orig. mag. xl2.2)
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However, no or slight inflammatory reaction was found from 1 to 3 weeks in Group 1. This suggested that there was no mfection through the root canal into the periapical area. TeDCPD cement did not stimulate periapical tissue. The cement was surrounded by macrophages (M(|)) and multi-nucleate giant cells (MGC) at 3 weeks after preparation, and had mostly disappeared or been displaced by new bone at 5 weeks (Figure 1). Reconstruction of alveolar bone was dependent on the disintegration of cement. MGC and Mcj) must be mduced if cement is to disintegrate in periapical areas. Expression of MGC is induced by intracellular digestion of resistible particles [5]. The cement in periapical tissue would be disintegrated by tissue fluid, and small particles of HAp and DCPD discharged on the cement surface during settmg. The particles might be absorbed by MGC and M(|). In group 2, moderate inflammatory reactions were observed at 1 week after preparation. At 3 weeks, hard tissue had deposited at the apex and reconstruction of the alveolar bone was observed (Figure 2). In Group 3, ZOE did not disappear and mild inflammatory responses continued (Figure 3). In Group 4, acute inflammatory reactions occurred durmg the early experimental period. No osseous healing or fibrous connective tissue was observed in the osseous defect at 5 weeks (Figure 4). It was suggested from the results that MGC and Mcj) would play hnportant effects to heal the osseous defect. And then, presence of a digestible material by phagocytes may be an essential factor to reconstruct alveolar bone defect. CONCLUSIONS In this experiment, alveolar bone was mjured mechanically inducing extensive bone defect. It was considered that, commonly, such defects would be repaired by fibrous connective tissue. This repair seemed to progress leisurely. However, the repair was promoted by using the calcium phosphate cements. The existence of non-absorbent in the alveolar bone prevented repair. It was concluded that osseous healing accompanied absorption of the cement in the alveolar bone. REFERENCE S 1. Pitt Ford, T.R. and Rowo A.H.R.
J. Endodon. 1989,15, 286-289.
2.
Yoshikawa, M., Toda, T., Oonishi, H., Kushitani, S., Yasukawa, E., Hayami, S., Mandai, Y. and Sugihara, F. In: Bioceramics Volume 7, Butterworth-Heinemann, Oxford 1994, 187-192.
3.
Yoshikawa, M., Toda, T., Oonishi, Mandai, Y. and Sugihara, F. Volume 9, Pergamon, Oxford 1996,453-456.
4.
Brown, W.E. and Chow, L.C.
5.
Glowacki, J., Jast, M. and Goldring, S. J. Bone Mineral Res. 1986,1, 327-331.
In: Bioceramics
In: UnitedStatesPatent1985, 4518430.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
SYNTHESE S OF RAPIDL Y RESORBABL E CALCro M PHOSPHAT E CERAMIC S WIT H HIG H MACR O OR HIG H MICR O POROSIT Y G. Berger, R. Gildenhaar, U. Ploska, M. Willfahrt Projectgroup "Biomaterials and Implants" of the Federal Institute for Materials Research and Testing, Unter den Eichen 87, D-12200 Berlin, Germany ABSTRAC T This study was aimed at preparing rapidly resorbable materials based on the main crystalline phase of calcium potassium sodium phosphate, Ca2KNa(P04)2, by using different techniques for obtaining different macro and micro porosity. One solution was achieved by using a technique to get spongiosa-like materials as described in the literature. A slurry of the material is inserted into a polyurethane sponge which evaporates during the sintering process leaving the sponge-shaped ceramics. The other solution is found in a new way. The calcium alkali phosphate is melted together with a sodium borosilicate glass. In the melt and/or during the cooling of the melt the material undergoes a phase separation of a spinodale nature. This leads to the possibility of leaching out the borosilicate glass almost completely with NaOH solutions leaving the crystallized calcium alkali phosphate skeleton. Both techniques offer some possibilities of an advantageous application as bone substitution material. KEYWORD S Calcium phosphate ceramics, resorbable ceramics, Ca2KNa(P04)2, porosity, processing. INTRODUCTIO N The reason for the development of a rapidly resorbable material based on Ca2KNa(P04)2 [1, 2] was that these glass-ceramics should possess a higher solubility than tricalcium phosphate ceramics granules which are used as a bone defect filler. Apart from the advantage of the actual increase of the solubility this material will retain some of the disadvantages of TCP ceramic granules when produced in the same manner to obtain compact particles. These particles will be impacted in the filled bone defect during the healing process, and further resorption will be prevented. Therefore, some techniques were sought to overcome these obstacles. In the literature some manufacturing methods have been described of obtaining hydroxyapatite (HA) as dense or macroporous ceramics, for instance by Monma and Takahashi [3], loku, Somiya, and Yoshimura [4], Xingdong et al. [5], Arita, Costano, and Wilkinson [6]. General aspects for processing porous ceramics as applied here in this paper were given by Saggio-Woyansky, Scott, and Minnear [7] using a polymeric-sponge method that typically produces open-cell structures similar to spongiosa. MATERIAL S AND METHOD S Material preparation of rapid resorbable ceramics The synthesis of various glass-ceramics containing Ca2KNa(P04)2 as the main crystalline phase has been described in detail elsewhere [1,2]. In this paper only one composition (code GB 14) of such glass-ceramics was used for all the fiirther preparation techniques (wt-%: 30.67 CaO, 43.14 P2O5, 9.42 NajO, 14.32 KjO, MgO 2.45). This glass-ceramics that melted in a platinum crucible at 367
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1550 C crystallized spontaneously even when shock-cooled. Material produced this way possesses a solubility measured in TRIS-HCl buffer solution of pH 7.4 at 37 C for 120 h that is eight times higher than alpha-tricalcium phosphate ceramics [2]. Variant 1: Spongiosa-like rapid soluble ceramics The GB 14 glass-ceramics was milled in an achate mill (Pulverisette 5, Fritsch Lie, Germany) several times. The particle size was measured (Master Sizer S, Malvern Instruments Inc., United Kingdom) to fmd out an optimal particle size distribution for further processing. This powder was mixed with about 40% water and a 0,3% polyethylenglycole binder (Uniox, Nippon Oils & Fats Inc., Japan) dissolved in this amount of water for making a ceramic slurry . The slurry was immersed in a polyurethane sponge that contain about 7 - 9 pores per cm (linear). The sponge was compressed to remove air, taken into the slurry, and then allowed to expand. When the slurry has carefully infiltrated the sponge, which can also be supported by repeating the com› pression/expansion step the provision of the slurry had to be removed also by corresponding compression of the sponge. In the next step the ceramic slurry was dried to deposit the material on the sponge. This process was realized by oven drying in the temperature range of 100 to 150 C for a duration of about 5 hours. In a next step the evaporation of volatile organics (180 - 350 C) and the bum out of the polymeric sponge were performanced. This process was bushed at temperatures between 650 up to 700 C for about 10 minutes. The fmal step was the sintering of the ceramics. It took place at higher temperatures: about 1050 to 1250 T for a duration of about 5 to 30 min. After this processing ceramics have been maintained as shown in Figure 1. Variant 2: Rapid soluble ceramics with high internal surface The glass-ceramic GB 14 was crushed into powder of about 60 ^im. A borosilicate glass V2 was melted in a platinum crucible at 1550 T (wt-%: 55.0 SiOj, 6.0 Na20, and 39.0 B2O3) and also crushed into a powder of about 100 jim particle size. After this, GB 14 and the glass V2 were
Figure 1. Scanning electron micrograph of the spongiosa-like rapid resorbable ceramics. The picture shows a ceramic segment of 3.63 cm by 2.03 cm.
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Figure 2. Scanning electron micrograph of the rapidly soluble ceramics with high internal surface before (left) and after leaching out (ri^t) the borosilicate glass with 2.5 N NaOH leaving a calcium phosphate ceramics based on Ca2KNa(P04)2 as the main crystalline phase. mixed in the mass ratio of 3:2 and molten in a Pt-crucible at about 1450 T for only 5 minutes. The melt was poured on steel plates and cooled at 450 C. During these processes the material crystallized spontaneously and the main crystalline phase was the expected Ca2KNa(P04)2 (XRD). The material was crushed into pieces of about 5 mm that were leached in 2.5 N NaOH at temperatures between 37 to 95 C for different times in the range of 0.5 to 10 hours. Figure 2 shows a material prepared by this technique. Methods The specific surface area was determined by using BET-method with nitrogen or krypton as analyzing gas (ASAP 2000, Micromeritics Inc., USA). Scanning electron microscope (SEM) investigations were perfomanced to determine the optimal sintering conditions and to control the morphologies of the shaped ceramic material (S-4100, Hitachi hic, Japan). RESULT S AlVD DISCUSSIO N Variant 1: Spongiosa-like rapid soluble ceramics The one decisive step in the process is to obtain a good sintered material that did not show any indications of overfired processing conditions. This means the sinter necks between the basic particles have to become stronger but the formation of a voluminous melting phase has to be prevented which would lead to the break-down of the ceramic bridges destroying the sponge, or rather spongiosa-like structure, of the shaped material. Optimal sinter conditions are 1150 C for only 10 minutes. Figure 3 shows how the particles are sintered. They are still separate as observable. Using these conditions it is possible to get material that consists of 70 up to 80% porosity. But there is another important parameter regarding the initial powder. The best results are given when the D50 value of the powder was in the range of about 8.5 ^m
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Variant 2: Rapid soluble ceramics with high internal surface One of the important steps realizing this variant was to find out a borosihcate glass composition that is highly soluble in an alkaline solution and did not tend to mix with the calcium phosphate glass-ceramics especially with the main crystalline phase of Ca2KNa(P04)2. The last point is supported by using relative short melting times. But it is interesting that the melt is clear. It is only during cooling that the melt crystallizes spontaneously as for the case of GB 14 without mixing with the borosihcate glass. The variation of the alkaline solution treatment leads to different values of the measurable internal surface determined by BET with nitrogen as analysis gas which is used when the specific surface was expected to be higher than 0.5 mVg.
Figure 3. SEM of spongiosa-like rapidly soluble ceramics (1150 "C, 10 min).
Interesting results were yielded by the use of 2.5 N NaOH solution at 80 C. In dependence on the solution time and the particles used the borosihcate glass could be removed out of this composite. For instance, specific surface values of about 8 mVg for a solution duration of 1 day up to 29 mVg for 17 days could be obtained. Chemical analysis of the leached material shows that only 0.16 wt-% B2O3 are remaining which would be important for an application as bone substitution material. CONCLUSIO N The rapidly resorbable material can be formed in a spongiosa-like shape with 70 up to 80% macro porosity as-well as granules that possess a high internal surface depending on the chosen leaching out conditions. Both material forms could be used later on for bone substitution. ACKNOWLEDGEMENT S This project is partly (materials with high internal surface) supported by the German Research Foundation (DFG), Bonn. The authors would like to thank also Mrs. Birgid Strauss for making the SEM investigations. REFERENCE S 1. Berger, G., Gildenhaar, R. and Ploska, U. In: BioceramicsVolume 8, Pergamon, Oxford 1995,453-456. 2. Berger, G., Gildenhaar, R. and Ploska, U. Biomaterials,1995,16, 1241-1248. 3. Monma, H. and Takahashi, T. Gypsum & Lime,No. 226, 1990, 143-147. 4. loku, K., Somiya, S. and Yoshimura, M., J. Mater.Sci. Lett, 1989,8, 1203-1204. 5. Xingdong, X. et al. ClinicalMaterials,1989, 4, 319-327. 6. Arita, I.H., Costano, V.M. and Wilkinson, D.S. J. Mater Sci: Mater,in Medicine, 1995,6,8-13. 7. Saggio-Woyansky, J., Scott, C.E. and Minnear, W.P. Amer.Ceram.Soc. Bull., 1992,71,1674-1682.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
COMPOSIT E BIOCERAMIC S MAD E OF MACROPOROU S CALCIU M PHOSPHAT E CERAMIC S FILLE D WIT H A SELF-SETTIN G CEMENT . HISTOLOGICA L EVALUATION . Patrick Frayssinet* "", Alain Lerch*, L. Gineste**, N. Rouquet*. Bioland, 132 Rte d’Espagne, 31100 Toulouse, France See de parodontologie, Ecole Dentaire, Chemin des maraichers, Toulouse, France Laboratoire du Tissu Osseux et des Pathologies Osteoarticulaires, Universite Paul Sabatier, France
ABSTRAC T The compressive strength of macroporous calcium phosphate ceramics was increased by filling the pores with a self setting calcium phosphate cement. The resulting material was implanted in sheep condyles and subjected to histological analysis after 20, 60 and 120 days. A progressive ingrowth of bone occurred as the cement was degraded. At four months, all the cement had been replaced by bone. Somefi-agmentsof cement were still embedded in the newly formed bone. INTRODUCTIO N Macroporous calcium phosphate ceramics have a very weak compressive strength ranging,fi-om0.5 to 10 MPa dependmg on the porosity percentage, pore size, chemical composition and grain size,. Furthermore, these materials are fi-agile and brittle, making it impossible to shape the material during surgical operations. Macroporous ceramics act as a scaffold and guide for the bone regeneration tissue. The so called bioactivity of these ceramics is linked to the material surface area and porosity. We have developed a composite ceramic made of macroporous ceramic in which the pores are filled with a DCPD based self-setting cement, confering the material with a compressive strength similar to that of a dense ceramic. The obtained material can also be mill-shaped during surgical operations. The DCPD-based cement is a highly soluble material which disappears fi-om the material within a few weeks thus liberating the pores for bone growth within the material. We implanted cylinders of composite ceramic in bone defects drilled in sheep bones, which were then histologically analyzed after various periods of time rangingfi-om20 to 120 days MATERIAL S AN D METHOD S Material to be implanted: The calcium phosphate macroporous ceramics consisted of 75% HA and 25% of P-TCP. The porosity was 70% and the mean pore size 500 ^m. All the pores were interconnected. They were filled with a DCPD- based calcium phosphate self-setting cement which set in situ.Once set the cement was composed of P-TCP particles dispersed within a DCPD matrix. The ceramic was shaped into cylinders 8 mm m diameter and 15 mm in length. 371
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Implantation protocol: A 9 mm hole was drilled in the external condyle by a lateral approach in 12 sheep. A macroporous cylinder loaded with cement was implanted in one site and control material consisting solely of a calcium phosphate macroporous ceramic was implanted at the same site in the opposite leg. 20, 60 and 120 days after implantation, four sheep were sacrificed by nembutal injection, and the condyles were removed. Histologicalanalysis : The samples were fixed in formaldehyde solution, dehydrated in graded ethanol solutions then embedded in PMMA. Thick sections were made using a low speed diamond saw and ground before being stained with a fiiscin-toluidine blue solution. RESULT S At 20 days: none of the implants had been osteointegrated. A few woven bone trabeculae were forming in the soft tissue surrounding the implants. These trabeculae generally originated fi-om the bone cavity edges and progressed toward the implant. The ceramic implants showed invasion of the pores with fibrous tissue and some trabeculae were sometimes found in contact with the outer ceramic surfaces. The pores of the composite cylinders had been obturated by the cement. Bone trabeculae had often formed at the implant surface and some had become inserted in the cement which showed degradation marks in these zones (Fig. 1). Some grains, a few micrometers in size, had been released making hollows in the cemented pores in which an extracellular matrix had formed. These grains had been phagocytosed by mononucleated macrophages in the implant periphery. At 60 days,the control materials showed partial integration with bone ingrowth at the surface of the external pores. The central part of the ceramic still contained fibrous tissue in which some mononucleated cells containing ceramic particles were visible. The composite showed a progressive ingrowth of woven bone in the external pores in which the cement was slowly being replaced by bone tissue (Fig. 2, 3). Most of the cement crystallites located close to the solid surface were coated by a proteinaceous substance. These crystallites were being phagocytosed and degraded by macrophages. Many osteoblasts were differentiating at the surface of mineral aggregates. Based on the localization of the osteoblasts at the surface of a mineral support, a preferential differentiation process seemed to be occurring on these surfaces. Not all thefi-agmentsof the cement were being degraded by macrophages. Some were included in the bone matrix which had been formed within the pores. Some degradation of the ceramic scaffold had also occurred. At 120 days: Bone tissue had grown into the entire cylinders which showed major signs of degradation. All the pores of the composite ceramics were free of cement, which was found as mineral particles within macrophages. Some islets of macrophages were present in the bone marrow cavity of the bone tissue. Although some islets of macrophages had phagocytosed bone tissue, no exaggerated bone resorption could be evidenced in close proximity to the material. The control ceramics were totally integrated and showed signs of resorption. DISCUSSIO N These composite materials exhibit very attractive mechanical properties. The Young’s modulus is 500 MPa, and the compressive strength 20 MPa. This must be compared with the Young’s modulus of 80-120Gpa of dense calcium phosphate ceramics and the compressive strength of IMPa of the
Composite Bioceramics Made of Macroporous Calcium Phosphate Ceramics: P. Frayssinetet al.
Fig. 1 : 20 days after the implantation, the ceramic (C) still appears as a dense material. Newly formed trabeculae are ingrowing in a pore in which the cement is degrading (>). Toluidine blue staining. Bar: 300^m.
Fig. 2 : After 60 days, many pores were invaded by bone trabeculae which were fixed at the cement surface. At first, there was an ingrowth of a loose connective inside the cement (>> ), some cement particles were included in the newly formed bone (i^). The ceramic walls were not degraded yet (cer). Toluidine blue staining. Bar : 150 ^m.
Fig. 3 : At high magnification, there is a degradation front in the cement which contains many macrophages loaded with cement particles (>>). Behind this front, there is a bone formation zone which shows osteoblasts (^) differentiation from the fibroblast-like cells of the loose connective tissue. These cells form preferentially at the surface of cement remains (cem). Toluidine blue. Bar: lOO^m.
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macroporous control ceramics without cement . The Young’s modulus is not so different from that of bone tissue which ranges from 1.4-50 0 MP a for the distal tibia (2) and the material can be shaped by milling during surgical intervention . The slow replacemen t of the cement by bone tissue is though to transfer the mechanica l strength to the bone as it was shown that the strength of porous calcium phosphate implant was highly correlate d with the amount of pore space which had become occupied by bone (3). The biological propertie s of the ceramic are not altered by the presence of the cement . The histomorphometr y measurement s (data not shown) indicate that the amount of bone and the ossificatio n rate was even higher in the cemente d ceramics. Some characteristic s of the histologica l reaction are interestin g when they are compared to the osseointegratio n process at the ceramic contact. The formation of an osteoid tissue at the material surface is precede d by the depositio n of a protein matrix which could not always be related to synthesis by cells in close proximity. This matrix, which had infiltrated the material micropores was not mineralize d and seeme d to be a prerequisit e for osteoid formation by osteoblasts . The bone formation process was very active in the cement degradation zones. The presence of macrophages that had phagocytose d the calcium phosphate crystals did not trigger an activation of osteoclast s as shown by the absence of osteolysi s of the woven bone that had invaded the ceramic. Furthermore, it seeme d that bone formation was increase d by the presence of this cement but its is unclear why this should be so. It could be linked to degradation of the calcium phosphate cement . The degradation of calcium phosphate material has already been describe d as favoring bone formation in contact with such materials (1). The complexit y of macrophage activation and the high number of products that can be synthesize d by these cells could partially explain these results. CONCLUSION S The addition of a self-settin g calcium phosphate cement to a macroporous biphasic ceramic can help to obtain ceramics with enhance d mechanica l propertie s but no alteration of their biological properties . REFERENCE S 1. Ducheyne, P., Beight, J., Cuckler, J., Evans, B., Radin, S., Effect of calcium phosphate coating characteristic s on early post-operativ e bone tissue ingrowth, Biomaterials 1990, 11531-540 . 2. Goldstein, S.A., The mechanica l propertie s of trabecular bone, dependenc e on anatomic location and function. J. Biomech. 1987, 20 : 1055-106 1 3. Martin, RB., Chapman, M.W., Holmes, RE., Sartoris, DJ,, Shors, E.C., Gordon, JE., Heitter, D C , Sharkey, NA, Zissimos, AG , Effects of bone ingrowth on the strength and non- invasive assessemen t of corralline hydroxyapatit e material. Biomaterials 1989, 10 : 481-488 .
Bioceramics, Volume10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposiumon Ceramicsin Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PHYSICA L PROPERTIE S OF AN APATITI C CERAMI C CONTAININ G TRICALCIU M PHOSPHAT E PREPARE D BY THE WAY OF A CEMEN T Z. Hatim**, M. Freche*, J i . Lacout* * Laboratoire des Matfiriaux, Physico-Chimie des Phosphates, ENSCT-INP UPRESA CNRS 38, rue des 36 Ponts, 31400 Toulouse, FRANCE ** Laboratoire de Chimie-Physique, Universit6 Chouaib Doukkali, Facult6 des Sciences, B.P.20, El Jadida, MAROC ABSTRAC T Biphasic ceramics containing both hydroxyapatite (HAP) and tricalcium phosphate (TCP) can be prepared by way of ionic cements. Two kinds of synthesis were examined: (i) dispersion of TCP particles in a cement before sintering, (ii) preparation of a non-stoichiometric cement, then decomposition into HAP and TCP during sintering. Only the second route leads to fairly dense, homogeneous ceramic with an increase of mechanical resistance with the amount of TCP. The particle size does not really influence the mechanical properties. Whatever the synthesis route, the overall rate of dissolution is strongly influenced by the composition of the ceramic and increases with the TCP content. KEYWORD S Bifrfiasic c^amics, hydroxyapatite, dissolution, mechanical resistance. INTRODUCTIO N The biocompatibility of hydroxyj^atite ( HAP : Caio(P04)6(OH)2) and tricalcium phosphate a or p (TCP: Ca3(P04)2) has attracted great interest in the potential of dense or porous materials for the filling of osseous defects [1,2,3]. The difference of solubility and biological resorbability between these two calcium phosphate materials [4,5], has led to the preparation of two-phase ceramics, with various global solubilities, offering the possibility of pore formation due to the faster resorbability of TCP. Such biphasic materials are now commonly used and the properties of the final two-phase ceramic are dependent on the HAP/TCP ratio [6], the sintering conditions, the homogeneity of the dispersion of the two phases and the particule sizes of each phase. Dense or porous hydroxyapatite c^amics can be obtained by diffo-ent methods [7]. One, recently developped method is the sintering of a calcium phosphate cement previously prepared and shaped Two methods of introducing TCP into the HAP matrix were used in the present study. In the first, various amounts of TCP granulates were mixed with the cement paste before sintering. In the second, a cement paste with different Ca/P atomic ratios ( 1.59 to 1.67) was sintered; here, the formation of the two phases (HAP and TCP) occurred during heating. With the first route, the final ceramic was formed by a dispersion of particles in an apatitic binder, in the second route, the HAP and TCP phases WCTCfinelydispersed throughout each other. Two important properties were examined : the compressive strength and the dissolution rate for a given pH value 4,3. Dispersion was visualized by microscopy. 375
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MATERIAL S AND METHOD S Ceramic preparation The cement was prepared by addition of a liquid containing calcium and phosphate ions to a mixture of solids : tricalcium phosphate (Ca3(P04)2) and tetracalcium phosphate (Ca4P209). The quantities of the different compounds were accurately determined in order to obtain a precise Ca/P atomic ratio. First route Particles of TCP were finely mixed with a cement paste with a Ca/P = 1.67. Three particule sizes ( 180-250,250-355 and 355-500 micrometers) and four proportions of TCP (0,10, 30, 50 and 70 % weight) were used. Then the cement paste was moulded and shaped to form cylindrical samples (diameter : 5 mm, length : 6 mm). The cement was hardened at SOX for 7 days in deionized water. The cylindrical samples were then heated to 1250X in order to obtain thefinalceramic. Second route Cement pastes with four different Ca/P atomic ratios (1.67, 1.65, 1.63 and 1.59 ) were prepared and moulded; the cylindrical samples were hardened at SO’^C for 7 days and sintered at 1250^C. In this case, after hardening, the cement formed was a nonstoichiometric apatite. During heating decomposition of this non-stoichiometric apatite occurred and HAP and TCP were formed according to the chemical reaction : Caio.x(P04)6.2x (HP04)2x(OH)2 - > (l-x)Caio(P04)6(OH)2 + 3xCa3(P04)2 + 2xH20 The amount of TCP (10, 30, 50, % weight) in the final ceramic is determined by the initial Ca/P atomic ratio. Physical and chemical analysis Ultimate compressive stress was determined using a press, (Hounsfield, serie S) making the average on three specimens. X-ray diffraction (Inel, CPS 120 ) was used to check the final composition. In vitro solubilities were determined using a pHstat (Methrom) which maintained the pH at a constant value (4.3) by addition of a recorded volume of HNO3 (0.00IM). The slope of the curve : volume of acid versus time, gives an indication of the rate of solubilization of the ceramic and allows comparisons to be made. Electron and optical microscopy observations were also made. RESULTS AND DISCUSSION [VIechanical 1testing .
Load(N)
150 0 1
150 0
135 0
135 0
’^
120 0
105 0
105 0
900
900
750
750
600
800
450 150
0
450
^X
300
300 150
^ 00
0.1
02
03
0.4
0
(C)
00
0.1
02
0.3
0.4
0
01
0.2
0.3
04
Displacement (mm) Figure 1. Compression behaviour of (a) ceramic : stoichiometric HAP (b) ceramic : containing granulates (first route) (c) ceramic : from non-stoichoimetric cement (second route)
Physical Propertiesof an Apatitic Ceramic Containing TricalciumPhosphate:Z. Hatim et al.
377
Table 1. Compressive stress of co-amic prq)ared from stoichiometri c cement (a) according to PTC P granulate sizes (b) according to propcition of P-TCP granulates.
(a)
1 p-TCP granidate sizes (urn) 1 Ultimate compressive 1 stress (MPa)
(b) 1
p-TCPgranulates (% ) weight 1 Ultimate compressive 1 stress (MPa)
180-250
250-355
355-500
50
56
45
0
10
30
50
48
50
76
45
1
Figure 1 presents the comp^ssive strength versus the displacemen t at constant speed for three representativ e specimens . Difference s were noted betwee n the ceramics prepared according to the two routes. The specimen s prepared by sintering a mixture of tricalcium phosphate particles with stoichiometri c cement (first route) exhibite d a similar behaviour to that of a foam with a short linear elastic deformatio n at low stress, followed by a collapse plateau due to the failure the individual struts (figure 1(b)). This can explain the large dispersion of the results. In contrast, specimens prepaid by decompositio n of a non-stoichiometri c apatitic cement (second route) exhibited a curve caracteristi c of a dense ceramic with a distinct point of failure (Figure 1(c)). The diff^ence s betwee n the two cement s are surely due to the porosity of the first one. This was verified by electro n microscq)y. Concerning the ceramic prepared according to the first route, the main results are r their sizes or their prop(»tions in presente d in Table 1. The addition of TCP particles whateve the ceramic did not ^preciably modify the compressive strength of the material. Large holes, observed by SEM , contribute to the heterogeneit y of the ceramic. In some cases particles of TC P were visible : they were closely associate d with the surrounding apatitic phase. It is possible, considering the binding, that the mechanica l behaviour is due more to the presence of pores than to the presence of TCP particles. With the second route, the ceramics were denser : microparticle s of TCP and HA P were homogeneousl y dispersed. SEM showed a ceramic with a homogeneousl y dispersed microporosity (pore diamete r of 1 micrometer) . The HAP/TC P ratio strongly influence d the mechanical properties . The values of the failure strain, largely increase for 30 % (194 MPa ) and 50 % (155 MPa ) of TCP . Comparison of the apparent density of the ceramics shows that a ceramic of pure stoichiometri c cement (Ca/P =1.67, and TCP = 0% ) presents a density of 2, a ceramic containing particles of TCP ( whateve r the amount and the size of the particles ) a density of 2.2, and a ceramic prepared from a non-stoichiometri c cement a density of 2.5 to 2.8. The higher the density is, the greater the resistance to failure is. The mechanica l resistance is particularly increase d with the second route when TCP is well dispersed .
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Dissolution
testing
Table 2. Dissolution of first route material: (a) according of P-TCP granulate % (b) according of p-TCP granulate size.
(a)
1 p-TCP granulate (% ) weight 1 Volume of acid added 1 (ml/min) x 10^
0
10
1 30
50
1 70
13
18
27
29
37
(b)
1 p-TCP granulates sizes (um) 1 Vdume of acid added 1 (ml/min) x 10^
180-250
250-355
29
22
100 1 41
1 355-500
1
29
The dissolution curves of the ceramic prepared according to thefirstroute are not very regular : this can be attributed to the disaggregation of some particles during the dissolution measurements. Nevertheless, a general slope was determined. All the results are reported in Table 2. The overall solubility of all the specimens increased with the proportion of TCP. For example, solubilization was doubled when the proportion of TOP in the material was increasedfrom0% to 30%. The range of particule sizes used in this study did not allow its influence to be clearly established. The dissolution phenomena of TCP and HAP are separate : the dissolution rate of TCP, at pH = 4.3, is three times greater than of HAP. So, the dissolution of the ceramic is selective. This favours the disaggregation of part of the ceramic, but also creates pores in the ceramic bulk. CONCLUSION Addition of TCP to apatitic ceramics preparedfromionic cements leads to compounds which present an overall solubility greata* than of pure apatitic ceramic. When TCP is added as particles no modification of the mechanical resistance s^pears; in contrast TCPfinelydispersed in ceramics prepared by sintering of non-stoichiometric cements, enhances the ultimate compressive stress; this can be attributed to a greats homogeneity and densification. Nevertheless, concerning biomaterial, application advantageous properties are often confored more by bioresorbability and rehabitation aq)acity than by high mechanical resistance properties. So, biphasic c^amics containing TCP particles could be of greater interest.
REFERENCES 1. Brown. W.E., and Chow. 1. C. Ann. Res. Mater, Sci; 6,213-226, (1976) 2. Ducheyne. Bioceramics 4, (1991), 135, Bonfield ED. 3. LacouL J.L., and Mejdoubi. E., Brevet Fr., 92.09019/PCT/Fr. 4. Daculsi. G., and Passuti. N., Bioceramics II, Onishi H et Heimcke ed Ishiyakuro ed Tokyo, (1990). 5. Klein. CP.A.T., Driessen. A.A., J.Biomed. Mater. Res., 17(5), pp. 769-786 (1983). 6. Toriama. M., Kawamura. S; Shiba. S, Yogyo Kyokai Shi 95 (4): 92-94 (1987). 7. Mejdoubi. E., Lacout. J.L., and Hamad. H., Bioceramics, 8:457-460, (1995).
Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CYTOCOMPATIBILIT Ca/P RATIOS .
Y OF CALCIU M PHOSPHAT E COATING S WIT H VARIOU S
Patrick Frayssinet* , Laurence Arbore*, Nicole Rouquet* * Bioland, 132 Rte d’Espagne, 31100 Toulouse, France. Laboratoire du Tissu Osseux et des Pathologies Osteo-articulaires.
ABSTRAC T Established cell lines and primary osteoblast cell lines were grown at the surface of HA-coating with Ca/P ratios ranging from 1,659 to 1.765. Cell growth and differentiation at the coating contact did not vary with Ca/P. The Ca concentration and pH of the medium during the culture at the different coatings contact was similar. These results suggest that a low amount of CaO contaminating the HA-coating did not induce a cytotoxic effect. INTRODUCTIO N Calcium phosphate coatings for biomedical use can show very different characteristics depending on plasma spray parameters and the calcium phosphate powder. The temperature attained during the spraying can lead to decomposition of the molecule. Tricalcium phosphate, oxyapatite or calcium oxide may appear in the amorphous phase of the coating,. The presence of calcium oxide in the coating is a cause for concern as it reacts with water to form Ca(0H)2 which is a strong base when dissociated. The presence of CaO could lead to cytotoxicity in contact with the coating due to a pH decrease in close proximity to the ceramic material. CaO can be detected by the value of the Ca/P ratio which is increased over the theoretical HA ratios (1.67). We grew cell cultures with a primary and an established cell line at the surface of calcium phosphate coatings with various Ca/P ratios for twenty days to allow the dissociation of Ca(0H)2 and detect any cytotoxic effect. MATERIAL S AND METHOD S Materials tested: Titanium alloy discs 8 mm thick, 30 mm in diameter were grit blasted and plasma sprayed with different powder batches containing various amounts of tricalcium phosphate (range 0-2.8%) and calcium oxide (range 0.9-02%). The polystyrene of the culture wells and grit-blasted titanium discs were used as controls. Cell culture: An established fibroblast cell line (L-929) was grown at the surface of the coatings for 20 days. The samples were placed on the surface of polystyrene wells and sealed to the well with an agar solution. The cultures were carried out in triplicate. The cells were grown in a DMEM medium supplemented with 10% foetal calf serum and glutamine and a 5% CO2 atmosphere. The culture 379
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medium was changed once at 10 days, then three times a week and the cells were counted after 24, 48, 120, 240 375, and 480 hours. A primary cell line of chicken osteoblast s was grown for one week on the coating surface under the same conditions as describe d above. At the end of the culture period (7 days), the membrane alkaline phosphatase activity of the cell line was labeled histochemicall y using the azo-dye kit from Sigma Chemicals and the stained surface measured using an image analysis device coupled to a Reicher Polyvar reflecte d light microscope . Table 1 : characteristic s of the different HA-coatings Implant group Ca/P coating mass (mg) thickness (^m) 1 0.18 1.756 142.6 2 1.765 0.22 154.9 3 1.697 0.18 143 4 1.725 0.18 140.6
Ra(^m) 9.62 8.8 9.85 9.48
Rt(^m) 65.4 60.6 66.6 64.46
Rp(|xm) 24.5 23.2 24.3 24.8
pHandcalcium content measurements pH and calcium content of the culture medium were measured in the culture medium every day during the first 10 days using a calcium and pH probe RESULT S Cell culture: The growth rates at the surface of the different coatings were similar and not significantly different from the control (fig. 1). Culture medium pH: The pH of the culture medium in which the cell cultures were grown at the surface of the different coatings were not significantly different except for the control cells betwee n 48 and 240 hours. The pH increase d fromthe beginning of culture and decrease d from 120 hours onwards (fig.2). Fig. 1 : number of cells per unit area at the surface of the different materials vs time (hours)
-- -
-Ca/P 1,756 - C a /P 1.765
- - A - -Ca/P 1,697 - - X - .-Ca/P 1.725
-- -
#-
100
200
-titaniu m polystyren e
300
400
500time . (hours )
Cytocompatibilityof Calcium Phosphate Coatings With Various Ca/P Ratios: P. Frayssinetet al.
381
Calcium content intheculture medium : The calcium content s of the culture medium diflfere d significantly betwee n the calcium phosphate coated and control materials (fig. 3). The calcium content was constant in the medium in contact with control and increase d throughout the period of measuremen t for the medium in contact with calcium phosphate coated materials. No significant differenc e could be demonstrate d within the group of coated materials. Membrane alkaline phosphatase activity: No significant difference s were shown betwee n the surface areas occupied by cells exhibiting membrane phosphatase activity within the group of coated materials (table 1) table 1 : percentag e of the sample surface occupied by cells exhibiting a membrane alkaline phosphatase activity material % surfacelox
"""""^^l ip^lyiSi^^ 11.8 – 1.72
cli?\.691
CdJ? 1.725
Ca/P I J ^
Ca/P 1 765
4 – 1.89
1.21 – 1.19"
3.75 –1.47
4T25 – 1.56 "
Fig. 2 : culture medium pH measured at the material surface at different times of the culture
7 ,9
Ca/ P
’
1 ,75 6
Ca/ P
1 .76 5
C a /P
1 ,69 7
-Ca/P
1 ,72 5
tita n iu m
- polystyren e cultur e m edium
2
tlm.ft,’’our.)’» «
^^0
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Fig. 3 : calcium content of the culture medium at different times of the culture
120-
.*
Ji A
100-
? 3
-- --Ca/ P 1,756
-
80
- C a /P 1.765
- - A - - C a /P 1,697 -- --Ca/ P 1,725 B titaniu m
1 60 o
’ J^* ’ 1
- ^ ’
.
,..i...,..A...-. - ^rJl i linr V Vi/Giiuies
^
20 - r^^ifllSi r*
b ^^a
fl -
0
1%
V
50
"^
n
s ^
-V
"
100
’
150
n ’
200
0
Milieu
^
V
250
time (hours )
DISCUSSIO N AND CONCLUSION S Cell growth rate and differentiation remained unaltered by the increases in the Ca/P ratios up to 1.765 meaning that less than 5% CaO was present. The pH and calcium contents of the culture medium in which the cultures were carried out were not modified by changes in the Ca/P ratio. These results suggest that Ca(0H)2 is poorly soluble in the tested medium or that the CaO phase in the coating is poorly available to the cell or the extracellular fluids. It has been proved that the CaO phase forms preferentially close to the metal surface (1). It is probable that CaCOa is formed due to the presence of a high concentration of CO2 in the atmosphere. This high content of CO2 is probably responsible for the pH shift observed as soon as the culture medium is placed in the incubator. The decrease in pH after 120 hours was due to cell line metabolism. These results emphasize the difficulty in establishing specifications for calcium phosphate coatings. The spraying process may lead to the deposition of layers of complex materials, the characteristics of which may be difficult to determine. REFERENCES 1/ X, Ranz, L, Gobbi, F. Rustichelli, N. Antolotti, C. Rey, Structural problems linked to hydroxyapatite coatings.in : Ravaglioli (ed) Fourth Euro Ceramics, 8 : 19-26, Gruppo Editoriale Faenza Editrice S.p .A. - Faenza Italy.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
COMPARISO N OF RESORPTIO N AN D BON E CONDUCTIO N TW O CaCOsBON E SUBSTITUTE S
OF
J.C. FRICAIN*, Ch BAQUEY, B. BASSE-CATHALINAT, B. DUPUY INSERM U 443 146 Rue Leo Saignat 33076 Bordeaux-Cedex, France * UFR Odontologie de Bordeaux 2 16-20 Cours de la Mame 33000 Bordeaux (France) ABSTRAC T The aim of this study was to compare the behavior of natural coral and calcite (obtained by heating of coral) after implantation in the femoral condyles of New Zealand male rabbits. The methods used were decalcified and under calcified histology scanning electron microscopy, X -rays microprobe analysis, gravimetry and histomorphometry. Results show that coral and calcite were progressively resorbed and replaced by bone ; the phenomenon was slighdy slower with calcite.
KEYWORD ^ Coral - Calcite - Resorption - Bone substitute INTRODUCTIO N The two crystalline forms of CaCOs aragonite ( from natural coral ) and calcite (from natural limestone) have been used with success as bone graft substitutes (1-23). However, natural coral transformed into calcite by heating has never been tested in vivo in osseous sites. In vitro,calcite has shown the same cytocompatibility as native coral (4). The aim of this study was to compare the behavior of coral and calcite after implantation in bone. MATERIAL S AN D METHOD S Cylinders (0.5 x 1 cm) of natural coral or calcite obtained by heating at 500 C for 15 hours (5) were weighed then implanted in the femoral condyles of 18 New Zealand male rabbits; two holes were kept empty for the control. At D 15, 30 and 60 the condyles were retrieved and decalcified or not. The decalcified implants were used for decalcified histology (hematein, eosin, safran staining) and to measure calcite and coral resorption by gravimetry as was previously described (6). The undecalcified implants were fixed in 10 % buffered formalin, then embedded in methyl methacrylate, butyl methacrylate according to the Wolf technique (7). The semi-thin sections (30 |Lim thick) were made with a diamond saw according to the 383
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usual method (8) and then studied by undecalcified histology (solochrome staining), histomorphometry (evaluation of the pore surfaces using a computer assisted system for histomorphometry camera CCD Sony 95C, Biocom 2000), scanning electron microscopy (Hitachi S 2500) and X ray microprobe analysis (Ca, P and C quantification after coating sections with carbon). RESULT S * Undecalcified and decalcified histology have showed that control holes were filled by fibrous tissue at D 15, then osteoid and osseous bone at D 30 and 60. At D 60, holes were nearly totally filled by trabecular bone. For the coral and the calcite, results were similar. At D15, filling was at the periphery of the implant and the pores were colonized by fibrous tissue. Between old bone and the implants new trabecular primary bone had appeared. At D30, coral and calcite were partially resorbed and replaced by bone (Fig. 1). In the pores, new trabecular bone was visible in contact or not with coral or calcite (Fig. 2). Many osteoblasts were lining the osteoid and surfaces of the coral implant pores. At D60, bone formation increased from the periphery to the center of the implant which had nearly entirely disappeared. * HISTOMORPHOMETR Y Macroporous surface was (X – DS) 29.49 % –1,6 before implantation, 29.29 % – 1.2 at D15, 58.57 % – 3.7 at D30 and 89.43 % – 1.8 at D60 for the coral; 31.02 % – 1.5 before implantation, 28.59 % – 0.9 at D15, 48.2 % – 1.76 at D30 and 79.48 % – 1.8 at D60 for the calcite.
Figure 1. Photograph of coral after 30 days of implantation. Note at the periphery the implant invasion by newly formed bone and at the center the coral without any bone ( / ) (G xlO - solochrome staining undecalcified section 30 |xm thick)
Figure 2. Photograph of calcite after 30 days of implantation. In the pore, trabecular bone surrounded by osteoblasts (OB) is shown (G x 200 hematoxylin -eosin safran staining).
Comparison of Resorptionand Bone Conductionof Two CaCOs Bone Substitutes:J.C. Fricain et al.
385
* GRAVIMETR Y The percentages of weight loss during implantation were at D15, 30 and 60 :17.38 % - 39.06 – 3.7 - 83.8 – 7.3 for the coral and 11.75 % - 18.1 – 9.3 % 64 – 1.3 % for the calcite, respectively * Scanning electron microscopy and Xray microanaliysis.In the pore new bone was more often in contact with coral or calcite. Before implantation for coral and calcite, the percentage of Ca was equal to 100 %, 0 % for P and the Ca/P ratio was 1.62 for the bone control. At 15, 30 and 60 D the ratio varied from 10.15 – 43 to 3.32 – 0.2 then –2.12 – 0.23 for the coral and from 10.79 – 2.6 to 4.2 – 0.1 then 2.43 –0.16 for the calcite. The analysis of P, Ca and C at the interface of the bone and implants showed that there was decreased in phophorus in the implants by comparison with bone and that the level of Ca was constant (Fig. 3 et 4). DISCUSSIO N Natural coral is resorbed in contact with bone or soft tissue (2 - 9 - 10). A previous study showned that calcite was resorbed in contact with soft tissue (3). Our results (morphometry - gravimetry) demonstrate that calcite was resorbed in contact with bone. After D15, 30 and 60, on the samples studied, calcite was resorbed but slightly move slowly than natural coral. At D15, Morphometry did not show implant resorption. This was because with this technique, only changes in macroporosity are measured but not the modification of microporosity
Figure 3. SEM photograph of undecalcified calcite sections showing the direct bone (O) bound to calcite (cal). Simultaneous line analysis of calcium (Ca), phosphorus (P) and carbon (C) starting from left (Cal) to right (connective tissue (TM). Note the continous hight level of calcium content at the interface between calcite and bone.
Figure 4. SEM photograph of undecalcified coral section showing the direct bone (O) bounding to coral (Cor). Simultaneous line analysis of Calcium (Ca) Phosphorus (P) and Carbon (C). Starting from left connective tissue (TM) to right area (TM). Note the continous high level of calcium content at the interface between coral and bone.
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.The mechanism by which coral and calcite were resorbed is probably linked to the osteoclasts (2-11) but it is possible that phagocytic cells of the connective tissue participate in CaCOs implant resorption (11). Coral and calcite were progressively filled by new formed bone (histology and microprobe analysis) which grew in contact or at a distance from the implants. Walker et Katz (12) have proposed two mechanisms to explain the adherence of bone to calcite: formation of calcium carboxylate or calcium-sulfate compounds and bone adhesion to an overgrowth of hydroxyapatite on the calcite surface. More recendy, Muller Mai et al (13) proposed two processes to explain mineralization at the coral surface: degradation and precipitation in the hydroxyapatite form and mineralization due to the production of calcified globules by cells at the coral surface. However in our study, the presence of bone at a distancefromthe implant and the spontaneous regeneration of bone in the holes of controls indicate that bone cicatrization is not only due to the materials used. CONCLUSION The natural or transformed calcite coral implanted in the femoral condyle of rabbits was progressively resorbed and replaced by bone.The phenomenon was slighdy slower with calcite. REFERENCE 1. Guillemin G., Patat J.L., Foumie J., Chetail M. /. Biomed,Mater.Res., 1987, 21, 557-567 2. Guillemin G., Meunier A., Dallant F., Christel P., Pouliquen J.L., /. Biomed. Mater.Res.,1989, 23, 765-779 3. Fujita Y., Yamamuro T., Kotanis S., Otsuki C, Kobuto T., /. Biomed.Mater. Res., 1991, 25, 991-1003 4. Fricain J.C, Bareille R., Amedee J., Dupuy B., lADR, Orlando, 1997 5.Fricain J.C, Choussat P., Rouais F., Baquey Ch., Dupuy B.,FifthWorld BiomaterialsCongress, Mars 29-June 2, 1996, Toronto, Canada 6. Fricain J.C, Basse-Cathalinat B., Dupuy B., Actualitesen BiomatMawc VolumeVI Edition Romillat,1997 7. Wolf E., Roser K., Hahn M., WezkerUng H., Delling G.G., Virchow Archiv. Pathol.Anat.Histopathol., 1992, 420, 17-24 8. VanDerLubbeH.B.M.,Patklein C, De Groot K. Stain technology,1988, 63 (3), 171-176 9. Albustany K.R., Hunt J., William D.K., 11th EuropeanConferenceon Biomaterials,Pise (Italic), September 10-14, 1995. Livre du Congres, pp 469472 10. Fricain J.C, Roudier M., Rouais F., Basse-Cathalinat B., Dupuy B. J.Periodont. Res.,1996,33 , 463-469 . ll.Voigt C, Merle C, Muller-Mai C, Gross U. /. Mat. Sci.: Materials in medicine,5, 688-691 12.Walker M.M., Katz J.C, Bull.Hosp. JrDis, Orthop.Inst.,1983, XLIII, 103. 13.Muller-Mai C, Voigt C, de Ahneida Reis S.R. /. Mater. Sci. Mater. Med., 1996, 7, 479-488.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
RELIABILIT Y OF DUAL ENERG Y X-RA Y ABSORPTIOMETR PHOSPHO-CALCI C BIOCERAMIC S IN RABBI T
Y IN EVALUATIO N OF
J.X. Lu\ O. Legrand^ B. Flautre\ A. Galliu^, M. Descanqjs^ B. Thierry^, P. Hardouin^ andB. Sutter^ 1. IRMS and 2. Service de Medecine Nucleaire, Institut Calot, 62608 Berck-Sur-Mer,¥KP^CE. 3. CRITT Ceramiques Fines, Z.I. du Champ de 1’Abbesse, 59600 Maubeuge,FRANCE.
ABSTRAC T Dual Energy X-ray Absorptiometry (DXA) allows direct measurement of bone mineral content. We evaluated its interest for evaluation of porous ceramics before and after bone implantation. HA and p-TCP implants were inserted in 3 mm diameter cavity of rabbit femur with 3 delays (TO, T12 and T24 weeks) for DXA and histomorphometric analysis. Our results show excellent new bone contact in interface implant/bone for the two ceramics. Bone ingrowth and biodegradation were more important in p-TCP than in HA. Comparison of the sites, bone ingrowth activity was cortical > cancellous > marrow site. But biodegradation activity was marrow > cancellous > cortical site. After comparison of the two methods, we consider that DXA can bring complementary inquiries in the evaluation of porous ceramics. Its non invasive and atraumatic character should permit in vivo longitudinal survey, and analysis of biodegradation in resorbable ceramics and bone rehabitation in non resorbable ceramics. KEYWORDS : Bioceramics, Bone, Dual energy absorptiometry (DXA), Histomorphometry. INTRODUCTIO N During last years, biomaterials based on calcium phosphate ceramics have been extensively used as osseous substitute in experimental trials and human clinic. This development for clinical practice have necessitated non invasive evaluation. There is numerous methods for postoperative biomaterials measurement. Histomorphometry is helpfiil but traumatic. Conventional imaging modalities are imprecise for quantitation. The aim of this study was to test if Dual energyX-ray absorptiometry (DXA) could bring quantitative evaluation in bone biomaterials in an e?q3erimental model in rabbit, comparatively to histomorphometry. DXA uses a X-ray tube with two different energy levels, enabling correction for the absorption by soft tissue. DXA allows precise, accurate, and non-invasive direct measurement of bone mass in himian and animals ^^’^\early bone changes after total hip arthroplasty ^^\ With adapted software, and high resolution collimators, DXA make possible evaluation of small samples, histologic sections or bones ^’’’^l MATERIA L AND METHOD S Bioceramics:Hydroxyapatite (HA) and Beta Tricalcium Phosphate (P-TCP) with same porosity (50%, measured by Hg porosimeter), pores size (100-300^m) and porous interconnections (30387
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lOO^m, measured by morphometry), were used. Ca/P ratio was 1.55+0.03 for p-TCP and 1.6 4 –0.03 for HA (measured by Inductive Coupling Plasma). The cylindrical implants were 3 mm diameter (dia.) and 6 – 0.2 mm length. Animals:20 female New Zealand rabbits, 10 month old (adult), 4.0 – 0.3 Kg body weight. Surgical procedure:We used a method deviated from Pasquier et al}^\ In rigorous asepsis conditions and under general anaesthesia, the implants were inserted in a cavity of 3 mm dia. and 6 mm depth in the middle diaphysis (cortico-medullar site) and the extern condyle (cancellous site) of both femurs. Delays: rabbits were euthanazied by overdose anaesthesia initially (TO, 4 rabbits), 12 weeks after implantation (T12, 8 rabbits) and 24 weeks after implantation (T24, 8 rabbits). The total femurs were removed without soft tissue and fixed in 10% neutral buffered formol during two weeks. 8 samples for each ceramic and for each site were removed at T12 & T 24 delays. DXA : In vitroprocedures’.DXA accuracy and reproductibility (precision) were evaluate by comparison with pure weight (PW). 10 HA and 10 p-TCP bioceramics samples were weighted after desiccation. DXA was performed samples soaked in a tissue equivalent material (75^ ethanol, 4 cm depth). We used the manufacturer supplied ultra-high resolution device: line spacing 0.254 mm and point resolution 0.127 mm (Hologic QDR-IOOOAV, Waltham MA USA). Bone Mineral Content (BMC) in mg of equivalent HA or p-TCP was calculated. Ex vivo procedures:The femurs were measured with same technique in vitro,BMC and Bone Mineral Density (BMD) = BMC per projection surface area, mg/mm^ were obtained for regions of interest (ROI) in implanted areas: cancellous (Cn-S) (3.05 x 5.97 mm^), cortical (Ct-S) and medullar (Ma-S) (3.05 x 1.65 mm^), and also in reference (ref) areas: contralateral condyle and proximal diaphysis: (Fig. 1).
D Reference area m litkplantation area
Medullar area D Cortical area
Figure 1.Interest areas for measure Histomorphometr y (HMM) : After DXA analysis, a undecalcified bone preparation was used for each specimen. Residual pores volume (RPV, %), new bone volume (NBV, %), residual material volume (RMV, %) were measured on two 50 ^mi sagittal implant sections with Van Gieson’s Picro-Fuchsine staining. Ratio of material degradation (RMD = 100% - measured RMV / initial RMV) was calculated. Statistical analysis: Results were expressed by means and standard deviations. Impaired bilateral /-test and paired Wilcoxon-test were used to compare the two ceramics, and the bone sites with reference bone by
Reliability of DXA Absorptiometryin Evaluation of Phospho-Calcic Bioceramics:J.X. Lu et al.
389
delay. Correlations between NBV, RMD and BMD were studied with Pearson correlation coeflScient (r). RESULT S AND DISCUSSIO N DXAin vitro(Table1)\ The coefiBcient of variation (CV) of the measure is excellent. Despite narrow range of PW values (from 43.4 to 50.0 mg for p-TCP, 57.0 to 61.5 mg for HA), BMC and PW correlated very well (p<0.001) in HA and p-TCP. Nevertheless, we found significantly difference between DXA and PW, demonstrating a systematic enor due to DXA, comparable for p-TCP or HA (BMC/WP ratio). Table 1.Implantsmeasurementsbeforeimplantation CV
Sample
PW(mg)
BMC(mg )
r
BMC/P W Ratio
HA
0.70 %
10
58.54 + 1.23
59.12 –1.94
0.984
1.010
p-TCP
1.08%
10
47.48 + 1.83
47.92 + 1.99
0.873
1.009
Implant
DXA ex vivo (Table2): To compare DXA results in the different sites, we used BMD which avoids variation in ROI size. Table 2. BMD (mg/mm^Jof the implantsafterimplantation Sample In cortical site In medullar site In cancellous site T24w T24w T24w T12w T12w T12w 8 0.7610.07 0.7210.05 0.6710.05 0.6510.06 0.4710.07 0.4410.06
ref P-TCP
8
0.8710.05 0.8010.13 0.8410.09 0.8010.12 0.73 10.07 0.5510.09
HA
8
0.8810.10 0.8810.06 0.9210.08 0.9610.09 0.7810.03
0.7910.07
In reference bone, BMD decreased with time in the three sites. This decrease could be due to surgical procedures and increase in animals age. Whatever the delay and the bone site, BMD is higher in implantation area than in ref bone area. This demonstrate a correct selection of measured area and persistence of biomaterial even at 24th week. BMD decrease with time in p-TCP particularly in Ma-S (^<0.001) showed ceramic biodegradation. Thus, osseous recolonization could not be highlighted. On contrary, an weak increases of BMD e?q)ressed new bone formation in HA. HMM HMM have shown a excellent biocompatibility in vivo for HA et p-TCP. Fig. 2. New Bone FoimAtioii in Time
Fig . 3. Material Degradation in T i me
6 0.
HA
50
BTC P
100 1
1 BHA 80
BTC P
40
20 H 20
10OH
0
Cn-12
Cn-24
a-12 a-24 Site and delay
Ma-12
Ma-24
Cn-12
Cn-24
a-12 a-24 Site and delay
Ma-12
Ma-24
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In HA , whateve r the site, biodegradatio n was not showed (no RM V and RPV modification) . In p-TCP, new bone formation and ceramic degradation was depending on bone site. In Ct-S, RP V and RM V decrease d with time (p<0.05), NBV increase d (^<0.05). In Ma-S RPV increased with time (p < 0.05). These results indicate the modificatio n ways of p-TCP: new bone formation is higher in Ct-S, material degradation prevails in Ma-S (Fig. 2 and 3). Correlations. At 24th week there was a significant correlatio n betwee n NBV and RM D in Cn-S (r = 0.810; /7<0.05 ) and betwee n BM D and RM D (r = 0.782; p<0.05) in Ma-S only for p-TCP. No correlation was observed in HA. CONCLUSIO N This preliminary study showed that DX A is a precise and accurate method for mineral content measuremen t of small samples of phosphate-calci c ceramics in vitro. When these samples are implanted in rabbit femur, the new bone formation activity is higher in cortical bone site, material biodegradatio n prevails in the medullar site. DX A allows analysis of implant inside bone, and materials modification , according to information brought by HMM . Our results confirm the potentia l interest of DX A in bone substitute evaluation t’^’^-^’^l DX A can bring complementar y inquiries in the evaluation of porous ceramics. Nevertheless , DX A cannot replace HMM : for a given implantation site in a whole bone, BM C adds measuremen t of material degradation, new bone formation, peri-implant modifications , and could not distinguish betwee n the mineralisatio n within the implants and their periphery. This major disadvantage is counterbalance d by its non invasive and atraumatic character, which should permit in vivo longitudinal survey, and complementar y dynamic analysis of biodegradatio n in resorbable ceramics and bone rehabitatio n in non resorbable ceramics. This technique will be expedien t for evaluation efficacy of calcium phosphate as bone substitute especiall y in vivo. REFERENCE S 1. Johnston CC , Slemenda CW , Melton LJ. NEngJMed 1991, 324:1105-110 9 2. Grier SJ, Turner AS, Alvis MR . Invest Radiol(1996) 31:50-6 2 3. Sutter B, Legrand O, Hardouin P, Bascoulergue G.J BoneMinRes 1994 9:S275-B18 4 4. Denissen H, Verhey H, De Blieck J, Corten F, Klein C, Vanlingen A. J PeriodRes 1996, 31:265-27 0 5. Shaoul J, Leichter I, Weiss AM , Chavel AC , Nyska A, Mosheiff R, Liebergall M , Klein BY . EurJ ExpMusculoskel Res 1995, 4:201-20 6 6. Pasquier G, Flautre B, Blary MC , Anselme K, Hardouin P. J MatScien1996, 7:683^69 0 7. Denissen H, De Blieck JD, Verhey H, Klein C, Lingen AV . J BoneMinRes 1996, 11:638-64 4 8. Corten HGA , Caulier H, Van Der Waerden JPCM , Kalk W, Corstens FHM , Jansen JA. Biomaterials 1997,18:495-50 1
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TH E EVALUATIO N OF DEGRADABILIT BIOGLASS fi IN-VITRO
Y OF MEL T AND SOL-GE L DERIVE D
D.C. Greenspan^ J.P. Zhong^ X.F. Chen^ and G.P. LaTorre^ 1. U.S. Biomaterials Corporation, One Progress Blvd. #23, Alachua, FL 32615, USA 2. University of Florida, Gainesville, FL 32611, USA ABSTRAC T Melt-derived Bioglassfi has been used with success clinically for over 10 years. It has been well recognized that its rapid bonding with living tissue and promotion of bone growth are associated with the induction of hydroxy-carbonate apatite (HCA)[1]. Newly developed sol-gel derived Bioglassfi with the unique feature of a micro-porous structure has exhibited not only the rapid induction of HCA but also a significant degradability when their pores reach a certain size[2]. In this work, molten and sol-gel derived Bioglassfi have been tested in a simulated body fluid (SBF) and a comparison of the HCA formation and degradability has been made in vitro. While sol-gel derived Bioglassfi showed a more rapid induction of the HCA layer than molten derived Bioglassfi it also clearly showed much more resorbability, which could be very beneficial in clinical bone-grafting application. KE Y WORDS : Bioglassfi, Sol-gel, Degradability, Hydroxy-carbonate apatite, in-vitro INTRODUCTIO N The 45S5 composition of melt derived Bioglassfi has been successftiUy used clinically as a monolithic implant to reconstruct the ossicular chain[3], and to maintain the alveolar ridge after tooth extraction[4]. The particulate form of the 45S5 composition has been used commercially to regenerate bone lost as a result of periodontal disease[5].The ability to control the in-vitroand invivo bioactivity of melt derived glasses by altering their composition is well documented[6,7]. Although it is possible to control the reactivity of the glass by altering its composition, it is difficult to control the degradability of the glass by this route. The ideal synthetic bone graft particulate would be one which completely resorbs within a desired period of time and is replaced by natural tissue. Previous work on sol-gel derived glasses indicate the ability to control the pore texture of these materials which may influence their in-vitroand in-vivodegradability. In this present study the in-vitrodegradability of 58S sol-gel derived glass particulate is compared with 45S5 melt derived particulate to determine the influence of porosity on the degradation process.. MATERIAL S AND METHOD S The 45S5 meh derived glass was prepared as previously described in the literature[8]. The 58S sol-gel derived glass was prepared by acid hydrolysis of tetraethoxysilane, SiCOCjHs), with additions of triethylphosphate, OP(OC2H5)3, and calcium nitrate as previously described[9]. 391
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The 58S was heat treated to 700C in ambient air with a resulting average pore radius of 43A. Both materials were then ground and sieved to the desired particle size range of 710-300|im . Some physical characteristic s of these materials are listed in Table I. All samples were reacted in simulated body fluid(SBF#9)[10 ] in an orbital shaker at 175 RPM and 37 C for time periods from 1 hour up to 28 days. This dynamic testing model was determine d to yield reproducible results for in- vitrotesting of bioactive particulate[ll] . The sample solutions were exchange d at time intervals of 6 hours, and 24 hours, and then every 2 days to prevent the ionic saturation of the solution. The samples were then recovere d using vacuumfiltrationwith 5|i m filterpaper and then rinsed in acetone to prevent further surface reaction. Particulate samples were then submitted for FTI R analysis and solutions were submitted for inductive coupled plasma(ICP) analysis. Table I Physical Characteristic s of Bioactive Glasses
45S5
Bulk Density (g/cm^) 2.65
Average Pore Radius (A) N/A
Surface Area (MVg ) 0.02
Pore Volume (cmVg) N/A
58S
0.99
43
207.00
0.41
Composition
RESULT S AND DISCUSSION : The results of the solutions analysis for cumulative soluble silicon can be seen in figure 1. The results indicate that initially both materials leach silicon rapidly over the first several hours. After 6 hours reaction time the sol-gel derived material continue s to leach silicon at a high rate exceedm g the melt derived material. After 2 days the leach rate for the melt derived material appears to level off at a concentratio n of 170ppm while the sol-gel derived material continue s to leach Si at a higher rate. After 6 days it appears that the leaching rate for the sol-gel derived material levels off to a concentratio n of 350 ppm. The results of calculating the weight loss for both materials from the solutions data is shown in figure2. You will note that there are negative weight losses for phosphorus and calcium which indicate an actual weight gain for both materials. This is due to the early precipitatio n of calcium phosphate on the surface of the particulates . The curves indicate that the sol-gel derived material experience s a greater weight loss of Si and a faster weight gain of calcium and phosphorus due to a greater degradability and a more rapid formation of calcium phosphate when compared to the melt derived material. The FTIR spectra in figure 3 confirms this with the sol-gel derived material forming crystalline HCA with in 6 hours while the melt derived particulate forms crystalline HCA betwee n 6 and 24 hours exposure to SBF. The greater in-vitro resorbability of the sol-gel derived material can be attribute d to the pore structure, high surface area and low bulk density as compared to the melt derived material(Table I). When a bioactive glass is expose d to solution a series of surface reaction occur including stage 1, ion exchange , and stage two, network dissolution[12] . These two stages contribute to a materials resorbability . A high surface area material enhance s these processe s by increasing the accessibilit y of reactive species on the surface of the glass to the surrounding solution. The pore structure can also influence the rate of formation of HCA on the surface of bioactive glasses. The effect is dramatic in this study as we are comparing porous and nonporous materials but can also be seen in porous materials of different pore textures[2] .
Degradability of Melt and Sol-Gel Derived Bioglass^ In-Vitro: D.C. Greenspanet al.
360 300 240
I
180 +
2
4
6
8
10
Time (Days) Figure 1 Si Release of 58S vs. 45S5
70
1 1 :£
50 -d
30 10 YJT m
^ ^ -10
H
^?
^
l’^"**^^ ^
-30 Time (Days) 58S in SBF
Time (Days) 45S5 in SBF
Fig. 2 Weight loss of Si, Ca and P of 58S versus 45S5 in SBF
^
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4 45S 5 Unreacte d % R
2
r\
si-o-si
588 Unreacte d
Si-O-Si
45S 5 SBF 6 hours % R
% R
2
5
45S 5 SBR.2 4 hours / \
HCA
100 0 Wavenumbers (crrvl)
100 0 Wavenumbers (cm-1 )
Fig. 3 FTIR reflection spectra of 58S versus 45S5 in SBF CONCLUSION S The 588 sol-gel derived bioactive glasses had a much higher in-vitrodegradation rate and formed HCA faster than the 45S5 melt derived material when reacted in SBF. Recent in vivo studies indicate enhanced bone mineralization for sites treated with these sol-gel derived materials when compared to melt derived material[13]. The ability to control the degradability of these porous glasses by altering their composition and pore structure justifies further studies of these materials for use as a resorbable synthetic bone graft particulate. REFERENC E 1. Hench, L.L., J. Am. Ceram. Soc, 1991, 74[7] 1487-510 2. Pereira, M.M. et al., J. Am. Ceram. Soc, 1995, 78[9] 2463-68 3. Wilson, J., et al., "Bioglassfi Middle Ear Devices - The Year Clinical Results," Present at 2P’ Society for Biomaterials annual meeting, March 18-22, 1995, San Francisco, CA 4. Stanley, H.R., et al.. Encyclopedic handbook of Biomaterials and Bioengineering, Part B: Application, Vol. 2, eds. D.L. Wise et al.. Marcel Dekker, 53, 1995 5. Zamet, J.S., Darbar, U.R., et al., J. Dent. Res., 76 (lADR Abstracts) #2219, 291, 1997 6. Hench, L.L., and LaTorre, G.P., Bioceamics 5, 1992, 67-74 7. Wilson, J., Nolletti, D., CRC Handbook of Bioactive Ceramics Vol. I, Bioactive Glasses and Glass-Ceramics, eds. T. Yamamuro et al., CRC Press, 283, 1990 8. Filgueiras, M.R., et al., J. Biomed. Mater. Res., 1993, Vol. 27, 445-53 9. Li, R., et al., J. Appl. Biomater., 1991, 2, 231-39 10. Kokubo, T. et al., J. Biomed. Mater. Res., 1990, 24, 721-34 11. Warren, L.D. et al., J. Biomed. Mater. Res.: Applied Biomaterials, 1989, Vol. 23, No. A2 201-209 12. Clark, D.E. et al., "Corrosion of Glass" Books for Industry and Glass Industry, NY, 1979 13. Wheeler, D.L. et al., "/« vivo Evalution of Sol-Gel Bioglassfi, Part I: Histological Finding," Presented at Society for Biomaterials annual meeting, April 30-May 4, 1997, New Orleans, USA
COMPOSITE CERAMICS
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
UPGRADIN G OF HYDROXYAPATIT E CERAMI C INCORPORATIO N OF a-TRICALCIU M PHOSPHAT E
BIOCOMPATIBILIT
Y
BY
S. Sarig, F. Apfelbaum, and F. Kahana Casali Institute of Applied Chemistry, School of Applied Science and Technology, The Hebrew University of Jerusalem, 91904, Israel
ABSTRAC T The object of the present study was to produce a novel composite of apatite-a-tricalcium phosphate (HA a-TCP) with improved biocompatibility as compared to available commercial products. The calcium phosphate powder was precipitated from dilute solutions irradiated by microwave. The XRD of the powder was remarkably similar to that of bone mineral. a-TCP was detected in ceramic sintered from this powder at 700^C. Biological analyses confirmed the enhanced biocompatibility of the ceramic in tissue culture experiments. KEYWORD S a-Tricalcium Phosphate, low temperature ceramic, enhancement of biocompatibility. INTRODUCTIO N Hydroxyapatite (HA)ceramic is preferable as bone graft substitute due to its similarity to the chemical composition of bone mineral and subsequent biocompatibility and lack of toxicity. However, most investigators believe that calcium phosphate materials are devoid of intrinsic osteoinductive properties. Recently it has been shown that ceramics containing a-TCP evoked bone formation in the pores of the ceramic (1). a-TCP crystallizes in the monoclinic space group W / a , each formula unit occupying 180 A^. b-TCP crystallizes in the rhombohedral space group RSc, and its unit formula occupies 168 A"’. a-TCP is therefore a "looser" structure than b-TCP and has a higher internal energy (2) which entails higher reactivity in aqueous media and enhanced resorbability. The crystal structure of a TCP has been worked out in detail (3). Phase diagrams indicate the thermodynamically stable pure phases transition ( a - to b-TCP in the present case) at 1125^C. However, the actual phase that forms under any given condition is dictated by kinetic rather than thermodynamic considerations. In the present case the decisive factor may be the degree of crystallinity of the source material. MATERIAL S AND METHOD S The powder was precipitated from CaCl2 10 mM solution (A) and NaH2P04 6 mM solution (B). Solution A was used without admixture, whereas solution B was made up to concentration of 25ppm with respect to L-aspartic acid and ISOppm with respect to NaHCOa. Equal 397
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volumes of solutions A and B (250 ml each) were introduced simultaneously into a 1000 ml beaker and put into a microwave oven (Sharp 700W) at maximum power for 5 minutes. The irradiated mixture was then quenched in ice for 30 minutes and the precipitate fihered through a milipore filter (0.45m), washed, and dried overnight at 55^ C. For the preparation of the ceramic, the powder was closely mixed with ammonium hydrogen carbonate (NH4HCO3), in proportion of 20% per weight. Addition of the carbonate at this stage was made to provide the material with suitable porosity after the ammonium carbonate decomposes on heating. A part of the mixture was ground for 2 mins. Then, portions of 200 mg each were introduced into a pellet dye (Graseby Specac) and compressed under pressure of 4 tons/cm^. Series of the so formed pellets were heated to 500, 700 and 900^C and kept at the respective temperature for one hour. RESULT S AND DISCUSSION . The XRD pattern of the powder prepared according to the above procedure is shown in fig. 1. This pattern is notably
1588 [counts ]
Pouder froM usual
procedur
1888
588
1888
588 -\
8 .8
58
[*2e]
68
Fig. 1. XRD pattern of the powder precipitated while irradiated by microwave. This pattern is similar to the XRD of bone mineral. Fig. 2. XRD pattern of the ceramic sinteredfi*omthe above described powder at 500^C. Note that the undifferentiated group of peaks between 31.77 and 32.9 remain unresolved. similar to that of bone mineral (2). It is interesting to note the lack of background in the XRD pattern ( fig 1 ), which could have indicated the presence of amorphous calcium phosphate. When heated to 500 C, the special crystallinity did not change (fig.2). Only when heated to 700 C, transition to b-TCP and hydroxyapatite took place. But the most interesting feature of this transition was the appearance of a-TCP as evidenced by the peak at 30.71 . It was fiirther
Upgrading of Hydroxyapatite Ceramic Biocompatihility:S. Sarig et al.
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sustained by two smaller characteristic peaks at 24.1 and 12.1 (fig-3) which are absent in both apatite and b-TCP diffractograms. When heated to 900 C, there was transition to b-TCP and to hydroxyapatite, with improved crystallinity, but a-TCP peaks were absent.
4B8B [ counts] 3500 -I
a-TCP
3000 2500 2000
a-TCP
a-TCP
1500 A 1000 -\
0 .0
JL^J^^JJUJ ^ n
10
1 20
1
1 30
r
50
[•29]
60
Fig.3. XRD pattern of ceramic from the microwave precipitated powder, sintered at 700 C. Note the differentiation of the 31.77, 32.2 and 32.9 peaks and the 30.7, 24.1 and 12.1 peaks characteristic to a-TCP. It was established that other modifications of parameters such as increase in the concentrations of the initial solutions of calcium chloride and sodium phosphate did not produce the powder which upon sintering yielded the special spacious a form. Other deviations from the procedure described above such as increasing the volume of employed solutions preserved the a forming potential, provided the irradiation period was extended in proportion to the increase of volume. When the volume was increased, for instance, sixfold and the irradiation lengthened to 30 minutes, the ceramic obtained on heating to 700 C contained a-TCP, but when the irradiation time was 15 minutes, the spacious form did not appear in the ceramic. The a-form has not been obtained from powders obtained without irradiation. In the literature thermal crystallization of a-TCP at temperature as low as 600 C was addressed. However, it was achieved in a single phase, from amorphous calcium phosphate which was obtained by a very complex procedure (4). It seems that the right amount of energy introduced into the crystallizing system of calcium phosphate, in the present simple procedure, predisposed the precipitate to form later the spacious a-form, stabilized in the solid phase. The superior reactivity of the composite ceramic, as recognized in tissue culture experiments (5) should be advantageous in its use as bone graft substitute.
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SUMMAR Y Calcium phosphate precipitate was formed from dilute solutions under microwave irradiation. The XRD of the powder had apatitic character with undifferentiated peaks and was similar to bone mineral pattern. On sintering at 700^C transitions to other phases were obtained, among them the high temperature, reactive a-TCP. Deviations from the experimental parameters of preparation such as: concentration of the source solutions, temperature of sintering, microwave enegy input per volume of solution, prevented the appearance of the a-form. The ceramic containing a-TCP was found to be superior to commercial materials in tissue culture experiments. ACKNOWLEDGMEN T The authors wish to thank the "Yeshaya Horowitz Association" for the support of this study. REFERENCE S 1. Klein C , de Groot K., Weiqun C , Yubao L.,and Xingdong Z., Biomaterials1994, 15, 31-34 2. Elliot J.C, Structure and Chemistry of the Apatites and Other Calcium Orthophosphates, Elsevier Science B. V. 1994 3. Mathew M., Schroeder L.W., Dickens B., and Brown W. E. Acta Cryst. 1977, B33, 1325-1333 4. Kanazawa T., Umegaki t., and Uchiyama N., JChem.Tech. Biotechnol 1982,32,399-406 5. Ben-Bassat H., Klein B. Y., Lemer E., Azoury R., Rahamim E., Shlomai Z., and Sarig S. Cells and Materials,1994, 4, 37-50
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PREPARATIO N OF HYDROXYAPATITE/COLLAGE
COMPOSIT E MATERIAL S CALCIU M N BY COPRECIPITATIO N METHO D
O.I. Slivka, V.P. Orlovskii Kumakov Institute of General and Inorganic Chemistry of Russian academy of sciences, Moscow, Leninskii pr. 31, 117907 Russia.
ABSTRAC T Interaction of calcium hydroxyapatite and collagen by coprecipitation method at the 20-450 Q temperature interval (pH=ll-12) m system CaCl2 - (NH4)2HP04 NH4OH - H2O - collagen was investigated. It was shown by the IRS and X-ray difl&action analysis that the nonsintered solid phases are consisted of collagen and poorly crystalline calcium hydroxyapatite without the other calcium phospate impunities. The difference between the crystalUnity degree of solid phases depended on temperature and initial form of collagen in reaction mixture was found. KEYWORDS : calcium hydroxy^atite, collagen, coprecipitation, composite. MATERIAL S Aqueous solutions of CaCl2, (NH4)2HP04, ammonia and collagen in powder form or in soluble form (solution in acetic acid) were used as initial components. Interaction in the system was studied at a variable amounts of collagen and at variable temperature (20 and 40-45^ C), at pH-10. Method of hydroxyapatite cc^recipitation at the stoichiometric molar ratio Ca/P=1.67 [see ref]. RESULT S The analysis of nonsintered composites by IRS and small-angle XRDS has shown that the general hydroxyapatite structure was kept. The analysis of IR-spectra has not shown the remarkable distinctions between the composites obtained under different conditions, namely: reaction temperature, forms of protein’s introducing in reaction mixture and the protein concentrations. However it was found by the powder X-ray diffraction analysis that: 1) all nonsintered composites has shown the amorphous structures but with all basical peaks characteristic of the calcium hydroxyapatite and with some additional peaks probably concerning to collagen; 2) at 40-45^ C and the introducing of collagen in the soluble form the crystallinity of corresponding composites was higher in comparison with the other specimens. 401
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SUMMAR Y The composite material consisting of calcium hydroxyapatit e and collagen was obtained. The powder XRD-study has shown that this product had an amorphous structure. The conditions of coprecipitatio n guarantee d a some higher crystallinity degree of nonsintere d product were found. The percent content s of Ca and P in the composite are correspondin g to the ones in a stoichiometri c hydroxyapatite , but their molar ratio Ca/P is less than stoichiometri c value. Preliminary biological testing of this composite has given a good results. REFERENCE S Orlovskii, V.P., Ezhova, ZhA , Rodicheva, G.V., Koval’, E.M., Sukhanova, G.E., and Tezikova, L A , ZA. Neorg,Khim,1992, 37, 4, 881.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BON Y REACTIO N OF SEVERA L KIND S OF Ca^P-COLLAGE N CONJUGATE D SPONGE S H. Oonishi\ F. Sugihara^, K. Minamigawa^, Y. Mandai^, K. Nagatcani^, S. Kushitani^ H. Iwaki^ N. Kin^andE. Tsuji^ ^ Department of Orthopaedic Surgery, Artificial Joint Section and Biomaterial Research Laboratory Osaka-Minami National Hospital, 2-1, Kidohigashi-madii, Kawachinagano-Shi, Osaka, 586, JAPAN ^ Nitta Gelatin Inc., Research Laboratory, Osaka, JAPAN ^ Osaka Prefectural Industrial Engineering Research Institute, Osaka, JAPAN ABSTRAC T Several kinds of caldtim jAosphate (a-TCP, TeCP or OCP)-coll^gen conjugated sponges(Ca»Psponges) were pr^arcd by the reaction of calcium phosphate(Ca»P) showing hydration activity with type I collagen able to form a gel in a short time under the physiological condition. As a control, HA-oollagen sponge was also pi^aied similaily. The objective of this stucfy is to find out a reasonable Ca»P-sponge to bony tissue reaction as wdl as the hemostatic effect on bleeding bone. The collagen gel was prq)aredfrom solutions of pepsin-solubilized type I collagen in BGJb medium in the pies^ice of each Ca»P to obtain Ca»P-sponges. Prior to implantation, XRD and SEM observations revealed that the Ca»P-sponge was an ipatite-collagen composite having a honeycomb structure with pores of microns in diameter. The sponges were filled into 5 mm holes drilled in the femoral condyle of mature rabbits. At one week after implantation, only OCP-sponge hadbegimto be incoiporated into bone and to be iq)laced by bone. At three weeks, OCP and a-TCP-sponges formed trabecular bone, TeCP- sponge was incorporated into bone and HA-sponge did not contactwith the bone. At six weeks, HA-collagai had been incorporated into new bone only at its periphery. OCP-oollagen was incoiporated into the new bone eariiest and showed the best biocompatdbility. All the sponges showed good hanostatic effect. In conclusion, OCP-sponge has both the b^t osteoconductive and resorbable characteristics and good hemostatic effect. KEYWORDS : a-TCP, OCP, TeCP, low crystalline apatite, apatite-coUagoi conjugate, bioresorption INTRODUCTIO N How can we make a bone-replacement material using calcium phosphate(Ca»P) and collagen? A composite of colls^oi and Ca*P might be quite useful as a hemostatic filling material which is fittable to any type of bone defects. In addition, if the composite has a high biocompatibility, it would be expected that the filled is refrfaced by new bone, partially or completely. The sponge-like composite was devdopedby the reacticm of chemical active Ca»P showing hydration activity with type I collagen able to form a gel in a short time under the physiological condition. At 9th this symposium, we named thisreactionsystem as a "dynamic biomimetic method". ^ using this m^od, we have rqx)ited that tetracaldum phosjAateCTeCPXl]-, a-tricaldum phosphate(a-TCP)[2-4]- andoctacalciumphosphate(OCP)[4-6]-collagen conjugated sponges had both 403
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collage n solutio n [1 % (w/v)]
[B] [C]
SxBQJ b medium containin g Ca»P particle s [2.5 % (w/v) OCP or 5.0 % (w/v) TeCP, a-TCP an d HAl reconstitutio n buffer (HEPES ) solutio n mixed col d [Al, [B] and [C] at a ratio of 7:2: 1
i
incubate d in a STC incubato r for 3 days
i
lyophiKzed
i
UV ray-inradiate d for 2 hrs .
Figure 1. Method of preparing (3a»P-collagen conjugated sponge desired osteoconductive characteristics and local hemostatic effectivaiess. The objective of this study is to find out a reasonable Ca»P-sponge to bony tissue reaction as well as the hemostatic effect on bleeding bone, comparing with a control, HA-collagen sponge[l], qualitatively. MATERIAL S AND METHOD S The method of preparing calcium phosphate(Ca»P)-oollagen conjugated sponge and HAcollagen sponge is shown in Figure 1. A pepsine-solubilized atelopq>tide collagen solution(Ce// matrixLAfi, NittaGelatin Inc., Osaka, Japan), a mixtureofB(jJbmedium(SIGMACo., St. Louis, USA)and Ca«P and a reconstitution buffer solution of HEPES(Dojin Chemicals Co., Kumamoto,
Figure 2. SEM jAotograjA of a-TCPsponge matrix
Figure 3. Micrograph of decalcified section 6 weeks after implantation of TeCPsponge.M; matrix of TeCP-sponge, N; new bone, arrows; osteoblasts, arro^iiead; osteocyte (H.E. stain, x400)
Bony Reaction of Several Kinds of Ca-P-Collagen ConjugatedSponges: H. Oonishi et al.
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Japan) were mixed under an ice-cooled condition and then incubated in a CO2 incubator at 37 C, and the resulting gel was lyqphilized to obtain spcMige-like products(OCP-sponge, TeCP-sponge, aTCP-sponge andHA-sponge). To strengthen the cross-linking force of the collagen fiber, 20W UV ray irradiation was carrfed out for 2 hrs. Theparticle sizes of all Ca»Ps used were 10 i-im. Prior to implantation, the obtained sponge-like composites weie analyzed by means of X-ny diffraction(XRD) and scanning electron microscq)e(SEM). These sponges were filled into a hole of 5 mm in diameter drilled bilaterally through femoral ccxidyle of mature rabbits, and the hemostatic effect was evaluated during sui^ery.The rabbits were sacrificed by intravenous injection of overdoses of Nembutalfiat 1, 3 and 6 weeks after implantation of the sponges.The condyles w«e trimmed out en blockincluding the implants and immediatdyfixedin 10 % {^osphate-neutral formalin solution. Decdcified and undecaldfied sections were made for light microscopic, SEM and backscattered electron imaging observation. RESULT S AND DISCUSSIO N Structure of sponges and hemostati c effect OCP-, a-TCP- and TeCP-sponges were found to be low oystalline gpatite-oollagen conjugates with traces of eadi Ca»P. The ciystal structure of HA was undianged in HA-sponge. All sponges had hcmeycomb structures with pores of microns in diameter. At higher magnification of OCP- anda-TCP-sponge matrices, it was observed that the agglomerates, which were composed of sand-rose sh£q)ed lamdlae of low oystalline apatite-collage conjugate, gathered with collagen fibers(Fig. 2). The matrices of OCP-, a-TCP-, and TeCP-sponges were strongly suggested to be reconstituted apatite-collagen composites like a native bony matrix.
Figiu^ 4. Micrograph of decalcified secticoi 6weeks afterimplantation of a-TCP-sponge. N; new bone, P; low oystalline £^atite particles, arrows; osteoblasts, curved arrow; multinucleated giant cell (H.E. stain, x400)
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Complete hemostasi s was obtained immediatel y after pressing with each sponge on bleeding bone. This hanostasis might be induced by the pressing effec t as well as the hemostati c mechanism due to the platelet aggregation ability of the collagen in the sponges. Comparison of bony reaction At one week after implantation , cmly OCP-sponge had b^gun to be incorporate d into and replaced by new bone. At three weeks, OCP-sponge formed laminar new bone of 100-70 0 \imin dqpth from the defecte d bony wall, a-TCP- and TeCP-sponges contacte d with the wall and were incorporate d into the new bone at their peripheries . On the other hand, HA-sponge did not contact with the bone still 3 weeks postoperatively . At six weeks, HA-sponge was incorporate d into the bone only at its periphery and a laige amount of HA still ranained at the central area.In growth of bone trabeculae was observed in the caiter of OCP-, a-TCP- and TeCP-sponges, and almost all the low oystalline apatite-particle s had been incorporate d into than 6 weeks postoperativdy . In TeCPsponge, active osteoblast s were observed at the edge of its matrix(Fig. 3). Especially for OCP - and a-TCP-sponges, multinucleate d giant cdls existe d adjacent to the gathered partides after 3 weeks. The degradation of these particles was likdy to dep«id on solution- and cell-mediate d processes . The latter process was bioresorptio n by multinucleate d giant cells interactin g with osteoblast s (Fig. 4) and probably by macrophages . It was hypothesize d that the low oystalline apatite partides bound with osteocalcin[7,8 ] were resoibed by Htvdosteoclasti c cells and incorporate d into the trabeculae by osteoblast s within 6 weeks after implantation . CONCLUSIO N Eveiy kind of Ca»P-collagai sponges showed good honostatic effect . OCP-collagen conjugate d sponge has both the best osteoconductiv e and resorbable characteristics , incorporate d into and replaced by new bone earliest.
REFERENCES
1. Sugihara, F., Minamigawa, K., Oonishi, H., Mandai, Y., Nagatomi, K., Yasukawa, E., Kushitani, S. andTsuji, E. In: Bioceramics Volume 7, Butterworth-Heinemann , Oxford 1994, 193-198 . 2. Toda, I., Kitayama, N., Sugihara, F., Minamigawa, K,, Gonda, Y., Suwa, F. and Oonishi, H. y. Jpn.Soc,Biomat., 1996, 14, 112-11 6 3. Toda, I., Kitayama, N., Sugihara, F., Minamigawa, K.,Suwa, F. and Oonishi, H. In: Bioceramics Volume 9, Elsevier Science 1996, 465-468 . 4. Takahashi, K., J. TokyoDent Coll,Soc, 1997(in press) 5. Sugihara, F., Oonishi, H.Minamigawa, K., Mandai, Y., Tsuji, E., Yoshikawa,M. and Toda, T., In: Bioceramics Volume 9, Elsevier Science 1996, 399-402 . 6. Kitayama N., J. Jpn.ProsthodonU Soc, 1997, 41, 86-93 7. Glowacki, J., Rey, M., Glimcher, M J., Cox, K.A. and Lian, J., J, CellBiochem.1991, 4 5, 292-302 . 8. Chenu, C , Colucci, S., Grano, M., Zigrio, P., Barattole, R. andZallone, A.Z., 7. CellBioL 1994, 127, 1149-1158 .
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
In Vitroand In VivoTests of Newly Developed TCP/CPL A Composites
Masanori Kikuclii’, Sung-Baek Cho^ Yasushi Suetsugu’, Jun/oTanaka’ Takayuki Kobayashi^, Masaru Akao\ Yoshiliisa Koyama^ and Kazuo lakakuda^ ^National Institute for Research in Inorganic Materials, l-l Namiki, Isukuba, Ibaraki 305, Japan. ^Division of Inorganic Materials, Institute for Medical and Dental Engineering, Tokyo Medical and Dental University, 2-3-10 Kanda-Sumgadai, Cliiyoda-ku, Tokyo 101, Japan.
ABSTRAC T Biocompatibility of p-tricalcium phosphate co-polymerized polylactide composite was tested in vitroand in vivo. The composite formed carbonate hydroxyapatite microcr) stals on its surface when soaked in a simulated body fluid and showed no cytotoxicity in a cell culture test using MC3T3-E1 cells. Animal experiments using German shepherds for oral applications and beagles for orthopedic application suggested that die composite is applicable as a guided bone regeneration membrane and bone substitute. KEYWORD S ptricalcium phosphate, co-polymerized polylactide, composite, cell culture test, guided tissue regeneration, guided bone regeneration INTRODUCTIO N ptricalcium phosphate (TCP) is well known as a bioactive ceramic and is used in medical and dental lields. The dense ceramic, when implanted into a bone, can be substituted by biological bone; however, this ceramic is britde and needs a long time for the substitution. It can be simply used as a structural material. On the other hand, some kinds of polyester are known to be biodegradable biomaterials; however, such polymers are too soft for artificial bone materials and have no osteoconductivity. In a previous study [I], the authors reported the preparation and mechanical properties of calcium phosphates CPLA(co-polymerized polylactide)[21 composites which had good properties for artificial bone materials. In this study, the TCP (TLA composite was tested in vitro and in vivo to check its biological properties. MATERIAL S AND METHOD S TCP was prepared by a wet mediod using Ca(OH)2and RTO4 as starting materials. The precipitates obtained were dried at 293K for 24 hours and calcined at 1073K for 3 hours. CPLA was prepared by the co-polymerization of L-lactide with fatty polyesters. The mechanical properties and solubility of CPLA were controlled by regulating the amounts and species of added polyesters. Hard (CPLA-H) and soft (CPLA-S) types of CPLAs, which have a mean molecular weight of 100,000 to 200,000, were used in this study. TCP and CPLA were mixed by a thermal kneading method using a LABO PLASTOMILLfi (Toyo Seiki). CPLA was melt in the plastomill at 453K for 1 minute, and TCP 407
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powder preheated at 453K was added and mixed with the CPLA melted for 10 minutes at 20 rpm for CPLA-S and at 5 rpm for CPLA-H. Then, the composites obtained were formed into a plate, a block, a circular prism and a sheet by a heat press at 453K and kept at 293K for 30 minutes to crystallize CPLA. The molecular weight change of CPLA after the thermal treatment was evaluated with gel permeation chromatography. The three-point bending strength was measured with a universal testing machine of 15 mm in span and at lOO/^m min crosshead speed using samples of 3x5x20 mm^ in size. The samples were cut with a diamond saw using ethanol and soaked in physiological saline. Simulated body fluid(SBF) soaking and cell culture tests were performed for plates of 3x5x10 mm^ in size. Thin film X-ray diffraction patterns and infrared spectra of composite surfaces were observed after the SBF soaking test. Cytotoxicity was tested by observing cellular morphology and cell growth rate of MC3T3-E1 cells. Composite sheets were implanted as oral guided tissue regeneration(GTR) and guided bone regeneration(GBR) membranes after shaping during surgical operation. The tissue guided around implants were extracted at 4,8 and 12 weeks after implantation and observed histologically. Blocks, 4x4x 12 mm"^ in size, and circular prisms, 4
1400
1200
1000
800
wave number / cm^
Figure 1. Infrared spectra of TCP/ CPLA composite surface after soaking in simulated body fluid.
Figure 2. Phase contrast micrograph of MC3T3-E1 cells cul› tured with TCP CPLA composite at 4 days after incubation.
In Vitro and in Vivo Tests of Newly Developed TCP/CPLA Composites:M. Kikuchi et al.
control
10"
O
409
TCP/CPLA
10^ 0
1
2 Time / day
3
4
Figure 3. Growth curves for MC3T3-E1 cells. composite had higher bioactivity than TCP ceramics. Phase contrast micrograph and cell growth curves for MC3T3-E1 cultured with the composite are shown in Figs. 2 and 3. Cells were in close contact with the composite and no morphological change was observed. The growth curve was almost the same as that obtained for the blank control. These results suggested that the composite had no acute cytotoxicity.
Figure 4. Histological section of periodontal tissues 12 weeks after GTR operation stained by villanueva bone staining.
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Figure 5. Histological section of jawbone 12 weeks after GBR operation, stained by villanueva bone staining. Figures 4 and 5 show the histological section of GTR and GBR experiments at 12 weeks after operation (villanueva bone staining). In Fig.4, the upper side of the arrows indicates regenerated tissues in which dentine, cementum and periodontal bone were found; further, fibrous tissue, which will probably change into a periodontal membrane, was regenerated. Bone regeneration took place more clearly in jawbone, as shown in Fig.5; the composite membrane could regenerate bone defects of 10 mm^ in area and 10 mm in depth. In a series of animal tests, pure CPLA membrane could not regenerate a similar jawbone defect. Though the composite membranes were degraded in 4 weeks after implantation, the bone defects were regenerated up to their original height. Accordingly, TCP powder in the composite could plausibly enhance cell differentiation by releasing Ca and PO4 ions. These results indicate that the TCP/CPLA composite is utilizable as GTR and/or GBR membrane. CONCLUSIO N TCP/CPLA composite had good biological properties and was easily shaped during surgical operation. The bioactivity of the composite was much higher compared to pure TCP ceramics and had no cytotoxicity. Membranes enhanced tissue regeneration for a dental defect and bone regeneration for a jawbone’s defect, probably due to Uie solution of TCP. The composite will be utilizable not only as a bone substitute but as a GTR and GBR membrane. ACKNOWLEDGMEN T The CPLA plastics were provided by Dainippon Ink Chemicals, Inc. Ihe authors are grateful to Mr’s. Yutaka Tashiro, Yasutoshi Kakizawa and Toshiki Shikata at Dainippon Ink Chemicals, Inc. for supplying CPLA and for measuring its molecular weight
REFERENCES
1 Kikuchi, M , Suetsugu, Y., Tanaka, J. and Akao., M. Bioceramics,Volume 9, Pergamon, Oxford 1996, 395-398. 2 Kakizawa, Y, Kagaku to Kogyo, 1995, 48 [9], 1070-1072.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
OCCLUSIO N OF DENTI N TUBULE S BY 45S5 BIOGLASS fi Leonard J. Litkowskil, Gary D. Hackl, Hollie B. Sheaffer^ and David C. Greenspan^ 1 Department of Restorative Dentistry, University of Maryland Dental School, 666 West Baltimore Street, Baltimore, MD 21201, USA 2 USBiomaterials Corporation, One Progress Boulevard, Alachua, FL USA
ABSTRAC T The currently accepted theory for tooth hypersensitivity is the hydrodynamic theory based on the belief that open dentin tubules allow fluid flow through tubules exciting nerve endings. Clinical replicas of hypersensitive teeth reveal numbers of open tubules. It is the purpose of the present study to evaluate Bioglassfi compositional, size range and techniques to enhance the bonding to dentin surfaces and occluding open tubules. Experiments were performed using standardized human dentin slabs in vitro cut from extracted teeth. SEM and FTIR evaluations were performed on the dentin surface after treatment with Bioglassfi compounds. Results showed an increase in tubular occlusion compared with non-Bioglassfi containing controls. KEYWORD S Bioglassfi, Tooth hypersensitivity. Dentin, Dentin occlusion INTRODUCTIO N Tooth hypersensitivity is a common problem which affects about 40 million adults in the United States, 10 million of which can be considered chronically affected (1). It is estimated that some 17% of adults in the U.S. have at least one or more sensitive teeth. The teeth may be sensitive to cold, heat, air or sugary foods. The incidence of tooth hypersensitivity increases with age. The increased incidence is believed to be related to the general increase in exposed root surfaces of teeth as a result of periodontal disease, tooth brush abrasion or cyclic loading fatigue of the thin enamel near the dento-cementum junction. The currently accepted theory for tooth hypersensitivity is called the hydrodynamic theory. This theory is based on the belief that open dentinal tubules allow fluid flow through the tubules. This flow excites the nerve endings in the dental pulp. Clinical replicas of sensitive teeth viewed in the scanning electron microscope (SEM) reveal varying numbers of open or partially occluded dentinal tubules. Tubules generally are not seen at the tooth root surface because of the cementum covering the tooth root, or because of a smear layer of dentinal debris 2-5 microns in thickness that covers the tooth surface and masks the tubules. When the smear layer of the tooth is present, the fluid flow that can occur through the dentin is only a few percent of that possible following acid removal of the smear layer, thereby "opening" the tubules. There is a growing body of evidence that occlusion of the dentinal tubules of a sensitive tooth, whether by resin infiltration, varnish coat or more recently by crystallite precipitation, results in reduction or elimination of the hypersensitivity. The duration of relief, however, is highly variable. Hypersensitivity usually reappears because of tooth brush abrasion, presence of 411
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acid challenges in the mouth or degradation of the coating material. Desensitizing dentifrices containing potassium oxalate have been found to provide temporary tubule occlusion (2). Potassium oxalate is thought to react with the smear layer to increase its resistance to acid attack as well as reduce the permeability. It is thought that the crystals produced when dentin is treated with potassium oxalate are calcium oxalate (3). Previously, all materials have used biologically inactive inorganic or organic components that will occlude the open tubules for a limited time period. Normal habits including the eating of acidic foods and vigorous toothbrushing will remove the materials from the tubules allowing fluid flow and a recurrence of sensitivity. In addition, literature has demonstrated with some hypersensitivity agents that simple rinsing with water significantly reduces the number of occluded tubules. Therefore, there is a need in the dental field for a material that would chemically react with the surface of dentin and intimately bond to tooth structure, which would significantly reduce the reopening of dentin tubules due to contact with oral fluids and potentially remineralize the surface. Bioactive and biocompatible glasses have been developed as bone replacement materials. Studies have shown that these glasses will induce or aid osteogenesis in a physiologic system (4). The bond developed between the bone and the glass has been demonstrated to be extremely strong and stable (5). The bioactive glass formulation has been widely tested in bone and soft tissue. Toxicology evaluation of the glasses has shown no toxic effects in bone or soft tissue in numerous in vitro and in vivo models (6). However, the glass has been reported to be bacteriostatic or bacteriocidal most likely related to the change in pH induced by the dissolution of the ions from the surface of the glass and lack of bacterial adherence to the glass surface (7). The bonding of the glass to bone begins with the exposure of the glass to aqueous solutions. Na+ in the glass exchanges with H+ from the body fluids causing the pH to increase. Ca and P migrate from the glass forming a Ca-P rich surface layer. Underlying this Ca-P rich is a layer which becomes increasingly silica rich due to the loss of Na, Ca and P ions, which will form an apatite layer. The utilization of bioactive glasses for bonding to tooth dentin has not been previously described. It is the object of this study to apply Bioglassfi containing material to dentin surfaces to evaluate the occlusion of tubules to reduce or eliminate hypersensitivity and quantitate remineralization. An evaluation of particle size, carrier vehicle, contact time and concentration for the most efficacious delivery system is currently on-going. Determination of the tubular occlusion was evaluated by SEM and remineralization by Fourier Transform Infrared Spectroscopy (FTIR). MATERIAL S AN D METHOD S in vitroexperiments were performed using a standardized slab of human tooth dentin from extracted teeth. These discs were cut from the extracted teeth using an Isomet diamond saw (Buehler Ltd.). The discs were 1.0 mm thick and the size of the tooth. The occlusal surfaces were ground on a series of wet silicon-carbide papers ranging from 320 to 600 grit. This was done to standardize the test surfaces. The surfaces were treated with 37% phosphoric acid for 60 s to remove the smear layer created during the grinding process and open all the dentin tubules(see figure 1) The surface was rinsed with distilled water for 20 s and dried with a blast of oil free air. Each slab was split in half and the experimental material placed on one-half of the specimen and the other half used as the etch control. Scanning electron microscopy was performed on the treated surface in each group. The slabs were mounted on scanning electron microscope stubs using silver paste. All specimens were
Occlusion of Dentin Tubulesby 45 S5 Bioglass^^:L.J. Litkowski et al.
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vacuum dried, sputter coated and examined in a JEOL-T200 scanning electron microscope. Fourier Transform Infrared Spectroscopy (FTIR) was performed on samples treated as above. The spectra are presented in Figure 4. RESULT S AN D DISCUSSIO N Figure 1 is a representative view of open tubules of the etched dentin prior to application of the material. Figure 2 is the treated side of the same sample after two minute application of the Bioglassfi compound. As can be seen there is an almost total occlusion of the tubules. Fourier Transform Infrared Spectroscopy (FTIR) was performed on samples treated as above. The spectra are presented in Figure 3. Sample 1 is the control and is a representive view of the spectrum of hydroxy carbonate apatite. The shape of the peaks between wave number 1150 to 500 is very characteristic. In sample 2 the peaks are disrupted after treatment with the acid etchant, especially in the less than 800 range. This is indicative of a loss of the mineral components of the tooth structure, Calcium and Phosphorous. Sample 3 shows a partial remineralization of the Ca and P on the tooth structure. This was treatment of two minutes with a small particle form of the Bioglassfi compound. Sample 4 was treated with the optimal size and shape mixture of the Bioglassfi compound and shows an almost complete remineralization of the tooth surface. This is also the formulation that was used in Figure 2.
Figure 1: Acid etch surface of a dentin slab with no treatment (500x)
Figure 2: Partially occluded tubules after treatment with Bioglassfi compound.
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Figure 3. Occlusion of tubules with complete closure on multiple tubules.
Figure 4: FTIR spectra, dentin slabs. Top to bottom: control, acid etch, partial treatment, optimal treatment.
SUMMAR Y It can be seen from the data presented that there is significant occlusion of tubules with the Bioglassfi compounds. In addition, early stage remineralization of the tooth surface can be seen with FTIR. In conclusion, incorporation of Bioglassfi into a compound for treatment of sensitive teeth could be efficacious based on in vitrodata. It will be necessary to perform a human trial on patients with hypersensitive teeth in order to confirm this hypothesis clinically. ACKNOWLEDGMEN T This study was partially funded by USBiomaterials Corporation.
REFERENCES 1. 2. 3. 4. 5. 6. 7.
Kanapka, Dent. Clin. North Am.,34:54 (1990)). Pashley et al., J. Periodont.,55:522(1984)) Pashley et al, Arc/z. Oral Biol., 30:731 (1985)). Hench etal, J Biomed.Mater.Res. 5:117-141 (1971) Piotrowski et al., J. Biomed. Mater. Res. 9:47-61 (1975)) Wilson et al., J. Biomed. Mater. Res. 805-817 (1981). Brink, M., et al. Bioceramics Volume 8 Hench & Wilson Eds.Elsevier Science Ltd, Alden Press, Oxford 1995 pp.471-476.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BIONER T AND BIODEGRADABL E POLYMERI C MATRI X COMPOSITE S FILLE D WIT H BIOACTIV E Si02-3CaO PzOg-MgO GLASSES AND GLASS-CERAMICS R. L. Reis^’^ A. M. Cunha^ M. H. Feraandes^ R. N. Correia^ ^ Dept. of Metallurgical Eng., Univ. Porto, FEUP, Rua dos Bragas, 4099 Porto Codex, PORTUGAL INEB - Institute for Biomedical Engineering, Pra^a Coronel Pacheco 1, 4050 Porto, PORTUGAL ^ Dept. Dept. of of Polymer Polymer Eng., Eng., U. U. Minho, Minho, Campus Campus de de Azurem, Azurem, 4800 4800 Guimaraes, Guimaraes, PORTUGAL PORTUGAL Dept. of Ceramics and Glass Eng., U. Aveiro, Campus de Santiago, 3800 Aveiro, PORTUGAL 2
ABSTRAC T In this work bioactive glasses and glass-ceramics in the Si02-3CaOP205-MgO system were incorporated, in weight fractions up to 30%, into two matrixes: ultra-high molecular weight polyethylene (UHMWPE) and biodegradable starch/ethylene-vinyl alcohol blends (SEVA-C). The composites were processed by compression and injection moulding, after a previous compounding operation. The reinforcements were characterised by DTA, XRD, laser granulometry, and SEM/EDS. Two granulometric classes, below 30 \xmand between 30 to 50 |im, were used. The mechanical properties of the moulded composites were evaluated in tensile tests, and their bioactivity was assessed by analysing the respective surfaces (by SEM/EDS and thin-film XRD) after different immersion periods in SBF. The evolution of the solution pH and Ca, P, Si and Mg concentration was followed vs. time by ICP. Due to its water uptake properties, which enhance the accessibility of the solution to the inner particles of the bioactive fillers, the SEVA-C based composites exhibit a higher tendency to form a Ca-P film on its surface. It was possible to develop composites that exhibit a bioactive behaviour associate to a mechanical performance adequate to bone replacement applications. Keywords: Bioactive glasses; glass-ceramics; polymers; composites; biodegradable; bone replacement; 1. INTRODUCTIO N The use of implants to stabilise fractures, to handle difficult bone damage, and to perform augmentation or replacement procedures has become a common method in bone surgery. In most cases, a bioactive behaviour and an adequate mechanical performance are among the most important requirements for a soft or hard tissue replacement material. The demands upon the material properties largely depend on the site of application and the function it has to replace or restore. Recently the possibility of reinforcing polymers with bioactive glasses to develop new bionert or biodegradable soft tissue bonding materials, was proposed and evaluated (1-2). In this work materials that may be usefial in a broad range of both temporary and permanent hard tissue replacement were developed. Several types of bioactive materials, namely a glass, a glass-ceramic crystallised from it and hydroxylapatite have been used as fillers of a bioinert 415
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polyethylene and a novel starch based biodegradable polymer. The aim of this research was to develop bioactive composites exhibiting an adequate range of mechanical properties. 2. MATERIAL S & METHOD S Two different polymers were used as composite matrix: ultra high molecular weight polyethylene (UHMWPE, Hostalen GUR 412, Hoechst, Germany) and a biodegradable starch/ethylene-vinyl alcohol blend (SEVA-C, Mater-Bi 1128RR, Novamont, Italy). Information on SEVA polymers, that present a range of properties very appropriate for biomedical uses, was already disclosed (3-5). Both polymeric materials were reinforced with amounts up to 30% (by weight) of a new bioactive glass and a glass-ceramic obtained from it (6,7). Hydroxylapatite (HA), obtained from Plasma Biotal, Tideswell, UK, sintered at 1250 C for Ih, was also used as reinforcement for comparative purposes. The nominal composition (wt. %) of the selected parent glass (BGEl) is 30.0 Si02 - 52.75 3CaOP205 - 17.25 MgO. The glass ceramic (BGEIC) was obtained from BGEl by a single stage heat-treatment (970 C for Ih). The transition and crystallisation temperatures were determined by differential thermal analysis (DTA), and the crystalline phases were identified by X-ray diffraction (XRD). All reinforcement materials were crushed in a ball mill to an average particle size (laser granulometric analysis) of 6.5 |Lim for HA and two granulometric classes - below 30 ^im or between 30 and 50 [im,for BGEl and BGEIC. The shape and chemistry of the fillers was studied by scanning electron microscopy (SEM) and electron dispersive spectroscopy (EDS). Composites were processed by compression moulding (UHMWPE matrix), in a SatinEnsemble press, and injection moulding (SEVA-C matrix), in a Klockner-Ferromatik Desma FM20 machine, after mixing of the respective materials in a biaxial rotating drum. Several processing conditions were studied and the respective parameters optimised. The tensile samples produced were dumb-bell shaped specimens with a cross section of 2x4 mm both for compression and injection moulding. UHMWPE was reinforced with two granulometric classes of BGEl and BGEIC, being processed by compression moulding. SEVA-C was processed by injection moulding. In this case HA was also used as a reinforcement for comparative purposes. Only one granulometric class (([) < 30 |um) of BGEl and BGEIC was used in these trials. The morphology of the produced materials was analysed by scanning electron microscopy (SEM) and optical transmission microscopy. The composites were tensile tested in an Instron 4505 machine, using a resistive extensometer, in order to determine the secant modulus (E at 1% strain) the ultimate tensile stress (UTS) and the strain at break (8r %). The cross-head speed was 5 mm/min until 1% strain and then 50 mm/min to fracture. The fracture surfaces were examined by SEM. Finally, in order to assess their in-vitrobioactivity, composite samples were immersed, for several time periods, in a comparatively large volume of a simulated body fluid (SBF) solution. The materials surfaces were then analysed by SEM/EDS and thin-film XRD (1^ incidence angle). The evolution of the solutions pH and the Ca, P, Mg and Si concentrations were followed by analysing aliquots of solution, periodically taken from the test flasks, using induced coupled plasma (ICP). 3. RESULT S & DISCUSSIO N XRD spectra showed the absence of crystalline phases on BGEl. After heat-treatment at 970 C both whitlockite and forsterite were detected (6). DTA results clearly present the correspondent two exothermic crystallisation peaks. It was possible to attain a range of mechanical properties that may allow for the use of these composites on bone replacement applications. The type of reinforcement and its
Bioinert and BiodegradablePolymeric Matrix Composites:R.L. Reis et al.
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granulometric distribution has a deep effect over the achieved mechanical properties (see Table 1 and 2). For UHMWPE matrix, the reinforcement with the glass-ceramic, associated to the smaller particle size of the reinforcement origmates the higher modulus but, as expected, a lower UTS and 8p
Table 1 -Tensile test results for UHMWPE matrix compression moulded composites. Effect of the type of reinforcement (BGEl or BGEIC) and the respective granulometry. Material UTS E,o/. e, (MPa) (GPa) (%) UHMWPE 35.6 –0.8 0.54 – 0.08 414.5 –39.6 UHMWPE + 10% BGEl {^ < 30 |am) 27.5 –3.3 1.03 – 0.23 272.9 – 48.7 UHMWPE + 10% BGEl (30 |am < cj) < 50 \im) 28.7 –3.1 0.92 – 0.23 312.3 –42.3 UHMWPE + 20% BGEl ((j) < 30 |^m) 19.4 –0.3 1.67 – 0.45 122.1 –15.5 UHMWPE + 20% BGEl (30 jim < (|)< 50 \xm) 24.0 –1.3 1.52 – 0.66 258.7 –22.8 UHMWPE + 10% BGEIC ((j) < 30 \xm) 25.3 –4.4 1.88 – 0.65 246.4 – 49.3 UHMWPE + 10% BGEIC (30 ^im < (jx 50 ^m) 21.4 –0.9 1.15 – 0.05 215.4 –22.8 UHMWPE + 20% BGEIC ((|) < 30 ^im) 18.9 –0.7 1.91 – 0.15 64.4 – 14.4 UHMWPE -f 20% BGEIC (30 |am < d x 50 urn) 19.1 –0.4 1.37 –0.53 131.0–23.8 Table 2 -Tensile test results for SEVA--C matrix injection moulded composites. BGEl ((j) < 30 jam), BGEIC ((|)< 30 um) and HA ((|) < 6.5 |am) were used as reinforcements. Material UTS E,o/, e^ (MPa) (GPa) (o/o) SEVA 37.9 –0.9 1.74 –0.08 20.02 –5.61 SEVA -C+10% BGEl 31.8–1.1 2.52 –0.32 2.65 – 0.47 SEVA -C +20% BGEl 31.3–1.5 2.85 –0.39 2.62 –0.13 SEVA--C +30% BGEl 26.9 –3.9 3.39 –0.44 2.11 –0.31 SEVA--C+10% BGEIC 34.5 –1.3 3.18 –0.36 2.01 –0.35 SEVA--C +20% BGEIC 32.4 –1.7 4.07 –0.41 1.87 –0.28 SEVA- C +30% BGEIC 31.6–3.9 4.13 –0.48 1.46 –0.28 SEVA- C+10%HA 29.0 –4.7 2.51 –0.31 2.11 –0.61 SEVA -C+20% HA 28.1 –1.8 2.80 –0.45 1.86 –0.48 SEVA C+30% HA 22.2 –4.5 3.35 –0.56 1.53 –0.44 The best mechanical results were obtained for the injection moulded SEVA-C reinforced with the stiffer reinforcing materials (BGEIC). A typical SEM micrograph of a fracture surface of one of these composites is presented in Fig.l These results correspond to an improvement of around 2 5 % on the stif&iess obtained with the much smaller HA particles. Nevertheless, SEM observation evidenced, as reported m our previous work (2), a very poor glass/polymer interaction. It is expected that the Fig.l - SEM micrograph of the fracture properties may be improved by: (i) optimising the surface of a SEVA-C+20% BGEl compounding methods used prior to injection composite. moulding (2), (ii) reducing the bioactive glass particle size or (iii) making use of coupling agents.
418 Bioceramics Volume 10 P SEVA-C+30%HA -PSEVA-C+30%BGE 1 Ca SEVA-C+30%HA A
-CaSEVA-C+30%BGE 1 -MgSEVA-C+30%BGE1|
10 Time (hours ) 10 0
1000!"
-SiSEVA-G+30%BGE1
Fig. 2 - Evolution of the Ca, P, Si and Mg elemental concentration (ICP results) in the SBF solution as a function of the immersion time for 30% BGEl and 30% HA reinforced SEVA-C composites. Fig.3 - Ca-P film formed on a SEVA-C+ 20%BGE1C composite: a) EDS spectra b) XRD spectra showing (*) T|^Pii|i|i|i|i|i|i|i|i|i[TJHi|nnni|nmpp | apatite typical peaks.
35. I 45. 50. 55. 60. ( bioactivity tests, the formation of an apatite-like layer was observed both on PE and SEVA based composites for reinforcement contents of at least 20%. The formation of this Ca-P fihn was enhanced in SEVA-C matrix composites due to its water uptaking capability which favours the dissolution of the inner bioglass particles, specially Mg and Si (ICP results confirm then* concentration increase with time, see Fig.2). Nucleation of the apatite film on SEVA-C surface corresponds to a clear decrease of the Ca and P concentration in the SBF and is likely to be related to the Si release that (Fig.2) possibly, creates sites for the formation of Ca-P nuclei. The formation of Ca-P layer on the materials surface was detected for all the reinforced materials. Thin-fihn XRD analysis (Fig.3) evidenced the presence of the (211), (102) and (210) characteristics peaks of HA. This layer presented, after a certain immersion time, Ca/P ratios in the HA range (see Fig.3). 4.
CONCLUSION S Both bioinert and biodegradable matrix composites reinforced with bioactive glasses and glass-ceramics were successfiiUy produced. The developed composites show an adequate range of mechanical properties associated to a bioactive nature (for a minimum filler amount of 20%) and, consequently, have potential for being applied as new bone-bonding replacement materials.
REFERENCES 1.
M. Wang, W. Bonfield, L. L. Hench, in Bioceramics 8, ed. L. L. Hench, J. Wilson et al., Butteworh-Heineman, New York, USA, (1995), 383 2. R. L. Reis, A. M. Cunha, S. R. Lacerda, M. H. Femandes, R. N. Correia, in Bioceramics 9, ed. T. Kokubo, T. Nakamura et al., Elsevier Sc. Ltd., Oxford, UK, (1996), 435 3. R. L. Reis, A. M. Cunha, J. Mater. Sci.: Mater, in Med,6, (1995), 786 4. R.L. Reis, A.M. Cunha, P. S. Allan, M. J. Bevis, J. Polymers for Advanced Technologies, (1996), 7, 784 5. R.L. Reis, S. C. Mendes, A.M. Cunha, M. J. Bevis, Polymer International, (1997), in press 6. J. M. Oliveira, M. H. Femandes, R. N. Correia, in Bioceramics 5, ed. T. Yamamuro et al., Kobunshi, Konkakai, Tokyo,Japan, (1992), 7 7. J. M. Oliveira, M. H. Femandes, R. N. Correia, Biomaterials, 16, (1995), 849
Bioceramics, Volume 10 Edited by L, Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TH E HEALIN G OF SEGMENTA L BON E DEFECTS , INDUCE D BY BIORESORBABL E CALCIU M PHOSPHAT E CEMEN T COMBINE D WIT H rhBMP-2 : USIN G AS PAST E K.Ohura, CHamanishi, S.Tanaka, and N.Matsuda* Department of Orthopaedic Surgery, Kinki University School of Medicine, Ohno-Higashi, Osaka-Sayama, Osaka 589, Japan Taihei Chemical Industrial Co., Ltd., Higashi-Koraibashi, Chuo-ku, Osaka 540, J^an ABSTRAC T Bioresorbable calcium phosphate cement cylinders (0 2.5 x 5 mm) were impregnated with 6.28 or L26 |xg of recombinant human bone morphogenetic protein-2 (rhBMP-2) and implanted into segmental defects of rat femora after coating with the same cement. The high dose group yielded much bone formation around the cement and all the defects had the radiographic evidence of union at 3 weeks. The low dose group and the cement alone group united 60 and 20 % at 9 weeks, respectively. The mechanical analysis at 9 weeks showed that the defects of the high dose group failed 92 % torque compared to the contralateral control. KEYWORD S : bioresOTbable cement, calcium phosphate cement, pseudoarthrosis model recombinant human bone morphogenetic protein-2 INTRODUCTIO N We have reported that the defects of rat femora were healed completely by the implantation of bioresorbable calcium phosphate cement (BCPC) cylinders ( 0 4 x 5 mm) impregnated with 6.28 M.g of ihBMP-2 and were completely healed[l]. The purpose of this study is to investigate that the BCPC used as paste hardens in situ and heals the same defects like the cylinders. MATERIAL S AND METHOD S The bioresorbable calcium phosphate cement (BCPC) with the composition of 6-TCP (6-TCP100, Taihei Chemical Industrial Co., Ltd., Osaka, Japan )(<300 ^lm):34.9, 6-TCP(300«700 Hm):16.3, MCPM(monocalcium phosphate monohydrate): 10.9, CSH(calcium phosphate hemihydrate): 8.7, distilled water: 29.2 wt% was mixed and shaped into 0 2.5 x 5 mm cylinders. After sterilization with drying sterilizer at 180 ^C for 1 h, two doses (1.26 or 6.28 |ig) of rhBMP2 was soaked into these cylinders. Forty eight adult male Sprague-Dawley rats, weighing between 330-360g, were anesthetized with intraabdominal administration of Nembutal. With a lateral s^proach to right femur, A@a predrilled high density polyethylene plate (4 x 4 x 23 mm) wasfixedalong the anterior cortex of the femur with 0 1.2 mm ttu-eaded Kirshner wires. A 5 mm segmental defect was created in the region of the middle of the shaft with use of a dental burr. The cement cylinders with rhBMP-2 were coated with the same cement and implanted into the defects before hardening. The cement without rhBMP-2 was pasted into the defects and used as control. Under sedation, serial radiographs of right femur were made at 3, 6 and 9 weeks 419
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postoperatively . At each interval, 2 rats from each group were killed for histological study and femora were excised . Undecalcifie d specimen s were prepared and 100 |im sickness slices were microradiographed . These slices were groimded into 10 |im sickness and stained with May Griinwald Giemsa. Ten rats from each group were killed for the mechanica l test at 9 weeks. After removal of plates and pins, all of the femora, both involved and contralatera l were teste d to failure in torsion.
RESULT S AND DISCUSSIO N Large bone shells were formed around the bone defect of the high doseBMP group at 3 weeks (Fig.l-a). The shell envelope d the plate so that the size of the shell was larger on the side of the plate. The cement cylinder was still covered with the coated cement and new bone was formed on the surface of the cement (Fig, 1-b). The defect was bridged by new bone but the volume of bone formed on the cement was small because large bone shells was formed outside. The cement filled into the defect as paste became hard and kept its shape at 3 weeks (Fig.2). Only small amount of new bone was formed along the margins of the osteotomy . According to the resorption of the cement , bone advanced eccentricall y in the region of the defect opposite the plate, but the defect did still not unite at 9 weeks.
Figure 1. (a) Radiograph of the high BM P group at 3 weeks, (b) high magnification of Fig 1(a) (A: bone shell)
The defect of the high dose BM P group united at 3 weeks and the plate was covered with new bone (Fig.3). The bone formed at the place of the defect became dense and solid at 9 weeks. The cement alone group and the low and high dose BM P groups united 20, 60 and 100%, respectivel y (Fig.4). The failure torque of the high dose BM P group by the mechanica l test was 92 % of that of intact controls and higher than other two groups. W e have reported that the femora implanted with BCP C cement cylinders combined with 6.28 ^ig of rhBMP- 2 achieved 100 % of the failure torque compare with intact controls[l] . BM P had been soaked into the inner wall of these cylinders. Their BM P slowly oozed out from the cylinder surface so that new bone formation was resulted mainly on their
Healing of SegmentalBone Defects: K. Ohura et al.
Figure 2. Radiographs of the cement alone group
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Figure 3. Radiographs of the high dose BM P group
Q (0
-H C CQ
0)
Figure 4. The failure torque of defect s that had radiographic evidence of union at 9 weeks.
Cemen t BMP(L) BMP(H)
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surface. Then the cement was resorbed and bone remodeling occurred. However, the cement coated on the cement cylinder solved fast and BMP was released rapidly this time. New bone was formed mainly at the interface between muscles and the blood clots formed after the operation. The bone strength to sustain the body weight was same between both cases. Therefore, the bone formed aroimd the coated cement was large but not stronger than that formed on the cement cylinder. The role of carrier is to effectively deliver BMP by delaying its diffusion and provide a favorable environment for adherence and proliferation of a responsive population of cells[2]. The ideal BMP carrier should be biocompatible and biodegradable to minimize local tissue response and be resorbed as new bone begins to form. The bioresorbable calcium phosphate cement (BCPC) was biocompatible and was resorbed fast, when combined with rhBMP-2. Bone union of segmental defects of rat femora implanted with BCPC carrier combined with BMP started faster and finished more completely than the demineralized bone matrix carrier and the poly [D,L(lactide-co-glycolide)] (PLGA) carrier combmed with BMP [3, 4]. The BCPC which is simple and easy to deal with is thought to be one of good BMP carriers. SUMMAR Y The bioresorbable calcium phosphate cement combined with 6.28 |ig of rhBMP-2 was implanted into segmental defects of rat femora as paste. All of defects united within 3 weeks and the bone strength became nearly same as that of intact controls at 9 weeks. REFERENCES l.Ohura.K., Hamanishi.C, Tanaka.S., and Matsuda.N., Bioceramics (Proc. 9th Int’l. Sym.Ceram.Med.), 1996,9,247-250. 2.Cole.B.J., Yasko.A.W., Lane.J.M., Tomin.E., Peterson.M., Ron.E., Turek.T., and Wang.E.A., J. Bone Min. Metab., 1993, 8, 8244. 3.Yasko.A.W, Lane.J.M., Fellinger.E.J., Wozney.J.M., and Wang.E.A., J. Bone Joint Surg., 1992. 74A, 659-670. 4.Lee. S.C, Shea.M., Battle.M.A., Kozitza.K., Ron.E., Turek.T., Schaub.R.G., and Hays.W.C, J.Biomed. Mater. Res., 1994,28, 1149-1156.
DENTAL AND E.N.T. APPLICATIONS
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
IMPLAN T PLACEMEN T ENHANCE D BY A NEW BIOACTIV E MATERLV L E. Schepers, L. Barbier Biomat, Department of Prosthetic Dentistry, Catholic University of Leuven, Capucijnenvoer 7, 3000 Leuven, Belgium ABSTRAC T The purpose of this study is to determine the efficacy of bioactive glass particles of narrow size range (300-355 ^im, Biogran^) in the treatment of bone defects prior to implant placement. At both sides of the mandible of 6 beagle dogs partial edentulous areas were created with removal of the intra-alveolar septa to obtain large defects, instantly filled at one side with bioactive glass particles. The other side was left empty as a control side. After a healing period of 4 months, 3 oral implants were placed in the glass-treated area and 3 in the control zone. Half of the implants were left subgingivally for 3 months and than resected. The other half of the implants were loaded withfixedprostheses for 7 weeks and than sacrificed. A qualitative and quantitative histological analysis revealed substantially more bone tissue and a higher and longer lasting remodelling activity around implants placed in glass-treated areas INTRODUCTIO N Bone defects around oral implants can jeopardize the successfiil osseointegration of the implant. Bone defects are often related to inadequate alveolar bone at the time of implant placement, e.g., in situations where dehiscence or fenestration defects arise, or where implants are placed in not congruent extraction sockets, etc. Bone regeneration in these defects, using bone grafts or bone substitutes, would improve the long-term prognosis of the implant^ ^. Bioactive glass is a bioactive ceramic that can be used as a particulate material ^. Previous animal experiments revealed a superior response of bioactive glass particles of narrow size range (300-355 ^mi, Biogran ) compared to hydroxylapatite granules (Calcititefi and Interpore-200fi) ^"’*. More osteoconductive bone growth starting from the wall of the defects was seen around the bioactive glass particles than around the HA particles. In addition, trabecular bone growth was observed in the center of the defect. These bone trabeculae were associated with the bioactive glass particles, which exhibited an osteophilic nature, while mostly fibrous tissue separated the bone tissue from the hydroxylapatite particles. It was clearly demonstrated that the bioactive glass particles showed an internal erosion via small cracks. In these protective pouches new bone tissue was observed, which was not connected to any external bone tissue. These islands of newly formed bonefimctionedas nuclei for enhanced repair. The purpose of this animal study was to determine the efiicacy of bioactive glass particles of narrow size range (300-355 ^m) in the treatment of bone defects prior to implant placement. MATEIOAL S AND METHOD S All premolars and the first molar in both sides of the mandible of 6 beagle dogs were extracted. The intra-alveolar septa were removed to obtain large defects, instantly filled at one side with bioactive glass particles. The other side was left empty as a control side. After a healing period of 4 months, 3 IMZfi implants with a length of 10 mm and a diameter of 3.3 mm, were placed in the glass-treated (test) area and 3 in the untreated (control) zone. The implants in the first group were left subgingivally for a period of 3 months; the implants in the 425
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second group were additionally loaded with fixed prostheses for 7 weeks before sacrifice. Undecalcified serial sections, in vestibulo-lingual and mesio-distal direction of each implant, were ground and polished to a thickness of 50 to 60 ^im. The sequential administration of different fluorochrome labels allowed a quantitative analysis of non-remodelled and remodelled bone tissue at the mesial, distal, vestibular and lingual interface and at a distance of 3 mm from the mesial and distal interface. One-tailed paired Student’s t-tests were performed to determine the significance of differences between data. The level of significance was determined as P<0.05. A qualitative light microscopic analysis was accomplished by staining some sections with a combination of Stevenel’s blue and Von Gieson’s picro-fuchsin. RESULTS AND DISCUSSION Surgical and labeling treatments were uneventful. Oral implant placement in the glasstreated area was possible because these granules transformed fully and their hardness was similar to that of the surrounding bone tissue. The vascularisation of the glass-treated implant bed was more pronounced than in the control side. The primary stability of all implants was excellent.
m Remodelled bone tissue 0 Non-remodelle4 bone tissue
test
control
Interface
test
control
Distant
Figure 1. Mean amounts of non-remodelled and remodelled bone tissue at the interface and at a distance of 3 nun of the interface around implants placed in bioactive glass-treated (test) areas and in untreated (control) areas. The histological sections demonstrated an intimate contact of all implants with the surrounding bone tissue after the subgingival healing period of 3 months. Implants placed in control regions showed on average 33.1 % (SD=23.85) of direct bone contact with their surface (Fig. 1). Implants placed in areas treated with bioactive glass granules demonstrated about 20 % more bone tissue in contact with the implant surface: a mean of 51.3 % (SD=28.17). At a distance of 3 mm from the implant surface, a similar tendency was observed as at the interface: the untreated sides of the mandible showed on average 23.8 % (SD=19.24) of bone tissue, while the glass-treated side contained a mean of 55.4 % (SD=29.47) of bone tissue. All the differences, in amounts of non-remodelled and remodelled bone tissue or their sum, between implants placed in treated and untreated areas were statistically significant, despite the large standard variations and the small number of animals used in this experiment. There was statistically significantly more bone tissue at the interface than at a distance in the control
Implant PlacementEnhanced by a New Bioactive Material: E. Schepers and L. Barbier
All
regions (P = 0.045) . In the test zones, there was no significant differenc e betwee n the amount of bone tissue at the implant surface and at a distance (P = 0.188) .
mRemodelle d bone tissue HNon-remodelled | bone tissue
test control
Interface
test control
Distant
Figure 2. Mean amounts of non-remodelle d and remodelle d bone tissue at the interface and at a distance of 3 mm of the interface around implants placed in test areas and in control areas. All implants showed an intimate contact with the surrounding bone tissue after the subgingival healing period of 3 months and the functionally loaded period of 7 weeks. The implants in the control region demonstrate d on average 42.0 % (SD= 22.13 ) of direct interfacial bone contact (Figure 2). Implants in the test region showed on average 57.1 % (SD = 29.15) of interfacial bone contact, approximatel y 15 % more than in the control region. This differenc e was statisticall y significant. At a distance of 3 mm from the implant surface, approximatel y 25 % statisticall y significantly more bone tissue was found in the test sites than in the control sites: a mean of 61.6 % (SD = 27.95 ) in the test areas and a mean of 34.7 % (SD = 26.23 ) in the control areas. However, nor the implants in the test sites, nor the implants in the control sites, demonstrate d statistically significantly more or less bone tissue at a distance than at the interface . The amount of bone tissue in remodellin g around implants placed in glass-treate d areas was statisticall y significantly higher than in control areas at the interface at week 5, 8 and 11, and at a distance only in the last two time intervals. 8 and 11 weeks (Figure 3). In the test and control area, the remodellin g at the interface was more important than at a distance. Most intense remodellin g processe s were activate d within the first two to four weeks after implant placement . After the first month of healing, the remodellin g strongly decrease d to reach, after 3 months, a minimal level in the control region and a slightly higher level in the test areas. The remodellin g activity in the loaded period was less important than during the subgingival healing period. Only in the first month after placing the prosthesis , significant remodelling took place. This remodellin g was not specificall y higher in the test regions than in the control regions and tended to a normal turnover at the end of the experiment . Where implants were placed in glass-treate d areas, a considerabl e osteoconductiv e growth of bone tissue around the glass particles was observed . Around these implants, the amount of bone tissue increase d towards the implant surface, while the number and size of the glass particles decreased . Large numbers of glass granules were seen at a distance, most of them showing internal bone formation in the eroded internal protectiv e pouches, which appeared to be separate d from the externa l bone tissue. No direct contact of glass particles with the implant surface was detected .
428 Bioceramies Volume10
% §
I
25 T
Glass-treated sites 20 +
Untreated sites 15 + 10 +
S
^
5m
Om
Unloaded pericxi
Loadedperiod
Figure 3. Amount of remodelled bone tissue at the interface (IF) and at a distance of 3 nun (D) of implants placed in bioactive glass-treated areas and in untreated areas during the 3 month subgingival healing period and the 7 week functional loading period. Placement of the implants and loading resulted in a much higher remodelling activity at the interface than at a distance. As a result, granules in contact with the implant surface were totally replaced by bone tissue. In the inunediate vicinity of the implant surface, the glass granules were strongly reduced in number and size by the active remodelling processes during healing and loading of the implant. On the long term, the glass granules close to the implant surface will be replaced probably entirely by a dense bone structure, ensuring the integration of the implant. At a distance of more than 2 to 3 mm from the implant surface, the glass granules were larger in number and size. Granules in the more remote regions will be less subjected to a replacing remodelling activity than granules closer to the implant surface. Conclusions The use of bioactive glass granules of narrow size in extraction sockets prior to implant placement improved the osseointegration. After 3 months of healing subgingivally and 7 weeks of occlusal loading, implants placed in glass-treated areas showed more bone tissue at the interface and at a distance of 3 mm from the interface than implants placed in untreated areas. Most important bone remodelling around the implants in the test and the control areas occurred during the subgingival healing period. Once loaded, a new but less intense remodelling peak was observed within the first month of loading. Bioactive glass particles near the interface of the implants were completely replaced by bone tissue due to the high remodelling activity in this area. Reference s 1. Gotfredsen K., Warrer. K, Hjorting-Hansen E., Karring T. Clin Oral Impl Res 1991; 2, 172-178. 2. Block M.S., Kent J.N. Dent Clin North Am 1992; 36, 27-37. 3. Schepers E., De Clercq M., Ducheyne P., Kempeneers R. J Oral Rehab 1991; 18, 439-452. 4. Schepers E., Ducheyne P. J Oral Rehab 1997 (accepted).
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France. October 1997) '1997 Elsevier Science Ltd
BEHAVIOUR OF BIOACTIVE GLASS (S53P4) IN HUMAN FRONTAL SINUS OBLIT› ERATION Kalle Aitasalo, MD, DDS\ Jouko Suonp^a, M D \ Matti Peltola, M D \ Antti Yli-Urpo, DDS^ ^Department of Otorhinolaryngology - Head and Neck Surgery, Turku University Central Hospi› tal, FIN-20520 Turku, Finland, ^Institute of Dentistry, University of Turku, FIN-20520 Turku, Finland
ABSTRACT A prospective study was carried out on bioactive glass (BG) as an obliteration material in a series of osteoplastic frontal sinus operations on 22 patients. Preoperative examinations and postoperative follow-ups included plain films, computed tomography (CT), magnetic resonance imaging (MRI) and laboratory monitoring. The follow-up visits were scheduled at 3, 6, 12 months and 2, 3 and 4 years. No complications related to the obliteration material were ob› served. All wounds healed well without signs of infection. The repair of defects filled with BG granules was closely monitored with repeated postoperative digital ROI selection of CT tomo› graphy image areas. Plain films and CT revealed no resorptive changes at the BG-bone inter› faces. Contrast enhanced MRI indicated ingro\\1:h of fibrovascular tissue in spaces between the granules. ROI selection showed minor changes during the first year. Thereafter, the BG-bone interface remained stable. Two sinuses were explored one year postoperatively, and the retrieved specimens were evaluated histologically. The spaces between the BG granules were filled by fibrous tissue and in some areas by immature bone. The granules were partly incorporated with bone and with fibrous tissue. No inflammatory reactions, lymphocytes or granulocytes, were observed. All laboratory tests were within the normal range during follow-up. No infections were seen and no reoperations were necessary in our (22) obliteration cases. INTRODUCTION A bioactive material elicits a specific biological response at the interface between the material and tissue, resulting in the formation of a bond between them [1]. Bioactive glass (BG) is a glass-ceramic material. It has proven to be a biocompatible, osteoconductive and safe mate› rial [2, 3]. In spite of extensive clinical use, only a few long-term studies on it have been pub› lished to date. In a study on rabbits, spongy bone defects were filled with BG and long-term tissue reactions were observed using histological, histomorphometric, SEM and EDXA methods [4]. In our previous study, massive amounts of BG were used as obliteration material in osteo› plastic sinus surgery with good clinical results [5]. Patients with chronic frontal sinusitis have been regularly monitored by means of clinical examinations, CT-scans, laboratory tests, and in some cases biopsy samples have been obtained. These methods do not show the exact conditions 429
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in the obliterated area and BG granules. The conditions particularly for new bone formation in the cortical bone cavity are probably different from those in spongy bone cavities obliterated with BG during orthopedic surgery. In addition, information is needed about the long-term effects of large BG masses for evaluation of BG as an obliteration material in frontal sinuses. In the present study, a long-term model was developed for simulation of obliterated frontal sinuses. The long-term stability of BG in connection with human frontal sinus operation was also studied. MATERIA L AND METHOD S In 1991-1996, we used BG S53P4 as obliteration material in osteoplastic frontal sinus operations on 22 patients (mean age 43.6 years). All patients were suffering from chronic suppu› rative frontal sinusitis. The operative technique has been described previously (5). In this tech› nique it is important to strip all diseased mucous membrane from the sinus and from its recesses. The upper orifice of the nasofrontal duct is also obliterated with lyophilized dura or Lyoplantfi. BG granules of 630-800 ^m, and 800-1000 ^m, weighing 15-25 g have been used to obliterate the frontal sinuses. The follow-up visits were scheduled at 3, 6, 12 months as well as 2, 3 and 4 years. CT scans were taken during all these visits but MRI was only carried out at 3 and 6 months. The mean follow-up time was 2.8 years. Hematology and serum CRP, asparate transferase (SGOT) and creatinine were determined during follow-up visits. The density of the occluded sinuses was assessed in nine control regions of each CTimage by means of Region of Interest (ROI) selection using the CT PACE TM Scanner System (Figure I). The system is based determination of digital absorption coefficients of each pixel in the selected region which are then converted to Hounsfield units (H units). For valid comparison of these units of each of the nine 3-mm-wide regions, special care was taken to equalize the head position at each scanning session. All parameters of CT scanning (slice thickness, position of granule mass, radiation energy and time) were equalized. Obliterated cavity conditions were simulated in vitro using 10 BG S53P4 masses, each weighing 25 g, in simulated body fluid (SBF) in an incubator. Dissolution of Si and P was de› tected at week 1 and montly up to 6 months. The SBF solution was replaced monthly and its pH was 7.3–0.05. Five of the masses consisted of granules of 630-800 ^m and the other five con› sisted of granules of 800-1000 ^m. BG can be tested in SBF for various aspects such as bioactivity, durability, quality as› surance etc. [6], The SBF contains inorganic ions in concentrations close to those occurring in blood plasma buffered with triscintric acid. It allows not only resorption of the BG but also pre› cipitation of a calcium phosphate layer. Thus, it was possible to assess at least indirectly what occurs in BG obliterated areas and how stable BG is in the long term. On the other hand, the effect of dissolution and resorption of various clinical components on radiological ROI selection could be assessed to a certain extent.. RESULT S BG in vitrostudy.BG components were measured during incubation in SBF at I week to 6 months. With both granule sizes, the dissolution of silicon (Si) was highest during weeks I to 2 and clearly decreased between months 1 and 2. The dissolution of phosphate (P) was high at week 1 and clearly decreased between months 1 and 2 during the period when a calcium phos› phate layer formed on the BG granules. Afterwards, a slight increase occurred between months 3
Behaviour of Bioactive Glass (S53P4) in Human Frontal Sinus Oblite-Ration:K. Aitasalo et al.
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and 6. Si dissolution was high when pH was high (from week 2 weeks to month 2). Then it de› creased to a stable level. The weight of the tested BG granules with both granule sizes was 25 g. The mass of BG granules was almost the same as that used in clinical applications. The total weight loss of Si with granules of 630-800 ^m over 6 months was 733 mg and with granules of 800-1000 ^m 523 mg. The weight loss was about 3% in small granules (630-800 ^m) and 2.1% in large granules (800-1000 ^m). The loss of P over 6 months in granules of 630-800 ^m was 96 mg (0.4 wt %) and in granules of 800-1000 ^im it was 77.6 mg (0.3 wt %). These results seem to support our previous impression based on clinical studies that BG is a safe and stable material even when large masses (20-30 g) are implanted BG infrontal sinus obliteration.At follow-ups, the CT-scans of all 22 patients showed no marks of inflammatory processes or mucous membrane regeneration in the frontal sinuses (Figure 1). The BG particles accomplished total obliteration in all sinuses with only minor visi› ble changes of general softening during the follow-up period. During the first follow-up year, the decrease in average density was 14.8% (H units), mainly as a result of sinusal drying or fibrosis. Afterwards, the decrease was 7.8% to 2.2% over the remaining follow-up period. In five patients; two men and three women (mean age 45.6 yrs) four-year results are now available (Figure 2). Over the four years, the total decrease in H-units was 22.8% as a result of resorption of glass (2.2 to 24% per year) and the effect of tissue drying in the operation area. Histopathological exami› nation after 1-year follow-up of bone and glass samples showed fibrosis around the BG granules and a clear reaction layer. No new bone formation was seen between the glass particles. No in› flammatory changes or foreign body reactions were seen. On MRI, fibrotic tissue was seen to increase towards the centre of the frontal sinus at 6 months but no other changes were found. To date, no complications have occurred and none of the patients have been reoperated on.
Figure 1. Bioactive glass (BG) in the frontal Figure 2. Density of bioactive glass (BG) in frontal sinus two years after osteoplastic operation, sinuses after osteoplastic operation, assessed using ROI selection, results converted to Hounsfield units. Means of measurements from five sinuses.
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DISCUSSIO N The osteoplastic operation of the frontal sinus is limited to patients with chronic suppu› ration and failure to respond to other treatments. The patients of this study had a severe long› standing infection. The high failure rates in earlier studies and our own experiences have made us look for new, stable obliteration materials. According to the present study, bioactive glass seems to obliterate the frontal sinus better than fat and other occlusion materials. The stability of BG was also proven by two biopsy samples showing no inflammation or foreign body reactions in the sinus. Repeated CT scans of frontal sinuses were carried out during the follow-up period showed only minor visible changes with no signs of regeneration of mucous membrane or reinfections in the cavity or in the osseous laminas of the sinuses. The in vitro results of massive BG implantation also seem to support our previous impression that BG is safe and stable material both in vitro and in clinical situations. The material enhances tissue healing and may also promote new bone formation in bone cavities. On MRl, only early fibrous tissue ingrowth in sinuses was seen, in agreement with histological findings. Clinically, the BG gran› ules were well accepted without any adverse affects. All defects consolidated over 6 months. All laboratory tests were within the normal range during follow-up. The results indicate that BG alone can be used a filler in human frontal sinus operations. REFERENCES 1. 2. 3. 4. 5. 6.
Hench, L.L., Andersson, O.H. In: An Introduction to Bioceramics, Reed Healthcare Communications, Singapore 1993, 41-62. Wilson, J., Pigott, G.H., Schoen, F.J., Hench, L.L. Toxicology and biocompactibility of bioglasses. Journal of Biomedical Materials Research Volume 15, 1981, 805-817. Nagasae, M., Abe, Y., Chikira, M., Udagawa, U. Toxicity of silica containing calcium phospate demonstrated in mice. Biomaterials, Volume 13, 1992, 172-175. Heikkila, J.T., Salonen, H-R., Yli-Urpo, A., Aho, A.J. Bioceramics Volume 9, 1996, 123-126. Aitasalo, K., Peltola, M., Suonpaa, J., Yli-Urpo, A., Andersson, O.H., Varpula, M., Happonen, R-P. Bioceramics Volume 7, 1994, 409-414. Heikkila, J.T., Mattila, K.T., Andersson, O.H., Knuuti, J., Yli-Urpo, A., Aho, A.J. Bioceramis Volume 8, 1995, 35-40.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
ALLCERAMI C DENTA L BRIDGE S BY TH E DIREC T CERAMI C MACfflNIN G PROCES S (DCM ) F. Filser’, P. Kocher’, H. Luthy’, P. Scharer’ and L. Gauckler’ Swiss Federal Institute of Technology Zurich, Department of Materials, Institute for Nonmetallic Materials, Zurich, Switzerland University of Zurich, Center for Dental and Oral Medicine, Clinic of Fixed and Removable Prosthodontics and Dental Materials Science, Zurich, Switzerland KEYWORD S zirconia, green machining, rapid prototyping, prosthesis, all-ceramic dental bridge ABSTRAC T A new process using direct machining of ceramics (DCM) allows the manufacturing of allceramic dental bridges. Presintered preforms of ceramics, e.g. Tetragonal Zirconia Polycrystals (TZP) are machined in short time and then sintered to final shape. The preforms show homogenous shrinkage during final sintering. No further machining is needed after sintering to fiill density to achieve high strength and high accuracy. We report the process steps, strength of veneered ceramics, accuracy, and the development of an all-ceramic dental bridge by the DCM process. INTRODUCTIO N Today, dental bridge prostheses are mostly metal-ceramic composite structures: a metallic fi*amework for load bearing, coated with porcelain for aesthetic appearance. In recent years, the use of many metals in the oral cavity has been disputed due to their biological incompatibility risk. Therefore, an all-ceramic bridge would be of great clinical value. In the past, ceramics were widely applied in dentistry [1]. All of these ceramics show low bend strength and toughness implying design restrictions, non-reliability and complicated multistep manufacturing procedures for bridges. These drawbacks can be eliminated by TZP stabilised with 3 mol% yttria. It exhibits in its dense sintered state higher strength and higher toughness than other dental ceramics as classified by Liithy [2]. Despite of its good properties zirconia has not yet been used for all-ceramic dental bridges due to difficulties in shaping by usual techniques like casting and sintering, hot pressing or grinding in its dense form. To overcome these shaping difficulties a new process was developed to fabricate all-ceramic dental bridges by direct machining of presintered greenbodies. Furthermore a new veneer porcelain adequate for the coating of zirconia firameworks was developed. In the following this new manufacturing process will be described and also the possibilities of thermal and mechanical adjustment of the porcelain veneer to the TZP ceramic. APPROAC H Our approach is shown in Fig. 1. Homogenous ceramic green bodies are fabricated and are slightly sintered to porous blanks. The outer shape of the resin model of the bridge is digitized. These bridge contours are enlarged for compensating the sintering shrinkage and transformed to machine data used to control the milling machine. Then the enlarged bridge fi-amework is machined out of the porous ceramic blank. This enlarged presintered firamework is sintered to full density maintaining the required accuracy. In the last operation the all-ceramic dental bridge is veneered with porcelain to meet aesthetics in colour and translucency. 433
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C^amic powdef)
Q^resi n model^^
presintered l framework f
sintered framework
veneering
machinin g
Figure 1. Process steps of the Direct Ceramic Machining (DCM) process. MATEWAL S AND METHOD S Process and Accuracy TZP powder (Tosoh, Tokyo) is used in the DCC process [3] for fabricating the green body and presintered at 1075 C for 1 hour. The presintered bodies are cut and ground to cuboid shape. The density of each blank is determined. A resin pattern of the bridge framework is fabricated by the usual procedures, digitised, and enlarged. Machining data are generated for rough and fine milling for each side and the enlargedfi-ameworkis machined out of a blank. Theframeworkis then sintered to full density at 1500 C for 2 hours. Veneering is performed using the porcelain with thermal expansion coefficient (TEC) adjusted to zirconia. To evaluate shrinkage homogeneity during sintering we used an array pattern in a test blank with 3 mm x 3 mm quadrafic elevations of 10 mm height according to Fig. 2. Prior and after sintering the array pattern the geometry of the array pattern was measured and compared. Straightness of the edges in the sintered state was analysed using a projector. Veneer porcelain Bend strength of zirconia bars coated with veneer porcelain as function of the TEC mismatch of the veneer to that of zirconia was tested. Two types of zirconia bars (20 mm length, 6.25 nmi width) of 1 nmi and 3 mm thickness respectively were ground from a dense sintered zirconia plate, polished and annealed at 1450 C for 30 min. Two glasses of different TEC were mixed for the TEC adjustment of the veneer and used for coating of the zirconia bars. Firing was performed at 960 C for 2 min and twotimesat 920 C for 1 min each. The total thickness of the test bars was 4 nmi after polishing. All specimens were tested in three-point bending mode with the porcelain on the tensile side. For each porcelain 5 specimens were tested and the modulus of rupture (MOR) was calculated using the load when the first crack occured.
All-Ceramic Dental Bridges by the Direct Ceramic Machining Process(DCMj: F. Filser et al.
1^ presintered specimen
28 mm
435
H
^i^t’:\\VA’-\\\y//>.’,ht)>)>yk’//.’,’y
X shrinkage factor prediction sintered specimen
deviation Figure 3. All-ceramic bridge with dense sintered zirconia framework and porcelain veneer. For determining bend strength of bilayered structures as function of the veneer materials’ strength we used zirconia bars of 20 nmi length, 4 mm width and 2 mm thickness. Different veneer materials with MORs ranging from 64 to 270 MPa and elastic moduli of 9 and 70 GPa were applied. Four experimental porcelains (A to D) of similar elastic modulus of 70 GPa and TEC adjusted to zirconia were used. In addition we used one filled polymer (E) with a low elastic modulus of 9 GPa. The total thickness of the specimens was 4 mm. All specimens were tested with the veneer face in tensile mode (first crack).
Figure 2. Test grid specimen to determine shrinkage homogeneity.
RESULT S Process and accuracy In order to evaluate the shrinkage homogeneity we processed dental bridges and test grid specimens as shown in Fig. 2. The linear shrinkage was 22.80 %. The edges of the elevations of the test grid specimen were found not to bulge and remained straight. The average deviation between the predicted dimensions and the sintered array pattern is 19 |im – 12 jim standard deviation for 96 measuring points on the test grid specimen. The dimensional accuracy during 250
,2 CM
O
CI
u
*^porcelain " ^zirconia *-’-"
*^
J
Figure 4. Adjustment of the TEC a of the porcelain to the zirconia (azirconia = 110 [10"^ K*^])-
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240 T CM
^
200 1
B
D
160 Zr02 veneer veneer porcelain (E = 70GPa) veneer = polymer (E = 9GPa)
120 80
50
100
150 200 250 MO R of veneers [MPa]
300
Figure 5. Bend strength of bilayered structures as a function of the MOR of the veneers. sintering is better than 0.1 % of the total specimen length. This is more than sufficient for dental bridge application. Fig. 3 shows prototypes of an all-ceramic dental bridge with a dense sintered zirconia framework. The dental bridge is placed on the die which demonstrates the high accuracy achieved even with complex free form surface shapes (see Fig. 3 bottom). Veneer porcelain Bilayered structures with a large difference in TEC exhibit a large difference in bend strength for the analysed veneer porcelain thicknesses (see Fig. 4). Specimens with thin porcelain veneers always exhibit higher strength than thick ones. Bilayers with veneer porcelain of similar or slightly lower TEC than zircionia showed comparable bend strengths. For veneering a framework we recommend a porcelain with a slightly smaller TEC than TZP designated by the shaded area in Fig. 4. Bend strength of bilayered structures as function of the MOR of different veneer materials (A, B, C, and D being porcelains with the same TEC of 11.0 x 10"^ K^) and a polymer (E) are shown in Fig. 5. Higher MORs of veneer materials lead to high strength of the bilayers. For lower elastic moduli (i.e. from 70 GPa for B to 9 GPa for E) and similar MOR of both veneers an increasing bend strength of the bilayers is observed. SUMMAR Y A new process "Direct Machining of Ceramics" (DCM) was developed for all-ceramic dental bridges. It comprises ceramic preforms that are machined in a presintered state and subsequently sintered to full density. DCM is feasible for all ceramics. Examples for zirconia (TZP) preforms and complex shaped dental bridges were demonstrated in detail. Final machining in dense sintered state is obsolete due to high accuracy. A veneering porcelain was developed matching zirconia in the TEC behaviour. High strength can be expected for zirconia composites using high strength veneering glass or polymer veneering with low elastic modulus. Further steps comprise the characterisation of all-ceramic dental bridges in laboratory and in vivo. ACKNOWLEDGEMEN T This research is supported by the Swiss Priority Program SPP for Materials Research. REFERENCE S 1. Anusavice, K.J., Ceramic Transactions,1995, 48, 101-124. 2. Liithy, H., in CAD/CIM in AestheticDentistry.,W.H. Mormann, Editor, 1996, Quintessence, Chicago, 229-239. 3. Graule, T.J., Baader, F.H., and Gauckler, L.J., cfi/Ber.DKG, 1994, 71, 4, 317-323.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
GRINDIN G OF ZIRCONIA-TZ P IN DENTISTR Y -CAD/CAM-TECHNOLOG
Y FOR
THE MANUFACTURIN G OF FIXE D DENTURES It Luthardt^ W, Rieger^, R, Musil^ ^Friedrich-Schiller-Universitat, Jena, Department of Prosthetic Dentistry, Bachstrasse 18,07740 Jena (Germany), ^Metoxit AG, 8240 Thayngen (Switzerland) ABSTRAC T Due to the rising attention attracted by the biocompatibihty of the material used titan and allceramics-systems are gaining importance. The all-ceramics-systems being currently available are made of glass- and Al203-ceramics. The processing of Zirconia, based on CAD/CAMtechniques, is made possible by the Precident DCS-Systemfi (Ginbach Dental GmbH, Pforzheim, Germany). Regarding its biocompatibihty and material-testing features (flexural strength of 900 MPa compared to 450 MPa of AI2O3) yttrium-oxide-partially-stabilized Zirconia (PSZ) is of great interest as a material for the manufacturing of crowns and bridges [1][2][3][4]. The possibilities of grinding of ceramics such as tetragonal Zirconia-Polycristals (Zirconia-TZP) (Metoxit AG, Thayngen, Switzerland) by the dental technician shall be examined by the presented investigation. KEYWORD S dentistry, CAD/CAM, grinding of Zirconia-TZP, all-ceramics-systems, crowns and bridges INTRODUCTIO N In dental prosthetics distinct alloys are used without any or in combination with ceramics or composites for the manufacturing of fixed prosthesis and implant superstructures. All-ceramics fixed dentures gained interest and importance especially because of their advantages concerning the biocompatibihty of the materials used. Here especially the absence of base alloy-constituents which are necessary to build a bond oxide layer in ceramics fused to metal fixed dentures are of consequence. On the one hand the long-term success of the all-ceramics crowns and bridges is determined by the mechanical features of the frameworks. The resistance against the growing of cracks is especially indicated by the fracture-toughness of the material. So an increased Weibull’s Modulus is more important for the long-term success of all-ceramics crowns and bridges than a high bending strength. On the other hand good mechanical properties are necessary to extend the application of all-ceramics fixed dentures to the posterior teeth or being regarded from a different point of view to increase the safety of the restoration. 437
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Figure 1: The Zirkonia-TZP-blanks
Figure 2: Sample die of tooth 16
It is not possible to manufacture fixed dentures out of Zirconia-TZP by direct sintering on qualified dies because of the material properties of Zirconia-TZP. Therefore the manufacturing of Zirconia-TZP has to be made by grinding. Grinding- and milling- technology in dentistry require the use of CAD/CAM-Systems in almost all cases. MATERIAL S AND METHOD S The CAD/CAM-System used was the Precident DCS-Systemfi[5]. The material used for this investigation was yttriimi-oxide-partially-stabilized Zirconia [6] manufactured by Metoxit AG [Thayngen, Switzerland]. Figure 1 shows the Zirconia-blanks. Figure2 pictures one of the 6 sample dies of tooth 16 which was used in this investigation. The first step is the definition and recording of the preparation line, the margin of the preparation of the tooth. The digitising of the surface-data of the prepared tooth takes place inside this area. The recording of additional data is possible if the CAD/CAM-system used is able to process this information. Additional data are the contour or the contact surface of the tooth in front of and behind the prepared one. This data may be recorded by optical [7] [8] [9] or mechanical [10] systems. The data-set of the three-dimensional tooth surface will be processed alter the entire recording. The manufacturing-files are processed on the basis of the three-dimensional surface-data together with design parameters, as clearance and thickness of the material. The processed machining files are used to control the CNC-machine. In case of the Precident DCS-Systemfi frameworks out of titan are milled or Zirconia-TZP and InCeram are ground. After the CAD/CAM-manufacturing a dental technician has to adapt theframeworkto the model die [11]. Figure 3 presents one framework of Zirconia. For the investigation of the frameworks the machined blanks were separated, the parting planes were polished and SEM pictures were taken. Later, to avoid side eflFects by sample preparation modified blanks were used. For this Metoxit AG [Thayngen, Switzerland],
Grinding of Zirconia - TZP in Dentistry: R. Luthardt et al.
439
>.<w|i4-_
Figure 3: Framework 16 of Zirconia
Figure 4: Modified blank after grinding the inner surface of a framework
produced half-cylinders . The margins were bevelled , the parting planes polished and two halfcylinders adhesivel y bonded. This new cylinder was poured in epoxy-resin . Both the blanks and the modified blanks were machined with various grinding parameters using the Precident DCS-Systemfi. Figure 4 shows one half of a modified blank after grinding the inner surface of aframeworkand before SEM-examination. TH E SEM-EXAMINATIO N Microcracks reduce the long-term success of all-ceramics crowns and bridges. The more and the longer the cracks the worse are the mechanica l properties . So CAD/CAM-manufacturing of ceramics has to avoid the developmen t of cracks [12]. The SEM-examination should proof the presence of grinding-relate d rim-damage and microcracks.
Figure 5: Locations where the SEM-pictures were taken
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The points within the blanks where the SEM-pictures were taken are depicte d by figure 5. These are the inner top surface, the side of theframeworkand the area of the preparation-line . Under the aspect of manufacturing engineerin g the worst grinding relations are to be found on top of the inner surface. The side allows answers to the surface condition to be reached. The area of the preparation-lin e has to stand the highest force while chewing but has the smallest thickness of the framework. CONCLUSIO N The results of the examinatio n show the possibility to manufacture fitting fixeddentures using the Precident DCS-Systemfi, These results are confirmed by the SEM-pictures of the ground surface. To use CAD/CAM-Technology instead of the casting of alloys it is necessar y to shorten the manufacturing time of the ceramics. If the manufacturing time of fixed dentures can be decreased , the expansion of the application of all-ceramics-system s will be ob\ious.
REFERENCES: [I]
[2] [3] [4] [5] [6] [7] [8] [9] [10] [II ] [12]
Christel, P. et al: Mechanical propertie s and short-term in-vivo evalation of yttriumoxide partially-stabilize d Zirconia. J Biomed Mater Res 1989, 23, 45-61. HolscKW.,KapperU H. F.: Festigkeitsprufim g von Vollkeramische n Einzelzahnersat z fur den Front- und Seitenzahnbereic h Dtsch zahn^rztl Z 1992, 47, 621-623 . RiegerW.:Aluminium- und Zirkonoxidkeramik in der Medizin ID R 1993, 2/93, 116-120 . Stevens, K: Zirconia and Zirconia Ceramics, An introductio n to Zirconia Magnesium Elektron Publikation No. 113, Second edition Litho 2000, Twickenham/UK 1986. Luthardt, R Musil,R,: Das Precident DCS-Systemfi -CAD/CAM-gefertigter Zahnersatz aus Titan und Zirkonoxid Phillip 11996,13,217-225 . Maier,K R,: Leitfaden technisch e Keramik; WeikstoflBkunde II , Keramik, Selbstverlag Institut fiir keramische Komponenten im Maschinenbau, Aachen 199L Benz,C , Schwarz, P.: Wie genau ist der optische Cerec-Abdruck? Dtsch zahnarztl Z 1991, 46, 632-634 . Bose,M , Ott,K. K R.: Wissenschaftlich e Erkenntnisse tiber das Cerec-Syste m Dtsch zahnarzU Z 1994,49 , 671-673 . Zel,vanderJ, M ; C AD/C AM-Restaurationen in der Okklusion ZahnarzU Welt 1994, 103, 420-425 . Bieniek, K. W.:Computer- und andere automatisiert e Systeme zur Erstellung von Zahnrestauratione n ZM K 1994, 5/94, 6-14. Holmes, J.R.,et al: Considerations in measurement s of marginal fit J Prosthet Dent 1989,62 , 405408 . Marx,R,: Modeme Keramische WeiiestofiF e ftir ^sthetisch e Restauratione nVerstariamg und Bruchzahigkeit Dtsch zahnarzU Z 1993, 48, 229-236 .
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris. France, October 1997) '1997 Elsevier Science Ltd
ZmCONI A IMPLANT S WIT H A PLASM A SPRAYE D SiOj-HA BIOACTIV E COATIN G Alexandre Pedra, Patrick Sharrock Rimimplant, Reims and Laboratoire de Chimie Inorganique, Universite Toulouse III, France. KEYWORD S Zirconia, dental implants, coating, Hydroxylapatite-silica . ABSTRAC T Yttria stabilized zirconium oxide is used for its high mechanical strength to manufacture the core of a dental implant. Bioreactivity relies on a plasma-sprayed undercoat containing zirconia and silica and an exterior silica-hydroxylapatite surface layer. Histological and clinical evaluations demonstrate excellent osteointegration. Load bearing implants are functional. Good initial clinical results are important for long term serviceability and success rates. INTRODUCTIO N Tetragonal zirconia, ZrOz, stabilized with Y2O3 is attracting much attention as a high quality technical ceramic (1-4). The enhanced mechanical properties of zirconia ceramics characterized by a flexural strength near 900 MPa and a toughness of lOMPa makes it an ideal biomaterial for the manufacture of dental implants (5-6). Crystalline calcium phosphates, hydroxylapatite (HA) or tricalcium phosphate (TCP) are too brittle to be considered for the manufacture of load-bearing implants (7). This has led to the development of HA plasma-sprayed coatings (8). Failure to obtain an adequately adherent HA bioactive coating on the ceramic substrate led us to examine various undercoats on zirconia in order to obtain a functional gradient. The Prosiap (projected silica and apatite) layer was successful and resulted in the manufacture of the first zirconia based dental implant with bioactive coating to be marketed in France in 1993. We report here the physico-chemical characteristics of the RIM implant and some features of the clinical results obtained with prostheses in use on the ceramic implants. MATERIAL S AND METHOD S RIM implants were manufactured by machining PSZ green bocfy cylinders followed by sintering (ZrOz with 3mol % Y2O3 from Ceramiques Techniques Desmarquest) (9). The implants have a cylindro-conical shape with three lengths (13, 16, 19 mm.) and widths (2.8, 3.3, 3.8 mm.). The transgingival collars are polished, while the endosseous root portions are coated with the Prosiap deposit. X-ray patterns were studied directly on coated samples. A few grains were detached from the coating for FTIR spectroscopy. Seven adult rabbits were used for the in vivo experiments. The endosseous parts of the implants were cut off and sterilized by gamma radiation. Two implants were impacted in each tibial epiphysis area in various locations in order to study the biological response in medullar, spongy and cancellous areas. Rabbits were sacrificed at 3, 6, and 12 weeks post-operatively and non decalcified sections obtained by infiltrating, cutting grinding techniques. RIM implants were placed in patients requiring dental root replacement using classical progressive drills cooled with internal serum irrigation, and a final drill with an adapted conical shape and calibrated depth markers. 441
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3000
Figure 1. X-ray pattern of Prosiap before(a) and following immersion (b).
Figure 2.1.R. spectra of Prosiap before(a) and following immersion(b).
RESULTS AND DISCUSSION The Prosiap coating consists of a Zr02-Si02 plasma-sprayed deposit of 20 um thickness covered by an HA-SiOi layer 30 um thick.The resulting zirconium containing silica melt provides an adherent undercoat for the subsequent HA outer layer. Mechanical tests performed on cylindrical test-pieces yielded lower tensile strength values for HA (12.7 MPa) than for the undercoat (61.8MPa). In the zircon-zirconia system the phases are compatible in all proportions at temperatures up to 1675 C (10), which may explain the strength of the bonding layer. The HA+undercoat (56 Mpa) and silica-HA+undercoat (57 Mpa) both have good tensile strengths which may be explained by glass fiision and chemical bonding by calcium silicate formation at interphases.The X-ray diffraction spectrum of the undercoat shows large and strong reflections at Cu Kalpha (theta) values of 15 and 25 , indicative of tetragonal zirconia in the substrate with typical d values of 2.968 and 1.821 A (Angstroms). Smaller peaks at d = 3.162 and 2.840A reveal the presence of monoclinic zirconia in the bonding layer. The plasma-sprayed silica-HA coating shows an X-ray pattern illustrated in figure la. The coating transforms into a more typical HA following immersion in Ringer’s solution during one week, as shown in figure lb. There is a small but perceptible shift in the position of the main HA peak from d == 2.8136A in the deposit to d = 2.8122A in the recrystallized deposit, accompanied by a doubling of the intensity. The starting HA powder has d = 2.8117A, shifting to d = 2.8141 A after mixing withfilmedsilica and calcining at 900 C. The infrared spectra confirm the amorphous nature of the deposit and its tendency to recrystallize when immersed into Ringer’s solution. Figure 2 illustrates the corresponding changes. The absorption peak at 1650cm’^ in figure 2b is related to carbonates present in the recrystallized coating. The Si-O-Si vibrations are absent from the 2b spectrum suggesting the HA originates in fact from new crystals deposited on the coating. All implants demonstrated good osteointegration at all time periods. By three weeks a periosseous formation was present surrounding the implants and sealing the cortical implant access zones. A thin 50 um wide, layered bone apposition followed the implant surfaces in the medullar areas. Figure 3 illustrates this new bone -bioactive coating interface- No spalling of the coating was observed. At six weeks, neoformed bone was attached to the implant in cortical situation and followed the surface geometry of the implant into spongy areas. Signs of bone remodelling near the implants were visible, with no deposition offibrousencapsulation tissue. At
Grinding of Zirconia - TZP in Dentistry: R. Luthardt et al.
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Figure 3. X-ray view of two RIM implants in rabbit epiphyseal bone at six weeks. twelve weeks the implants are very firmly bound, with endosteal bone proliferating along the ends of the implants protruding in the tibiae. The bone marrow had a normal appearance and we could not detect any significant dissolution or resorption of the coating. The clinical uses of RIM implants have been described and are multiple (11). Partially and totally edentulous patients have been treated successfully. Extraction-immediate implantation has been performed as well as implantation with simultaneous bone augmentation with HA products. These procedures are usually hazardous and difficult with bare metal implants, but seen
Figure 4. Retroalveolar view of 4 RIM implants placed in the maxilla. Note the natural bone height 26 weeks post-implantation and the absence of visiblefibrousinterposition.
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to be well tolerated with Prosiap coated zirconia implants. The example illustrated here concerns a woman with seven implants placed simultaneously in January 1995. Following implant integration, temporary crowns were fitted which were replaced by permanent, individual, sealed ceramic crowns on standard inlay cores. An adjacent revitalized adjacent tooth was also treated. None of the implants were connected. The patient was checked regularly. The panoramic view in November 1995 confirmed the absence of periapical infection or peri-implant inflammation. In March 1997 retroalveolar X-rays revealed good bone behaviour 26 months post-implantation, without any interposition of fibrous tissue. Ceramics are well accepted physiologically and psychologically as well. Overall patient satisfaction is excellent CONCLUSIO N Stabilized zirconia ceramics provide adequate mechanical strength for load bearing dental implants. HA-silica bioactive plasma-sprayed deposits with a proper ceramic bonding imdercoat stimulate early bony integration. The one step gingival procedure allows simultaneous soft tissue healing and yields very good esthetic results and tight gingival seals effectively preventing peri-implant inflanunation and encouraging patient compliance with proper hygiene. We feel technical ceramics combined with a bioactive coating provide an ideal material for challenging dental reconstructions.
RERERENCES 1. H. Liu, Q. Xue, investigation of the crystallization of Zr02 (Y2O3 3 mole%) nanopowders, J. Mat. Res (1996), 917-921. 2. W. Burger, H.G.Richter, C.Piconi, R.Vatteroni, A.Cittadini, M.Boccalari, New Y-TZP powders for medical grade zirconia, J. Mater. Sci.,Mat. in Med., 8, (1997), 113-118. 3. C.L Curtis, D.T.Gawne, M.Priestnall, The processing and electrical properties of plasmasprayed yttria-zirconia, J. Mat. Sci, 29, (1994), 3102-3106. 4. M.Mattioli-Belmonte, P.Mengucci, N.Specchia, G.Golbi, S.Dubini, L.Simonelli, F.Greco, G.Majni,G.Giagini,C.Rizzoli, An experimental study in X-ray spectroscopy of the zirconium (Ca-PSZ) bone interace. Microanalytical evaluation of the osteogenic response, J. Mat. Sci., Mat. in Med, 8, (1997), 85-90. 5. A.H.Heuer, L.W.Hobbs, Science and technology of Zirconia, Advances in ceramics vol 3 American Ceramic Society, Westerville, Ohio,(1981). 6. M.S.Zolotar, C.A.C.Zavaglia, Fracture toughness and microstructure degradation of Y-TZP in aqueous physiological environnement, J. Mat. Sci. Mat in Med., 7, (1996),367-369. 7. J.Li, L.Hermansson, R.Soremark, High strength biofimctional Zirconia: mechanical properties and static fatigue behaviour of zirconia apatite composites. , J. Mat. Sci. Mat in Med., 4, (1993),50-54. 8. Y.Kawamoto, Y.Yokogawa, M.Toriyama ,S.Kawam\u^, T.Suzuki, Coating of Beta-tricalcium phosphate on Yttria-partially stabilized zirconia using magnesium metaphosphate as an interlayer. J. Ceramic. Soc. Jap. Int-ed, 99, (1991), 19-22 9. B.Cales, High reliable zirconia ceramics for orthopaedics. Fifth world biomaterials congress, Toronto, 1996, 177. 10. T.Itoh, Zircon ceramics prepared from hydrous zirconia and amorphous silica. J. Mat. Sci. Let, 13, (1991), 1661-1663. ll.J.Pedra, P.Sharrock, Un nouvel implant dentaire bioactif en zircone: I’implant RIM, Actualites en Biomateriaux, 4, (1995), 405-410 .
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the10thInternational Symposium on Ceramics in Medicine, Paris,France,October1997) '1997 Elsevier Science Ltd
EFFEC T OF TIM E AND TEMPERATUR E ON THE PRODUCTIO N OF POROU S ELECTROLYTI C HYDROXYAPATIT E COATING S N. Asaoka*+, S. Best+ and W. Bonfield+ *Central Research Institute, Mitsubishi Materials Corporation, 1-297 Kitabukuro-cho, Omiya, Saitama, 330, Japan +IRC in Biomedical Materials, Queen Mary and Westfield College, Mile End Road, London El 4NS, U.K.
ABSTRAC T Porous hydroxyapatite coatings were produced on pure titanium substrates by the electrolysis of clear solutions containing calcium and phosphate ions and complexing agents at a constant DC voltage. The effect of the electrolytic conditions, i.e. the temperature, duration and applied voltage, on the properties of the coatings were investigated. The coatings were analysed using Xray diffraction (XRD) and scanning electron microscopy (SEM). No deposit was obtained at applied voltages higher than 5V. The coatings prepared at temperatures higher than 60T were identified, using XRD, as low crystallinity hydroxyapatite and SEM revealed that they had a porous honeycomb-like structure. The pore size varied from 0.5 to 2 microns depending on the temperature at which electrolysis was performed. KEYWORD S hydroxyapatite, titanium, coating, electrolysis, complexing agent INTRODUCTIO N Hydroxyapatite is one of the most widely researched bioactive ceramics. However, its clinical use has been restricted to non-major load bearing applications due to its relatively inferior mechanical properties. In order to take advantage of the bioactivity of the material and to overcome the disadvantages of mechanical performance, a number of methods have been investigated of applying hydroxyapatite coatings to metallic substrates. One of the most popular techniques currently in use is plasma spraying [1] and the vast majority of commercially manufactured hydroxyapatite coated implants for dental and orthopaedic applications are produced using this route. However, there are some disadvantages with the use of thermal spray techniques such as phase decomposition, poor coating reproducibility and debonding at the coating/substrate interface. A variety of alternative techniques, are therefore under investigation including dipping and sintering [2], sputtering [3], HIP [4], electrochemical deposition [5] and electrophoresis [6], although, currently there do not appear to be any commercially viable techniques torivalplasma spraying in terms of mechanical strength or productivity. A process for the production of uniform monetite coatings on titanium substrates by electrophoresis in clear solutions has been reported previously [7]. In this study, results are reported from a new process to prepare porous hydroxyapatite coatings directly from solution. 447
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Figure 1 Experimental apparatus for electrolysis MATERIAL S AND METHOD S Reagent grade orthophosphoric acid, calcium hydroxide and a complexing agent were dissolved in distilled water to prepare clear electrolytic solutions. Ammonium hydroxide solutions were utilised to adjust pH. The electrode set was composed of a commercial purity titanium plate (50x20x1 nmi^) as a cathode and a pure platinum plate (50x20x0.125mm^) as an anode. The titanium and platinum substrates were degreased using acetone and were then placed parallel with each other at a separation of 10mm. The electrodes were dipped into a vessel containing \QOcvc? of the clear electrolyte at a depth of 15mm for electrolysis (Fig.l). When the solution reached a predetermined temperature (from 20 to 80T) a constant DC voltage up to 10 volts was applied between the electrodes. As electrolysis commenced, the alkaline solution was dripped into the vessel and the electrolyte was slowly stirred to be homogenised. The current and pH of the electrolyte were monitored during electrolysis. After electrolysis for up to 60 minutes, the electrode sets were gently rinsed in distilled water to remove excess electrolyte and dried at room temperature for several days. Then the dried coated substrates were weighed to calculate the net weight of the coating layers. Phase identification was made using X-ray diffraction (Siemens D50(X)) with a tube voltage 40kV, 40mA and scanning rate 3.2deg(2theta)/min.. The structure of the coatings were observed using a scanning electron microscope (JEOL JSM-6300F) with accelerating voltage 5keV. RESULT S AND DISCUSSIO N The initial pH of the orthophosphoric acid, calcium hydroxide and a complexing agent mixture was less than 4 and on addition of the ammonium hydroxide solution, the pH increased up to a final value of pH 8 with precipitation commencing at above pH 5. There was a marked effect of voltage on coating formation: at 60"C, the net weight of the cathode increased with increasing voltage up to 5V, but at voltages above this, no deposition occurred (Fig.2). Generally, the coating weight increased with increasing temperature (except for the reaction performed at SOT at 5V). The fact that no coatings were obtained on the cathode with voltages higher than 5 volts is possibly due to the vigorous bubbling resulting from the electrolysis of water molecules which was observed on the surface of the both electrodes. In this case there appeared to be more
Effects of Time and Temperatureon Production of Porous ElectrolyticHA Coatings: N. Asaoka et al.
449
electric power consumed in the electrolysi s than in depositio n and as a result, coating formation may have been prevented . The bubbling was associate d with a correspondin g increase in electric current. As shown in Figure 3, the current increase d rapidly with increasing applied voltage and temperature . This may have affecte d the physical adhesion of the coating layers, and the sudden decrease of the coating weight at 80"C and 5 volts in Figure 2 may also be explaine d by the partial decompositio n of the layer by the bubbling. Electrolysis temperatur e also affecte d the formation and appearance of the coatings: the colour of the coatings prepared at 60X or above, was pure white while those prepared at lower temperature s were light grey. However, extendin g the duration of electrolysi s did not necessaril y result in an increase in coating weight. Figure 4 indicates that coating weight increase d from 10 to 20 minutes, while it decrease d at 30 minutes. The results from X-ray diffraction indicated that the coating layers deposite d at a temperatur e of 60*C or higher were composed of low crystallinity hydroxyapatit e while those deposite d at temperature s lower than 60"C were mainly composed of brushite (CaHP04-2H20). 10 E S "S
^
I
_ 8 \-
o
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.
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2
3 Voltage
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Figure 3 Effect of temperatur e and applied voltage on the electric current
(min.)
Figure 4 Effect of electrolysi s duration on the coating weight at 60T, 5V
Figure 5 Porous hydroxyapatit e coating surface (SEM image), prepared at 60**C and 5V for 10 minutes. )
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At the beginning of precipitation, the pH of the solutions was about 5. Therefore the phase composition of the initial precipitates and the coatings would have been brushite which is more stable than hydroxyapatite at lower pHs. As the pH continued to increase throughout the reaction, hydroxyapatite would have become more stable. However, an activation energy would have been needed to be supplied to convert crystallised brushite to hydroxyapatite and it is likely that, only at relatively high temperatures, would this activation energy have been sufficient. SEM observation revealed that the hydroxyapatite coating layers were uniform with a porous honeycomb-like structure. The pore sizes appeared to be more strongly influenced by the reaction temperature than reaction duration and, as indicated in Figure 5, the average pore size for a reaction performed at SOT (3V) was approximately 2 microns where for a coating deposited at 6 0 T the average pore size was approximately 0.5 microns. The porous hydroxyapatite coatings observed in this study are in contrast to monetite coatings, prepared previously [7] which were reported to consist of densely packed uniform plate-like particles. This structural difference may reflect the morphological preference in crystal growth under the applied electric field. The hydroxyapatite coatings prepared at SOT had a larger pore size than those at 60X. This difference seemed to be dependent on the growth rate of the crystals with precipitation temperature. CONCLUSIO N Hydroxyapatite coatings can be prepared using an electrolytic process in solution containing calcium and phosphate ions with a complexing agent. At applied voltages of 5 volts or lower a fine white deposit resulted on the cathodic titanium plate. Coatings prepared at 6 0 T or higher consisted of low crystallinity hydroxyapatite, while brushite coatings were obtained at lower temperatures. The hydroxyapatite coatings had a very porous structure with pores of 0.5 to 2 microns in diameter. ACKNOWLEDGMENT S The support of the EPSRC for the IRC in Biomedical Materials and of Mitsubishi Materials Corporation for one of the authors (N.A.) are gratefully acknowledged. REFERENCE S 1. de Groot, K., Geesink, R., Klein, C.P.A.T.and Serekian, P., J. Biomed. Mater. Res., 1987, 21, 1375-1381 2. Yankee, S.J., Pletka, B. J., Luckey, H. A. and Johnson, W. A., Therm. Spray Res. Appl., 1991,433-438 3. Ong, J. L., Lucas, L. C , Lacefield, W. R. and Rigney,E. D., Biomaterials, 1992, 13 ,249254 4. Heroe, H., Wie, H., Joergensen, R. B. and Ruyter, I. E., J. Biomed. Mater, Res.,1994, 28, 343-348 5. Shirkhanzadeh, M., J. Mater. Sci. Lett., 1991,10, 1415-1417 6. Ducheyne, P., Radin., S., Heughebaert, M and Heughebaert, J. C , Biomaterials, 1990,11, 244-254 7. Asaoka, N. Best, S. M. and Bonfield, W. In Bioceramics Volume 9, Elsevier Science Ltd., Oxford, 1996, 289-292
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CALCIU M PHOSPHAT E FORMATIO N ON CHEMICALL Y TREATE D VACUU M PLASM A SPRAYE D TITANIU M COATING S S.-W. Ha^ K.-L. Eckert^ H. Gnrner^, E. Wintermantel^ ^ Chair of Biocompatible Materials Science and Engineering, Department of Materials, ETH Zurich, Wagistr. 23, CH-8952 Schlieren ^ Medicoat AG, Gewerbe Nord, CH-5506 Magenwil ABSTRAC T Carbon fiber reinforced polyetheretherketone (CF-PEEK) substrates were coated with titanium (Ti) by vacuum plasma spraying (VPS) and chemically treated in lOM sodium hydroxide (NaOH) or 30% hydrogen peroxide (H2O2) solution. After chemical treatment, the specimens were immersed in simulated body fluid (SBF) containing ions in concentrations similar to those of human blood plasma. Scanning electron microscopy (SEM), energy dispersive X-ray analysis (EDX) and diffuse reflectance fourier transformed infi*ared spectroscopy (DRIFT) were used to analyse the chemically treated VPS-Ti surface and the calcium phosphate layer formed during immersion in SBF. It was observed that a carbonate containing calcium phosphate layer was formed on the NaOH treated VPS-Ti surface during immersion in SBF, whereas no calcium phosphate precipitation occurred on the untreated and H2O2 treated surfaces. It is therefore concluded that vacuum plasma spraying with Ti and subsequent chemical modification in 10 M NaOH solution at 60 C for 2 hours is a suitable method for the preparation of bioactive coatings for bone ongrowth on CF-PEEK. KEYWORD S Vacuum plasma spraying, carbon fiber reinforced PEEK, titanium, calcium phosphate coatings INTRODUCTIO N Vacuum plasma spraying on CF-PEEK is currently being established to produce Ti and hydroxyapatite coatings in order to make this thermoplastic composite suitable as implant material [1]. It is assumed that such coatings will enable long-term fixation of this material in bone tissue. In the present study, two different methods, NaOH treatment or H2O2 treatment on VPS-Ti surfaces were peformed to provide nucleation sites for calcium phosphate formation in a simulated body fluid. The characteristics of the modified surface and the induction of in vitrobiological apatite for› mation were investigated. MATERIAL S AND METHOD S CF-PEEK (Ensinger GmbH, Germany) disks with a diameter of 10 mm and a height of 7 mm were sandblasted with alumina, cleaned with ethanol and deionized water and dried in a vacuum oven at 200 C for at least 7 days. VPS was performed with fine Ti powder (d5o = 25 |Lim) and coarse Ti powder (d5o = 120 |im). Chemical treatment of the VPS coatings was performed in 30% H2O2 at room temperature and in 10 M NaOH solution at 60 C. Soaking time was 2 hours for both treat› ments. After chemical treatment the specimens were washed in distilled water and immersed in sim451
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ulated body fluid (SBF) which was prepared as described in [2]. Immersion was carried out in 25 ml of SBF during 1, 4, 10 and 24 days in a laboratory shaker (Infors AG) rotating with 80 rpm at 1>TC and pH 7.4. After immersion the specimens were gently rinsed in distilled water and dried in air. Scanning electron microscopy (SEM) and energy dispersive X-ray (EDX) analysis was performed to characterize morphology and chemical composition of the coatings. The specimens were coated with platinum in a sputter coater before SEM and EDX analysis. Chemical changes of the substrates after immersion in SBF were examined by diffuse reflectance fourier transformed infrared (DRIFT) analysis (Perkin Elmer System 2000). Infrared spectra were obtamed in a wavenumber range of 4000-400 cm 1 RESULT S AN D DISCUSSIO N In fig. 1, the SEM micrographs of untreated, H2O2 treated and NaOH treated VPS-Ti surfaces are shown. After NaOH treatment the formation of a fibrous, needle-like structure was observed. No morphological changes were detected on the untreated and H2O2 treated surfaces. EDX spectra showed that Na was incorporated into the VPS-Ti coating after NaOH treatment (fig. 3 left) and it is assumed that the newly formed structures as observed with SEM contain sodium. Besides Na and Ti no additional elements were found on the NaOH treated samples. No changes in chemical composi› tion were observed on both, untreated and H2O2 treated surfaces. After immersion in SBF, the as-received and the chemically treated VPS coatings were analysed with EDX and SEM. Fig. 3 shows the EDX spectra of the untreated and treated VPS-Ti coatings. In both, untreated and H2O2 treated substrates, only the Ti peak coming from the VPS-Ti coating and a little amount of platinum (Pt) from the sputter coating was detected. SEM analysis revealed that no calcium phosphate layer was formed on the as-received and H2O2 treated VPS-Ti surfaces after 24 days of immersion in SBF (fig. 2). On the NaOH treated coating, formation of calcium phosphate was observed aheady after one day of immersion in SBF. After 24 days the VPS-Ti surface was completely covered with calcium phosphate precipitates (fig. 2 and 3). Fig. 4 shows the FTIR spec› tra of the NaOH treated VPS-Ti surfaces, which were immersed in SBF for 1, 4 and 24 days, show› ing the continous growth of a carbonate containing calcium phosphate layer, which is regarded to be similar to the chemical composition of biological apatite m the natural bone [3-5].
^"^ untreate d
Figure 1
H2O2 treate d
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SEM micrographs of untreated (left), H2O2 treated (middle) and NaOH treated (right) VPS Ti coatings showing the significant morphological change after treatment in lOM NaOH at 60 C for 2 hours compared to the untreated and H2O2 treated surfaces.
Calcium PhosphateFormationon ChemicallyTreatedVPS TitaniumCoatings: S.- W. Ha et al. 453
The results of the present study demonstrated that H2O2 treatment of VPS-Ti surfaces does not induce the formation of calcium phosphate. After NaOH treatment a small Na peak appeared in the EDX spectra. This small amount of Na could result from the formation of sodium titanate in accor› dance to the observations reported in [6]. In that work the formation of Na2Ti50ii on various pol› ished Ti alloy surfaces after alkaline treatment in NaOH was shown. During immersion in SBF calcium phosphate formation could have occurred due to a ion-exchange process, since sodium titanate shows a high ion-exchange capacity and all its Na^ ions are exchangeable [7].
untreated
Figure 2
NaOH treated
H2O2 treated
SEM micrographs of untreated (left), H2O2 treated (middle) and NaOH treated (right) VPS Ti coatings after immersion in SBF at 37 C for 24 days. On the NaOH treated surfaces, the formation of a calcium phosphate layer was observed. In contrast, no precipiation of calcium phosphate occurred on untreated and on H2O2 treated surfaces. before immersio n in SBF
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EDX spectra of untreated, H2O2 treated and NaOH treated VPS-Ti coatings before (left) and after (right) immersion in SBF. On NaOH treated surfaces, Na was detected, while no change of the chemical composition was observed on H2O2 treated specimens. After immersion in SBF, Ca was detected on NaOH treated substrates indicating the formation of a calcium phosphate layer.
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FTIR spectra of VPS-Ti coatings on carbon fiber reinforced PEEK after NaOH treatment and subsequent immersion in SBF at 37 C, obtained in the diffuse reflectance mode. With increasing immersion time, phosphate and carbonate bands became more distinct indicating the continuous precipitation of a carbonate containing calcium phosphate.
CONCLUSIO N The present study has demonstrated that vacuum plasma spraying of Ti and subsequent chemical modification in 10 M NaOH at 60 C for 2 hours is a suitable method for the preparation of bioactive coatings on carbon fiber reinforced PEEK. Carbonate containmg calcium phosphate, which has nearly the same morphology and chemical composition as biological apatite in natural bone, has been formed on the NaOH treated VPS-Ti coatmgs durmg immersion m SBF. The VPS-Ti coatings showed a very rough topography and a high surface area. This is assumed to positively affect the chemical modification of the Ti surface by NaOH treatment and the in vitrocalcium phosphate dep› osition on the modified VPS-Ti surface.
REFERENCES [1] [2] [3] [4] [5] [6] [7]
Ha S.-W., et al., in: Bioceramics,Vol. 10, Elsevier Science Ltd., Oxford, UK, 1997, in press. Kokubo T., et al, in: Bioceramics,Vol. 4, Butterworth-Heinemann, Guildford, 1991, 113-120. Maruno S., et al, Biomaterials,12, 1991, 225-230. Brophy G.P., and Nash J.T., The americanmineralogist,53, 1968, 445-454. Heughebaert M., et al., Journal of BiomedicalMaterialsResearch,22, 1988, 254-268. Krni H.-M., et al. Journal of BiomedicalMaterialsResearch,32, 1996, 409-417. Clearfield A., et al.. Journal of solid statechemistry,73, 1988, 98-106.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PROPERTIE S OF PLASM A SPRAYE D BIOACTIV E FLUORHYDROXYAPATIT E COATING S RANZ X.l, REY C.l, ANTOLOTTI N.2, HARMAND M.F.3, MORONI A 4, ORIENTI L 4, VIOLA 0.2, BERTINI 5.2. SCRIVANI A 2 1 Laboratoire dc Physico-Chimie des Solides INPT-ENSCT. CNRS-UPRESA 5071, 38 rue des 36 ponts, 31400 Toulouse, France 2 Hametal S.p.A. Biocoatings div.. Via G. di Vittorio 51,43045 Fomovo Taro, Italy 3 LEMI, Technopole Montesqieu, 33650 Martillac, France 4 Istituti Ortopedici Rizzoli, via Di Barbiano 1/10,40136 Bologna, Italy ABSTRAC T In order to improve both the resistance to in vivo biodegradation and the osteoconductivity of plasma sprayed coatings, we developed a bi-layer coating composed of a stable fluorhydroxyapatite plasma sprayed coating covered by a poorly crystalline carbonate s^atite layer analogous to bone mineral. In vitrotests showed that despite a very slight cytotoxicity, the presence of the carbonate layer improves osteoblast proliferation and colonisation. The bi-layer coating bone ingrowth and bone attachment behaviour is similar to the more soluble hydioxyapatite plasma sprayed layers. KEYWORD S Plasma sprayed coating,fluorhydrohyapatite,solubility, surface treament, osteoconductivity. INTRODUCTIO N Plasma sprayed hydroxyapatite (Caio(P04)6(OH)2: HA) coatings on titanium alloys are largely used on orthopaedic prostheses owing to their ability to enhance bone formation and to establish strong bone/implant bonding [1]. Unfortunately, HA decomposes during the plasma spraying process into more soluble compounds [2,3] which can induce release of debris and detachment of the coating [4]. On the other hand, the osteoconductivity of Ca-P coatings has been shown to be related to their partial solubility [5]. It has been reported in a previous paper that fluorhydroxyapatite coatings (Caio(P04)6FOH: FHA) decompose less than HA during plasma spraying and lead to more adherent coatings [6]. In this report we compare the dissolution of HA and FHA plasma sprayed coatings at constant pH, and we describe a surface treatment allowing the formation of a carbonateapatite layer on the FHA coating. The main chemical-physical characterisatics of this layer (FTER, X-ray diffraction, SEM, specific surface area) are given and its biological activity determined by in vitroand in vivo studies. MATERIAL S AND METHOD S All the plasma spray depositions of stoichiometric HA and FHA were performed by Biocoatings. The kinetics of dissolution was measured at constant pH and temperature (pH=4 and T=37X) in lactic acid solution. The dissolution rate of HA and FHA coatings were determined by the quantity of acid required to maintain constant pH. After 20 hours the samples were dried and weighed then, they were sonicated to release loose particles and weighed again; this test gave an indication of the disaggregation of the coating. 455
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Poorly crystalline carbonate apatite was precipitated on the surface of FHA from supersaturated solutions containing Ca(NC)3)2, (NH4)2HP04 and NaHCOs [7]. The in vitrotests WCTC performed according to European and French standards (ISO 10993-3; NF S 91-145; NF S 91-142). The in vivo implantation into sheep and the biomechanical behaviour and histological obsCTvations were carried out according to published procedures [8]. RESULT S AND DISCUSSIO N Substrate characterisatio n FHA has been shown to be one of the least soluble Ca-P apatites [9]. The physical-chemical characterisation of the two coatings is reported in table 1 [6]. Constant pH and temperature dissolution curves are represented in figure 1. For coatings with similar specific surface areas (around 0.55 m^.g-^), the data indicate a slower dissolution rate for FHA than for HA.
Samples HA FHA
Table 1. Characteristics of HA and FHA coatings Phase composition (%) Physical charactoistics Oxy. Amo. P-TCP TTCP Ca/P Density Porosity Crystal. (atomic) (m^g-l) (%) (%) 65 20 1.0 1.1 1,667 3.08 15.3 83 07 16 0.4 1.8 1,666 3.13 9.8 81
Although it is difficult to relate the kinetics of dissolution of the coatings to a precise cause, the presence of foreign phases more soluble than ^atite, especially amorphous calcium phosphate, appears as the main parameter (table 1) [2, 3]. Moreover, fluoridation of apatites is known to improve their resistance to acid dissolution [11]. However, dissolution is not the only mechanism involved in the biodegradation of Ca-P coated implants; disaggregation of the coatings, which results in the release of fine particles of apatite into the solution, also occurs. AftCT dissolution and sonication, we measured the loss of weight (Fig. 2). Disaggregation was seen to be lower for FHA than for HA. The parameters are essentially the same as for dissolution. A predominant role of the amorphous Ca-P phase can also be suggested. This phase is among the most soluble of the Ca-P compounds and behaves like a binder between crystalline particles. Furthermore, hydrolysis of CaO into Ca(0H)2 favours crack formation inducing the disaggregation of the Ca-P layer [12]. FHA gives less soluble and more resistant Ca-P plasma spayed coatings. If we consider that bone remodelling in the vicinity of the implant involves acid dissolution, it may be expected that FHA coatings will last longer than HA coatings. On the otherhand, the osteoconduction seems to be related to the solubility of the material [5]. The biological activity of FHA might then be limited although FHA seems to favour cell proliferation and to increase bone density [14]. Surface treatmen t The aim of is this treatment was to form a poorly crystalline apatite at the surface of the coating to improve the biological activity. After treatment, SEM analysis showed the presence of small sph^cal particles from 0.5 to 2 ^m in diameter covering the entire surface of the substrate (Fig. 3). The chemical composition determined by EDS indicated that calcium and phosphate were associated with sodium and carbon (carbonate). Furth^more the specific surface area increased by about 12 % after treatment due to the presence of small crystals on the substrate. The X-ray diffraction patton of the precipitate famed by homogeneous nucleation, in the same conditions, revealed a poorly crystalline apatitic phase (Fig. 4a). The average dimension of crytals det^mined from band broadening of the 002 and 310 X-ray lines, [15] (150 A in length, 45 A in width) confirmed the very small size and the expected high reactivity of the precipitate.
Propertiesof Plasma Sprayed BioactiveFluorhydroxyapatiteCoatings: X. Ranz et al.
Figure 1. Dissolution curves
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Figure 2. Weight loss after sonication
Figure 4b represent s the Fourier Transformed infrared spectrum of the precipitate . The poor resolution of the apatitic bands can be related to the poor crystallinity. The bands at 1488, 1424 and 870 cm’^ are due to carbonate ions in labile environment s indicating the high activity of the lay^ [16]. The cartxniate content (about 20% weight) appeared high forcarbonate apatite of type B and corresponde d more probably to sodium-containin g carbonate apatite [18]: Ca7Na2(P04)3(C03)3(OH). In order to evaluate the bioactivity of FHAS , several invitrotests were performed. FHA S coatings appeared to be weakly cyto-toxi c resulting in a partial attachmen t of the human osteoblas t cells : after 6 hours of incubation: 60 % cells attached for HA . and only 35 % for FHA S (Fig. 5). However, FHA S favoured cell proliferatio n during the first 6 days of incubation (Fig. 6). After 27 days, the cell density on FHA S and HA was respectivel y 85 % and 66 % (100 % for the polystyrene test sample). This observatio n could be partly assigned to the effect of fluoride on cell growth [14]. SEM analysis showed that FHA S significantly favoured cell density and dispersion and cyt(9lasmic extensio n in comparison with HA . FHA S coated pins were implanted and compared with HA coated samples in unloaded conditions. Osteointegratio n was assesse d by measuring the percentag e of contact and the percentag e of bone growth. For FHA S and HA bone growth was respectively : 74 % and 80 % after 15 days; 99 % and 98 % after 9 months. This activity could be related to the solubility of the carbonate layer for FHA S and the presence of soluble phase in the HA plasma coating.The quality of the bone/implan t interface was measured by comparing the intrusion torque and extractio n torque of screws 6 weeks after of implantation in loaded conditions [8], both FHA S and HA enhance d the bone/coatin g integration . The histological observation s performed in unloaded and loaded conditions confirmed the osteoconductio n and osteointegratio n properties of the two coatings however a faster degradation of the HA coating was noted. The FHA S coating appeared to be as;
310
Figure 3. SEM inmge of plasma spayed coating after surface treatmen t
Figure 4. a) X-ray diffraction pattern of (rf the pecipitate ; b) FTIR of Uie precipitat e
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Cells /m2 r NOOOO 5000 0
Days l
Figure 5. Cell attachment study
3
6
9
15 2 1 2 7
Figure 6. Cell proliferation study
efficient as the HA coating regarding bone integration but more resistant to successive bone remodelling. CONCLUSIO N Bioactive coatings have to comply with two opposite properties: long-term stability which ensures a good boneAmplant interface and biological activity to favour bone repair. The solution of a bi-layer coating composed of a plasma-sprayed deposit of FHA with good bonding to the metal substrate and a surface lay^ of carbonate ^atite ^>pears to be a good alternative although it may be improved. ACKWOLEDGMEN T The present work was a part of a Brite EuRam programme supported by the European Commission.
REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 11. 12. 14. 15. 16.
Geesink, R., De Groot, K. and Klein, C.P.A.T. /. Bone JL Surg,1988, 70B, 17-22 Klein, C.P.A.T, De Blieck-Hogervorst, J.M.A., Wolke, J.G.C. and De Groot, K. Advances in Biomaterials1990,9.277-282 Van Blitterswijk, C.A., Bowell, Y.P., Rach, J.S., Leenders, H. and Bakker, D. BoneBiomat.Interface1991,295-307 Frayssinet, P., Hardy, D., Hanker, J.S. and Giammara, B.L. Cells and Materials1995,2, 125-138 LeGeros, R.Z., Orly, I., Gregoire, M. and Daculsi, G. Bone-Biomat.Interface1991,76-88 Ranz, X., Gobbi, L., Rustichelli, F., Antolotti, N. and Rey, C. Ceram,Cells and Tissues Volume 2, Ravaglioli 1996, in prepartion Antolotti, N., Chil6, M., Casti, A., Carton, F., Rey, C. and Ranz, X. Patent PR96A000021, 1996 Moroni, A., Orienti, L., Stea, S., Visentin, M. and Nfaltarallo, C. In: BioceramicsVol.8, Pergamon, Oxford 1995,345-350 Mweno, E.C., Kresak, M. and Zahradnik, R.T. Caries Res. 1977,11,142-171 Christoffersen, M.R. and ChristoffCTsen, /. Calcif TissuesInt.1985,37,673-676 W«ig, J., Wolke, J.G.C, Zhang, X. and De Groot, K. In: BioceramicsVol. 8, Pergamon, Oxford 1995,169-175 Farley, J Jl., Wergedal, JE. and BayUnk, DJ. Science1983, 222, 330-337 Sherrer, P. Gdtt.Narch.1918, 2, 98 Vignol, C. Thesis,I.N.P. Toulouse, 1973
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedingsof the10th InternationalSymposiumon Ceramics in Medicine, Paris, France, October1997) '1997 Elsevier Science Ltd
LONGER-TER M MECHANICA L AND BIOLOGICA L EVALUATIO N OF TITANIU M ALLO Y COATE D WIT H APATIT E LAYE R W.Q. Yanl*, K. Kawanabel, T. Nakamural, T. Kokubo^ iDept. Of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, Sakyo-Ku, Kyoto, Japan ^Dept. Of Material Chemistry, Faculty of Engineering, Kyoto University, Sakyo-Ku, Kyoto, Japan
ABSTRAC T The objective of this study was the development and use of a biomimetic coating method, which deposits a thin and uniform apatite layer onto titanium (Ti) implants in simulated body fluid (SBF), for improving their bone-bonding ability. The mechanical characteristics and histology of commercially pure Ti and apatite layer-coated Ti-6A1-4V alloy were investigated in rabbit tibiae. Interface failure load was determined using a detaching test after periods of 6,10,24, and 32 weeks. The apatite layer-coated implants exhibited significantly higher failure load than the uncoated control at each time period (all p < 0.001). Histologically, the coated-implants bonded directly to bone via the uniform coating, with no intervening soft tissue. In uncoated controls, there was fibrous tissue intervening at the interface even at longer periods. SEM-EPMA demonstrated a Ca-P-rich layer on the coated implants and direct bonding at the interface. No sign of degradation and delamination of the coating was seen after periods of up to 32 weeks. Our results suggest that apatite layer coating on Ti alloys in situmay improve bone-to-implant bonding and make it as a promising coating material. INTRODUCTIO N Coating of bioactive ceramics on metallic prostheses for promoting direct bone bonding has been a subject of extensive research. The use of HA coatings prepared by plasma-spray methods have shown enhanced quality of bone apposition and excellent early clinical results. However, the integrity of the coating thickness and composition as well as adherence of the coating to metals are still practical problems[l]. There are concerns that the instability of HA-coatings and the generation of HA debris as a result of either degradation or delamination will reduce efficacy in the longer term [2]. We have previously evaluated the surface modification of Ti implants using chemical treatments and reported the formation of bonelike apatite on the treated implants in vitro and in vivo [3,4,5]. The bone-bonding strength of Ti implants was remarkedly increased via the bonelike apatite on the implant surfaces [3]. It is therefore expected that the use of biomimetic process to biologically coat apatite layer on Ti implants in situmay render metals bioactive and may also be the solution to concerns of physically processed coatings. In the present study, the bone-bonding behavior of apatite layer-coated Ti alloys produced by the biomimetic method was evaluated for a longer-term in comparison with commercially pure Ti implants using mechanical test and histological examination. * Present Address: Centre for Biomaterials, University of Toronto, 170 College Street, Toronto, Canada 459
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MATERIAL S AND METHOD S SBF, the ion concentrations of which were almost identical to those of human blood plasma, were prepared by dissolving NaCl, NaHCOs, KCl, K2HPO4, MgCl2, CaCl, and Na2S04 in ionexchanged and distilled water and buffering at pH 7.40 with Tris buffer. Rectangular plates of Ti-6A1-4V (10 x 10 x 2 mm) were used as substrates. The plates were abraded with #400 diamond paste, cleaned in an ultrasonic bath, and then treated in lOM-NaOH aqueous solution and heated at 600’’C prior to the coating process. Coating with apatite was performed using a biomimetic method. After soaking in SBF with a buffer (pH 7.40) at 36.5 C for 4 weeks, the specimens were removed from the solutions, gently rinsed with distilled water and dried at room temperature. To confirm the deposition of apatite layer on the substrate, the surface of the coated implants were examined by scanning electron microscopy (SEM) (Hitachi S-2500CX), energy-dispersive X-ray microanalysis (EPMA) (Horiba EMAX-3700) and thin film X-ray diffractometry (TF-XRD) (Rigaku, CN2651A). Untreated Ti plates were abraded with #400 diamond paste and used as control implants. All implants were sterilized in ethylene oxide gas before implantation. The plates of apatite layer-coated Ti6A14V alloy (n = 8 for each time point) and Ti control (n = 8 for each time point) were implanted into the metaphyses of mature rabbits for periods of 6, 12, 24 and 32 weeks. Under sterile surgical conditions, a longitudinal incision was made on the anteromedial aspect of the proximal metaphyses. Using a dental burr, a slightly oversized hole was made from the medial to the lateral cortex parallel to the longitudinal axis of the tibia. Then an implant was inserted into the hole by perforating the tibia from the medial to the lateral side. A coated plate was implanted in one leg of the tibia and an uncoated one was implanted contralateral leg as a paired control. At sacrifice, segments of the tibiae containing the implants were excised and subjected to a detaching test to evaluate the bonding strength of the bone-implant interface as reported previously [3,5,7]. The failure load was measured when either the implant became detached from the bone or the bone itself broke. After mechanical testing, all specimens were fixed in a phosphate-buffered formalin solution, dehydrated in serial concentrations of ethanol and embedded in polyester resin. Each large block was sectioned, perpendicular to the longitudinal axis of the tibia, with a diamond band saw (BS30(X), EXAKT, Hamburg, Germany). Sections (5(X) \\mthick) were polished with diamond paper and coated with a layer of carbon for observation using the scanning electron microscope furnished with an energy-dispersive X-ray microanalyzer (SEM-EPMA) and backscatter electron detector. The other sections were further ground to a thickness of 80 |im (Microgrinding MG-4000; EXAKT, Germany) and used for Giemsa surface staining and contact microradiography . RESULT S The SEM-EPMA analyses of the cross-sections of Ti4A16V implants showed that a uniform apatite layer (about 20 |im thick), corresponding to the Ca-P-rich layer, was formed and covered the whole implant surface (Fig. 1). No Ca-P-rich layer was found on the Ti control implants. By TF-XRD, the peaks of apatite were confirmed on the coated implants. In the detaching tests, specimen fracture usually occurred between the bone and the implant, but breakage in the bone was observed in some of the apatite-coated specimens. In uncoated controls, some specimens were separated spontaneously before the test, and hence the failure load was defined as 0 kgf. The failure loads of the coated implants were respectively 1.62–0.68
Longer-TermMechanicaland Biological Evaluationof Ti Alloy Coated With Apatite:W.Q. Yan et al. 461
kgf and 4.13–1.8 4 kgf at 6 and 12 wks, and reached to the failure load of 6.09–2.0 4 kgf at 24 wks and 6.0212.1 0 kgf at 32 wks. In the paired controls, the failure loads were only 0.0310.0 2 kgf and 0.8510. 1 Ikgf at 6 and 12 wks, and increase d to 1.9510.7 6 kgf at 24 wks and 2.0110.8 3 kgf at 32 wks. The effect of the coating treatmen t on the failure load was significant (ANOVA ; p = 0.0001) . Further statistica l testing indicated a significant differenc e in failure load betwee n the coated and uncoate d implants at each time interval (r-test; all p < 0.001) . Histology showed active bone formation on the coated implants and the bone in direct contact with the apatite coating, without any intervenin g soft tissue (Fig. 2). A high degree of bone/implan t contact was demonstrate d for the coated implants by backscatte r electro n microscopy and contact microradiography. In contrast few bone apposition, but many fibrous tissue areas, were observed for the uncoate d implants at the early stages. In the longer-ter m periods, the uniform apatite coating bonded to both bone and substrate with no sign of degradation and delamination . At 32 weeks, the apatite layer was replaced by thin lamellar bone, which remained bonded with the underlying implant. SEM-EPM A showed a uniform apatite layer with a high Ca and P level at the interface betwee n the bone and the coated implants and no Ca-P layer on the uncoate d control at all intervals.
mmm
v\
[Rl^^Qa
tJOeoou) Fig.l SEM-EPM A of the cross-sectio n of the coated implant. A Ca-P-rich layer is deposite d on the Ti alloy surface.
Fig.2. Giemsa surface staining of the coated implant (CI), showing a thin apatite layer bonded to the implant surface and bone (B), with no intervenin g tissue.
DISCUSSIO N Attainment of a more adherent , and stable, coating-implan t interface is considere d to be critical in the longer-ter m implant fixation. In this study, we coated apatite on Ti alloy using a novel biomimetic approach and obtained a thin and uniform coating easily without any complicate d technique . The surface structure measured by TF-XR D and SEM-EPM A indicated bonelike apatite formation and Ca-P-rich layer on the coated implant surface. The mechanica l and biological results demonstrate d that the coated apatite layer had a profound effect on the bone-bondin g behavior of Ti alloy implants. A s previously reported, the formation of a number of apatite nuclei on the substrate with some treatment s can be initiated in simulated body fluids. These apatite nuclei grow
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spontaneousl y once they are formed on the substrate, and eventuall y a continuous apatite layer is formed by consuming the calcium and phosphate ions in the fluid [4,5,7] . This coating process may be analogous to the process of in vivo apatite formation on bioactive materials, and it appears more bioactive than coatings from conventiona l physical methods which may change the physicochemica l propertie s of apatite due to the high heat involved in those processe s [1,2]. It is well known that chemical bonding betwee n bioactive ceramics and bone is achieve d through a bonelike apatite (Ca-P-rich) layer that formed on the bioactive material surfaces in the body. In the present study, the apatite layer-coate d implants showed direct bone bonding without any fibrous tissue interposition , compared with major areas of fibrous tissue around the uncoated implants. A significantly higher failure load was found in the coated implants than in the uncoate d control at each time period. SEM-EPM A and Giemsa surface staining analyses provided further evidence that the coated implants bonded to bone via the thin coating of bonelike layer with a high Ca and P level. These improvement s could be due to chemical bonding of the bone-implan t interface and strong adherenc e of the coating to the implant. Studies examining the effect of the apatite coating on bone-implan t bonding in the longerterm provided further insight into the nature of the coating/meta l interface . Our finding of the coated implant at the longer-ter m periods showed that the coating adhered to the underlying implant so tightly that fracture occurred within the cortical bone. Apatite layer appeared to be replaced by bony tissue during remodeling without any delamination . This biologic bony substitution might have facilitate d the stability of coating-implan t interface , leading to a longerlasting and strong bone-implan t bonding. Apatite layer-coate d implants produced by the biomimetic method proved to be an effectiv e approach in improving bone-implan t bonding. The use of this method to coat a apatite layer on Ti-based metal indicates that further studies are warranted to investigate the bone-bondin g Ti metal in a load-bearing condition as a more physiological and more durable orthopeadi c implants.
REFERENCES
LFiliaggi, M.J., CoombsJsf.A., Pilliar, R.M., JMiomedMater. Res. 1991,25:1211-1229 . 2.Bloebaum, R.D., Beeks,D., Dorr,L.D., Savory, C.G., Dupont, J.A., Hofmann,A,A., Clin. Orthop.1994,298:19-26 . MaterRes., 1997, 3.Yan, W.Q., Nakamura,T., Kobayashi,M., Kim,H.M., Kokubo,T. JJBiomed. in press. 4. Kokobo,T., Miyaji,F., Kim,H.M., Nakamura,T. J.Am.Ceram.Sic. 1996, 79:1127-1129 . 5.Yan,W.Q., Nakamura,T., Kobayashi,M., Kokubo,T., Kim, H. M., Miyaji, F., Bioceramics 1996, (9): 305-308 . 6.Kitsugi,T., Nakamura,T., Yan, W.Q., Oka,M., Goto,T., Shibuya, T., Kokubo, T., Miyaji, S., J.Biomed.Mater. Res.,1996, 32:149-156 . 7.Yan,W.Q., Kawanabe,K., Nishiguchi,S., Nakamura,T., Kokubo,T. Biomateriah (1997), in press.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
ELECTROPHORETI C COATING S OF POROU S APATIT E COMPOSIT E ONT O ALUMIN A CERAMIC S Kimihiro Yamashita, Eiko Yonehara, Jun-ichi Hamagami, Takao Umegaki Department of Industrial Chemistry, Tokyo Metropolitan University 1-1 Minami-Osawa, Hachioji, Tokyo 192-03, Japan
ABSTRAC T The coating of porous hydroxyapatite on alumina and zirconia ceramics was undertaken by the electrophoretic lamination method. The multi-layers were comprised of porous hydroxyapatite, intermediate hydroxyapatite, and adhesive layer of calcium phosphate glass. The open porosity and pore size of the surface layers were adjusted by the addition of graphite or alumina powders. In case of alumina additives, the surface layers were decomposed to tricalcium phosphate during sintering, while hydroxyapatite structure was maintained in graphite-added surfaces. K E Y W O R D S : Hydroxyapatite, Porous Apatite, Electrophoretic Coatings, Alumina Ceramics, Composite Coating, Bonelike Crystal Growth INTRODUCTIO N Electrophoretic deposition (EPD) combined with sintering is reportedly one of the practical coating methods of hydroxyapatite (Caio(P04)6(OH)2, HAp) [1-5]. We have carried out EPD coatings of apatite on the ceramics of alumina and yttria-stabilized zirconia for the biomedical use [5]. For this purpose, porous surface layers with HAp structure are desirable. The strong adhesion of the coated layers to substrates is cdso required. To achieve such complex aim, we carried out the EPD lamination of the porous surface, intermediate HAp, and adhesive layers on alumina ceramics. In the present EPD method, the adhesion layers to ceramic substrates were first deposited by the EPD of the mixed powders of HAp and calcium phosphate glasses (CaQP205, CP) with a lower melting point (1(XX)-12(X)**C) onto ceramic alumina substrates according to the previous work [4]. Then HAp intermediate layer was formed on the first layer. The porous surface layers were lastly formed by the co-deposition of (1) HAp and alumina powders or (2) HAp and graphite powders; the pores were generated during sintering. This paper mainly reports the EPD preparation of porous surfaces as well as the lamination procedure. MATERIAL S AN D METHO D Starting powders of HAp were prepared by a wet chemical method. CP powders for adhesion layers were obtained by crushing the dried agglomerates which were obtained by the reaction of CaC03 with phosphoric acid. The reactant powders had the atomic ratio of Ca/P=l/2. Because CP powders were scarcely deposited, they were co-deposited with HAp. The surface layers were formed by the co-deposition of HAp with alumina (40wt%) or graphite powders (lwt%). Commercial alumina powders with an average diameter (d)=ljim were used, and graphite powders with d=10 and 25jim were prepared by scratching, crushing and sieving from the bulk graphite. 463
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The EPD was undertaken by applying dc voltage of lOOV/cm for several min to those mixed or single powders suspended in a mixed solvent of acetylacetone and ethanol (Acac/EtOH). After drying, the multi-layered alumina substrates were sintered at ISOOT. The coated aluminas were characterized by XRD and IR analyses. In order to evaluate the effectiveness of porous surface, the coated specimens were immersed in a simulated body fluid (SBF) for a week. Scanning electron microscopy was undertaken for the observation of bone-like crystal growth at deep parts in a coated layer. RESULT S AN D DISCUSSIO N Figure 1 shows the laminated layers (A) consisting of adhesion (B), intermediate (C), and porous surface (D). The adhesive and intermediate layers were well densified by sintering. The adhesion layer, which was composed of HAp and CP, was reacted to calcium phosphate compound with a lower melting point (-1100**C). The adhesion strength was lOMPa, which was influenced by the surface roughness. Figure 1 also indicates that the surface layer was porous. The result was explained by the inhibition of densification of HAp powders by the reaction between HAp and alumina, which the XRD analysis attributed to the decomposition of HAp to tricalcium phosphate (TCP) phase and the formation of CaAl204. The pore size as well as porosity was dependent on the quantity of added alumina powders, however, the decomposition was accelerated.
Porous Surface Layer HA p Layer Adhesive CP Layer Ceramic Substrate Figure 1. A schematic view of EPD laminated apatite composite layers on a ceramic substrate (A), and the scanning electron micrographs of sintered layers (porous surface (B), intermediate HAp (C), and adhesive CP (D)). The each bar indicates lOjim.
ElectrophoreticCoatings of Porous Apatite Compositeonto Alumina Ceramics: K. Yamashitaet al.
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In relation to bioactivity, the decomposition of HAp phase is unfavorable. When the addition of alumina was reduced to 10wt%, some portion of HAp was maintained. We also created pores by the addition of graphite powders to HAp. The mixed powders of C and HAp were well deposited by EPD in Acac/EtOH solvent. Figure 2 shows SEM photographs of the sintered surface of C-added HAp, indicating the effect of the size of graphite powders on the production of porous surfaces. The pores were created by burning of graphite during sintering. Both pore size and porosity were dependent on the amount and size of added graphite powders. Figures 3 (A)-(P^ compares the effectiveness of porous surfaces for the crystal growth in SBF; whereas the crystal growth rarely took place at grains 20jim deeper from the surface in non-added HAp layer (C), bonelike crystals were observed even at the part 80m deeper from the surface at porous HAp layer {¥). This result is attributed to the easy immersion of SBF solution into a porous layer. SUMMAR Y The effectiveness of the electrophoretic deposition technique was shown for the composite coating of porous hydroxyapatite surface and adhesive layers on ceramic substrates. The creation of pores at a surface was carried out by the addition of alumina or graphite powders to hyckoxyapatite.
Figure 2. The scanning electron micrographs of non-added HAp (A), and porous HAp surfaces (B-D), which were sintered at 1300C. The pores of the surfaces were adjusted by the addition of lOmg graphite powders with the diameter (d)=10fim (B) and 25\km (C), and 20mg with d=25^m (D).
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Figure 3. SEM photographs of fractured surfaces of non-added HAp (A, B, C) and porous HAp (D, E, F) layers, which were immersed in advance in a SBF. The photographs were taken at the points from 10 (B), 20 (C, E), and SO^im (F) from the surfaces, respectively.
REFERENCES
1. Ducheyne. P., Radine S., Heughebaert M., Heughebaert J.C., Biomaterials,1990, 11, 244254. 2. Ducheyne P., van Raemdonc W., Heughebaert J.C, Heughebaert M., Biomaterials,1986, 7, 97-103. 3. Umegaki T., Hisano Y., Yamashita K., Kanazawa T., Gypsum & Lime, 1989, No. 218, 2428. 4. Ding X., Yamashita K., Umegaki T., J .Ceram Soc.Jpn., 1995, 103, 867-869. 5. Nagai M., Yamashita K., Umegaki T., Phos, Res. Bull, 1991, 1, 167-172. 6. Garvie R.C., Urbani C , Kennedy D.R., McNeuer J.C, J. Mat, ScL, 1984, 19, 3224-3228. 7. Christel P., Meunier A., Heller M., Torre J.R, Peille CN., 7. Biomed.Mat Res,, 1989, 23, 45-61. 8. Hayashi K., Matsuguchi N., Uenoyama K., Kanemaru T., Sugioka Y., 7. Biomed,Mat, Res,, 1989, 23, 1247-1259.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
OSSEOINTEGRATIO
N IN EXPERIMENTA L HA-COATE D FEMORA L STEM S
E. De Santis, G. Rinonapoli, C. Doria, A. Manunta, MC. Sbemardori University of Sassari, Italy, Orthopaedic Dept, C.P. 40, 07100 Sassari, Italy.
ABSTRAC T The purpose of the study is to evaluate the bone-ingrowth in experimental implants of HA-coated femoral stems. Twelve sheeps were submitted to hip hemiarthroplasty with a specially designed femoral component HA-coated in the proximal 2/3 of the stem. The animals were killed at regular intervals of time after surgery. Both femora have been removed, fixed in neutral buffered formalin and submitted to plain radiographs, CT-scan, DEXA and Scanning electron microscopy. The osseointegration around the implants occurred in eleven specimens, one stem was loosened by 30 days. At 15-30 days the apposition of woven inunature bone was evident, and mature lamellar bone at 45 days. At this time, the gap between the endosteum and the HA-coated surface wasfilledby bridges of lamellar bone; this bone-ingrowth was prevalent in the proximity of the grooves and on the posterior surface increasing with time, with marked condensation of new formed bone, widely remodeled trabeculae. The layer of HA-coating tends with time to resorb; it is prevalently a slow resorption, with gradual replacement of HA by new formed bone confirming its osteoconductive properties. KEYWORDS : Hydroxyapatite, Hip arthroplasty. Prosthesis, Sheep, Osseointegration, Osteoconduction INTRODUCTIO N The goal of cementless prosthetic implants is to obtain an adequate biological stability by bone ingrowth. For several years the aim was to find the material which could allow and promote an optimal osseointegration without biological damage. Hydroxyapatite (HA), with its properties of atoxicity, antigenic inactivity, no carcinogenicity, bioactivity and biocompatibility, is the most studied material in this field. The large amount of studies on the osteoconductive property of this ceramic have reported good results [1,2,3,4,5,6,7]. We investigated about this topic with some experimental studies [8], which widely confirmed what had been found by other Authors. In the present experimental study we investigated the process of osseointegration in HA-coated femoral stems. MATERIA L AND METHOD S Twelve 2-3 years old sheeps, weighing 35-38 Kg., were submitted to unilateral hip hemiarthroplasty, with a specially designed femoral component. The stem, made of a special 467
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titanium alloy with a low modulus of elasticit y (Tilastan), was straight and cylindric in design, with a cx)llar and two longitudinal deepe r grooves in the diaphyseal portion and 11 transverse grooves each side. The proximal 2/3 of the stem had a circumferentia l HA-plasma sprayed porous-coatin g (thickness 105-10 8 micron, purity 95%, cristallinity 70%, porosity 11%) with 50-110 \i pores. A 22 mm Co-Cr alloy head was used, with a 10-12 nun morsetaper. Under endotrachea l general anaesthesia , an anterolatera l approach to the hip joint was performed. A combination of streptomyci n and penicilline (1.200.00 0 units) was administere d subcutaneousl y for 7 days postoperatively . The animals were allowed imrestricte d weight-bearin g immediatel y after surgery and walked daily during the whole postoperativ e period. They were killed by a barbiturate overdose at regular intervals of time (15, 30, 45, 60, 90, 120, 180, 200, 270, 360, 540 days). Both femora and the pelvis were harveste d and the soft tissue removed. High definition radiographs (AP+LL) , CT-scan and DEX A were taken of the implants. All specimen s were fixed in 10% buffered formaline. Cross-section s were obtained by a high-spee d diamond saw in the proximal femur at 4 levels: intertrochanteric , upper, middle and distal portion of the diaphysis filled by the stem. The sections , prepared for the observatio n by the classic procedure of fixation in 2.5% glutaraldheyde , post-fixatio n in 1% Osmium tetroxide , dehydration in acetone , drying in the critical point in C02 and metallize d in gold, were observed under scarming electro n microscope (ZeissDSM962). RESULT S In the early phases (15, 30, 45 days) the HA-coating maintained its original thickness and appeared strongly adherent to the metal surface. The gap, present at HA-bone interface , was progressivel y filled by new formed bone. Apposition of woven inunature bone, was evident at 2-3 weeks, followed by lamellar new bone with a marked trabecular orientatio n at 4-6 weeks. The bone ingrowth occurred in proximal and middle levels of the stem, where HA-coated surface was present. The process was greater in the posterior part of the bone-implan t interface . The new formed trabeculae formed bridges in the gap betwee n the endosteu m and the HA-coated surface, especiall y in the proximity of the transverse grooves and close to the medial cortex (zone 7 by Gruen). The entity of the bone ingrowth increase d with time and we saw, at 180-20 0 days, at the metaphysea l level, a great condensatio n of widely remodele d bone trabeculae which appeared to become thicker, as shown by X-ray and CT-scan. In the diaphyseal HA-coated portion of the stem, we observed the same stages of the process of osseointegratio n that we had seen in the proximal metaphysea l portion, but were slower in their development . At 200 days, it was possible to detect a different aspect betwee n the compact host bone and the new-forme d lamellar bone, that appeared emiched by residuals of amorphous HA crystals. Wefrequently detecte d an evident line of separation betwee n the host bone and the new-forme d bone, and this is correlate d to the different patterns of bone that present different mechanica l properties . In the distal uncoate d portion, the implants were surrounded by intervenin g fibrous tissue that was transformed into lamellar bone at 270 days. The beginning and the evolution of the osteogeneti c activity at the bone-implan t interface was associate d with evident changes of the HA . An actual phenomeno n was the progressive reduction of the thickness of the bioactive coating, until its complete disappearance , which was observed at around 200 days from surgery (Fig. 1). The gap.
Osseointegrationin ExperimentalHA-Coated Femoral Stems: E. De Santis et al.
469
previously filled by the HA , was then replaced by lamellar bone, differentl y arranged in the different sections of the femur. The process of gradual replacemen t of the HA by bone was precede d by morphological and structural changes, already observable at 60 days, consisting of lacunae and cracks with varying orientation . At greater distance of time (90 days), the confluenc e of crack lines led to the microfragmentatio n of the HA . In some sections we observed, already by 60 days, the presence of coarse sods in the point of maximum contact betwee n bone and stem. After resorption had been complete d (270 days) it was possible to detect amorphous residuals of synthetic HA mixed with new-forme d bone. The latter bone was totally different from the host bone and fi-omthe bone formed in the early phases of the experiment . It was lamellar bone, enriched by residuals of HA and showed a very peculiar arrangement: the lamellae, not yet mechanicall y solecitated , had not the same regular arrangement of the host bone. The resorption of the HA layer led to the appearance of a gap betwee n the new bonefi-ontand the bare metal. This gap underwen t a progressive filling with time, simultaneousl y with the forward movemen t of the new bone layer, as observed at 270 and 360 days (Fig. 2). The adaptive bone remodeling led to thickening with increase d density at DEX A in the medial and anterior cortex and thinning in the lateral and posterior cortex; it started at 45 days and progressed with time, until 540 days. Spongiosizatio n was evident, beginning at 90 days, in zone 7 under the medial collar and anteriorly, in zone 8. It increase d until 540 days. Spongious hypertrophy in zone 1 appeared at 30 days, progressing imtil 90 days and decreasing after 180 days, as shown by CT-scan and DEXA ; this technique show also an early demineralizatio n of the great trochante r that progress until 90 days and decrease progressivel y so than after 540 the mineral density at this level appear normal.
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Figure 1: S.E.M. (500X): 200 days: a gap is present betwee n new formed bone (bo) and no longer coated titanium (ti).
Figure 2: S.E.M. (550X): 270 days: the new formed bone has almost completel y filled the gap betwee n bone (bo) and metal (ti).
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DISCUSSIO N AND CONCLUSIO N Bone ingrowth is an actual phenomenon in experimental titanimn stems implanted without cement. HA-coating could enhance the process and it seems to reduce the fibrous tissue formation. The tensile and fatigue properties of bioactive materials cannot compete with the mechanical properties of metal surgical alloies. This is demonstrated by the numerous cracks in the coating that we have seen by the SEM. A possible factor influencing the resistance of the HA-coating to failure may be the thickness of the coating. The risk of disruption is assumed to increase with the thickness of the HA-coating. Excessive thickness would induce delamination and fragmentation, which would cause the loosening of an osseointegrated implant by a breakdown within the coating itself. Wide bone-implant contact and an adequate initial stability are prerequisites for a favourable evolution of the bone-implant interface. With time, the gap between the endosteum and the porous-coated surface, is filled by bridges of mature lamellar trabecular bone. This new bone formation is prevalent in proximity of the grooves and on the posterior surface of the stem. These data are indicative for a wide and early fixation of the stem by bone ingrowth, as confirmed by the SEM. The experimental insertion of a stem in the proximal femur leads to a redistribution of periprosthetic bone related to an adaptive bone remodeling. The proximal femoral remodeling in this kind of stem leads to proximal spongious hypertrophy with a thickening of the medial cortex and thinning of the lateral cortex. The last bone remodeling response is related to design and stifihess of the prosthesis. In this stem, made by a titanium alloy, the particular design and the presence of a collar, are responsible for reported modifications of the femur. Low implant stifihess, related to the used material and to the presence of grooves in the stem significantly reduced the degree of "stress-shielding". Other factors, such as the initial stability (percentage of canal fill), the location and the extent of porous coating surface, may influence the adaptive bone remodeling. REFERENCE S 1. Collier J.P., Surprenant V.A., Mayor M.B., Wrona M., Jensen R.E., Surprenant H.P.: J, Arthroplasty,1993, 8, 389-393. 2. Dalton J.E., Cook S.D., Thomas K.A., Kay J.F.: J, Bone J. Surg.,1995, 77B, 97-110. 3. Engh C.A., Hooten J.P., Zettl-Schaffer K., Ghaffarpour M., McGovem T.: J, Bone J, Surg,,1995, 77A, 903-910. 4. GeesinkR.G.T.: Clin, Orthop.,1990, 261,39-58. 5. Geesink R.G.T., Hoefiiagels N.H.M.: J. Bone J, Surg.,1995, 77B, 534-547. 6. Osbom J.F.: BiomaterialsDegradation,MA. Barbosa PrincipalEditor,Elsevier PublishersB,V., 1991. 7. Soballe K., Hansen E.S., Brockstedt-Rasmussen H., Bunger C: J, Bone J. Surg.,1993, 75B, 270-278. 8. Fadda M., Espa E., Manunta A., Rinonapoli G., De Santis E.: Italian Journal of Orthopaedicand Traumatology,1996, XXI I (2), 269-276.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EFFEC T OF HYDROXYAPATITE-COATIN G ON TH E BONDIN G OF BON E TO TITANIU M IMPLANT S IN THE FEMU R OF OVARIECTOMIZE D RAT S T. Hara, K. Hayashi, Y. Nakashima and Y. Iwamoto Department of Orthopaedic Surgery, Factory of Medicine, Kyusyu University, 3-1-1 Maidashi, Higashi-Ku, Fukuoka 812, Japan
ABSTRAC T Twelve rats were ovariectomized, and 12 rats were control. After 24 weeks, all rats had HA-coated titanium cylindrical implants inserted into the medullary canal of the right femurs and uncoated implants inserted into the left femurs. Four weeks later, the implants were harvested and subjected to mechanical push-out test. The dissected left tibias were subjected to dual energy X-ray absorptiometry (DEXA). DEXA showed a 13.4% reduction in bone density in the ovariectomized group as compared with control. The push-out strength of HA-coated implants was higher than uncoated implants in both groups. No significant difference in the push-out strength of uncoated implants was measured between the control and ovariectomized group. HA-coated implants showed a 40.3% reduction in push-out strength in the ovariectomized group compared with control. INTRODUCTIO N Many investigators have used animal models to demonstrate the osteoconductive properties of bioactive ceramics, especially hydroxy apatite (HA)[1, 2]. However, the animals used in basic studies into the properties of HA-coating had normal bone metabolism and bone mass. Therefore, we studied bone-implant shear strength of HA-coated titanium implants in ovariectomized and normal rats. The purpose of this study is to confilm the efficiency of hydroxyapatite coated on the surface of cemendess prosthesis in osteoporosis. MATERIAL S AND METHOD S Twenty four female 12-weeks-old Wistar King A rats were split into two groups. One group was ovariectomized bilaterally (OVX group). The other group had sham operation (Control group). After twenty-four weeks, HA-coated titanium implants (HA-implant, length: 23.0mm, diameter: 1.4mm + 40.0 Aim, the surface roughness: Ra4.3–0.6 A6m, the thickness of HA coating: 20.0 uvo) were inserted tighdy into the medullary canal of the right distal femur in both groups and imcoated titanium implants (Ti-implant, same length, diameter: 1.4mm, the surface roughness; Ra4.3 –0.5 Aim) were identically inserted into the left femur. Four weeks later, rats were killed and bilateral femurs with implants and left tibias were harvested. The femurs in each group were subjected to mechanical push-out test (Figure-1). The dissected left tibias were subjected to dual energy X-ray absorptiometry (DEXA). 471
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Bioceramics Volume10 Load Rat femur Implant
Wood piece
Support plate
Cement
Figure-1. Push-out test Each femur was sectioned at the proximal level of the implant edge, and the distal 2 mm of the implant was exposed. Femures were fixed in the wood pieces with cement. Specimens were placed on the special support plate and subjected to push-out test. RESULT S One rat died after sham operation. No inflammatory reactions or infections were seen in any other rats. DEXA study: The BMDs (Bone Mineral Densities) of control and OVX group were 124.93 –7.85mg/cm2 and 108.19–8.67mg/cm2, respectively. Mechanical test: In the control group, the bone-implant shear strength of HA and Ti-implants was 27.45 –4.58kgf and 2.98– 1.62kgf, respectively. In OVX group, the strength of HA and Ti-implants was 16.39 –2.07kgf and 0.81 –0.08kgf, respectively (Figure-2). There were no cement fracture between the bone and wood after push-out test.
N.S. 30
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Figure-2. The result of push out test OVX: ovariectomized group, C: control group The data are given as the mean and the standard error of the mean.(* p<0.0001)
Effectof Hydroxyapatite-Coating on theBonding of Bone to TitaniumImplants:T. Hara et al. 473
DISCUSSIO N Past studies suggested that osteoporosis reduce the fixation between porous coated implants and bone, but dose not affect the mechanical and histological integrity of HA coated implants [3, 4]. In our study, Ti-implants which had no direct bonding to bone showed low shear strength in both groups because of low surface roughness. HA-implants showed high shear strength in both groups, but there was a 40.3% reduction of the shear strength of HA-implants in the OVX group compared with control group. This study suggests that a initial fixation of cementless THR is improved by using HA-coated prosthesis but there are some limitations to the extent of bone growth with HA-coated prosthesis in osteoporosis. ACKNOLEDGEMEN T We presented this study at the 43rd annual meeting of the Orthopaedic Research Society in San Francisco, California. This work was supported by a Grant-in-Aid for Scientific Research(C) (05807133), The Ministry of Education, Science, Sports, and Culture, Japan. REFERENCE S 1. Cook, S.D., Thomas, K.A., Kay, J.F. and Jarcho, M., Chn.Orthop. 1986,232,225-243 2. Hayashi, K., Inadome, T., Tumura, H., Nakashima, Y. and Sugioka, Y.,Biomaterials 1994,15,1187-1191 3. Hayashi, K, Uenoyama, K., Mashima, T. and Sugioka Y., Biomaterials 1994, 15(1): 11-16 4. Soballe, K., Hansen, E.S., Brockestedt-Rasmussen, H., Hjortdal, V.E., Juhl, G.I., Pedersen, CM., Hvid, I. and Bunger, C , J. Arthroplasty 1991,6(4): 307-316
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BIOACTIVE BONE CEMENT
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Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
OPTIMIZATIO N OF SETTIN G TIM E AN D MECHANICA L STRENGT H OF p-TCP/MCP M CEMENTS . P. Van Landuyt, C. Lowe and J. Lemaitre Laboratory of Powder Technology, Surface Chemistry and Biomaterials Group, Swiss Federal School of Technology, MX-Ecublens, CH-1015 Lausanne, Switzerland.
ABSTRACT
In this work, the formulation and processing conditions of p-tricalcium phosphate/monocalciu m phosphate monohydrate (P-TCP/MCPM ) cement s have been optimized with respect to setting time (ST) and wet diametral tensile strength (DTS). The effects of mixing liquid compositio n (Na4P2P7 ^^d H2SO4 concentrations) , initial MCP M content, and volume of mixing liquid were investigated . The observed STs and DTS s are correlate d to the converte d P-TCP fraction and the porosity in the hardened specimens . The statistica l analysis of the results reveals that a general decay of ST and DTS occurs as the storage time of the mixing solutions increases . This decay is ascribed to the progressive hydrolysis of pyrophosphat e to orthophosphat e ions in acidic solutions, which results in a loss of the setting retarding efficienc y of the mixing solutions. Longer STs were generally observed with the mixing solutions having the highest Na4P2P7 ^i^d the lowest H2SO4 concentrations . Increasing the initial MCP M content decrease s ST only slightly, the effect being stronger for higher liquid contents . DT S is mainly affecte d by the MCP M content , the highest content resulting on the e effect on DTS . average in a 50 % increase in strength. The liquid volume has no detectabl Increasing the Na4P2P7 concentratio n of the mixing liquid increase s DTS at low H2SO4 concentration , but decrease s the strength otherwise . DTS s are clearly correlate d to the fraction of initial p-TCP converte d into brushite, which is directly related to the initial MCP M content of the specimens . From these results, optimal cement composition s have been obtained, developin g at the same time maximum DTS s (= 2.5 MPa) and adequate STs (= 14 min).
INTRODUCTIO N
Upon contact with water, p-tricalcium/phosphate-monocalciu m phosphate monohydrate (p-TCP/MCPM ) mixtures are rapidly converte d into dicalcium phosphate dihydrate (brushite, DCPD ) ^ Previous invivoexperiment s on rabbits have shown that the resorption rate of hardened cement s depends on the stoichiometri c excess of p-TCP over MCP M ^ ; the end product then consists in P-TCP/DCP D mixtures. In order to meet the surgeon’s needs, the cement s should set in about 10-15 min and develop adequate mechanica l properties within one hour. The influence of powder characteristic s on cement propertie s has been previously studied and optimized ^» ^. In this work, sodium pyrophosphat e (Na4P207, NaPP) and sulfuric acid (H2SO4, SA) were added as setting retarders in the mixing solutions of p-TCP/MCP M cements . Their effect s on selecte d caracteristic s of the cements were investigated , togethe r with the effect s of initial MCP M content and volume of mixing liquid. The characteristic s observed were : setting time (ST), diametral tensile strength (DTS), porosity and fraction of P-TCP converted . 477
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Bioceramics Volume10
MATERIALS AND METHODS
Samplepreparation. The following raw materials were used : P-TCP (Bioland, ref. BTP005-4), MCPM (Albright & Wilson, IBEX), Na4P207 (Fluka, ref. 71920), H2SO4 (Merck, ref. 731). The cement specimens were prepared by mixing for 1 min the desired proportions of P-TCP and MCPM, with an aqueous solution containing the desired concentrations of H2SO4 (SA) and Na4P207 (NaPP). The solution concentrations and initial MCPM contents expressed as weight fraction of MCPM in the dry cement powders are given in Table I. An initial MCPM content of 44.8 %wt would be needed for a complete transformation of P-TCP to DCPD ; thus, p-TCP is in excess in the cements investigated here. Characterization techniques. The STs were measured using the Vicat needle method ^. The DTSs were determined on humid cylindrical specimens, after 24 hours ageing at 20 C in 100% r.h.: the specimens were shaped in cylinders ( 0 x h = 9–0.2 x 6...13 nmi), and tested on an Adamel Lhomargy DY31 testing machine with a displacement rate of 1 mm/min. The final p-TCP conversions were determined gravimetrically : the cement specimens were roughly ground, dried to constant weight in a desiccator and weighted again after 4 h calcination at 500 "^C . The calcination weight loss is directly related to the thermal decomposition of DCPD and residual MCPM present in the sample : hence, it can be used to calculate the fraction of initial MCPM and p-TCP converted into DCPD. The open porosity of hardened cements was measured by the Archimede’s method, using isopropanol as the immersion liquid. Statistics. The effects of mixing liquid composition, MCPM content and liquid volume were studied in 16 randomized runs, using the 2^ factorial statistical design summarized in Table I. Mechanical tests were performed in triplicate. The results were analysed using the ANOVA technique, the effects of the factors being adjusted by a multilinear model with interactions ^ ; the adjusted values were calculated according to the appropriate statistical model, taking into account the regression parameters found significant at the (1 - p) confidence level (p is mentioned in the figures). Residual analysis was used to check the adequation of the statistical models. Table I. Experimental design for the study of setting time and diametral tensile strength of calcium phosphate cements. Factor A) [NaPP] in the mixing liquid B) [SA] in the mixing liquid C) Initial MCPM content (dry powder) D) Mixing liquid volume
RESULTS AND DISCUSSION
Levels Low
High
0.05 M O.IOM 30 wt% 0.40 mL/g
O.IOM 0.20 M 40 wt% 0.45 mL/g
Preliminary statistical analysis. The aqueous mixing solutions containing both NaPP and SA were prepared at once at the beginning of the experiment. Then, the 16 cement specimens were prepared successively in the selected randomized order, which took about 4 h in total. Inspection of the residuals (i.e. the differences between the experimental responses and the results calculated according to the retained statistical model) highlights a chronological decay of ST and DTS. The ST decrease corresponds to a progressive decay of the retarding efficiency of the mixing solutions, which can be reasonably explained by the slow hydrolysis in acidic solutions of pyrophosphate into orthophosphate ions. For the
Optimizationof Setting Time and Strengthof P-TCPjMCPM § 1 Adj. Val. A Exp. Val.
Biocements:P. Van Landuyt et al.
479
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p < 0.05
^
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o
Treatment
Treatment
Figure 1. Setting times (log scale) of Figure 2. Diametral tensile strengths of specimens prepared at 20 C (see text). aged cement specimen s (see text). The treatment s are coded as follows : - Label ’a’ means that the highest level of factor A [Na4PP] = 0.05 M is applied, the lowest level [Na4PP] = 0.10 M being applied otherwise . - Label ’b’ means that the highest level of factor B was applied [SA] = 0.10 M the lowest level [SA] = 0.20 M being applied otherwise . - Label ’ab’ means that both factors A and B were applied at their highest levels, the other factors being applied at their lowest level. - The meaning of the other labels is deduced according to the same logic. 0.8
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Treatment Treatment Figure 3. Fraction of converte d p-TCP in Figure 4. Open porosity of aged cement aged cement specimen s (see text) specimen s (see text) subsequen t discussion, the experimenta l values of ST and DTS have been correcte d for this chronological drift. Setting time.The ajusted values of ST calculate d from the statistica l model are presente d in Figure \. STs range from 4 to 20 min, dependin g on the treatment . The most significant factors (p < 0.01) are [NaPP] and [SA]: increasing [NaPP] from 0.05 to 0.10 M induces a marked ST increase , whereas ST is systematicall y lower at higher [SA].
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Less important factors (p < 0.05) are the mixing liquid volume and the MCP M content: increasing the former slightly increase s ST, whereas slightly shorter STs are observed at higher MCP M contents . Hence, the longest STs are obtained by combining the highest [NaPP] with the lowest [SA]. Fractions of ^TCP andMCPMconverted. As expected , the fraction of p-TCP converte d increases significantly with the initial MCP M content (Fig. 3). Increasing [SA] also increases slightly the converte d p-TCP fraction. The converte d MCP M fraction . (77.1–2. 4 % ) is only marginally affecte d by the starting compositio n of the cements Openporosity. The open porosity ranges from 34 to 43 %vol , dependin g on the treatmen t (Fig. 4). Increasing the MCP M content or [SA], and decreasin g the liquid volume lower the cement porosity. As seen previously, higher initial MCP M and SA content s increase the converte d (3-TCP fraction, and hence the DCP D content of the hardened cement . Since DCP D has a lower density than the starting reactants , and since the apparent volume of the material remains constant, the solid volume fraction must increase at the expense of porosity. The decrease in porosity is larger as more DCP D is formed upon cement hardening. Diametral tensilestrength. Figure 2 shows that DTS is signifiantly improved by higher MCP M content s (p < 0.01). This effect is correlate d to the higher final P-TCP conversion , and hence the higher DCP D content of the cement s containing initially more MCPM . [SA] and [NaPP] interact slightly (p < 0.05): at lower [SA], DTS generally increase s with higher [NaPP], the opposite being observed at higher [SA].
CONCLUSIONS
ST is mainly influence d by [NaPP] and [SA] in the mixing liquid. A practical ST of at least 10 min can be achieve d with mixing solutions containing 0.10 M NaPP and 0,10 M SA. These additives have to be dissolved seperatel y in order to prevent the alteration of NaPP in acidic solution. DT S increase s with higher initial MCP M contents , which promotes higher degrees of PTC P conversion , higher amounts of DCP D and lower porosities in the hardened cements . The best compromise betwee n acceptabl e ST (= 14 min) an optimal DTS (= 2.5 MPa ) is achieved by cement s ’ac’ and ’acd’.
ACKNOWLEDGEMENT S
This work has been supported by the Board of the Swiss Federal Schools of Technology (PPM project n 4.2D). The support of the Robert Mathys Foundation, of Stratec Medical, and of Ciba Pharma (Switzerland) is gratefully acknowledged . Bioland (France) is gratefully acknowledge d for kindly providing the P-TCP powder used in this study.
REFERENCES 1 Mirtchi A. A., Lemaitre J., & Munting E. Biomaterials 10 (1989) 634-638 . 2 Ohura K., Bohner M., Hardouin P., Lemaitre J., Pasquier G. & Flautre B. J. Biomed. Mater. Res. 30 (1996) 193-200 . 3 Andrianjatovo H., Jos6 F. & Lemaitre J. J. Mater.Sci.-Mater.Med. 7 (1996) 34-39. ^ Van Landuyt P. & Lemaitre J. Bioceramics, Vol. 9 (Kokubo T., Nakamura T. & Miyaji F. eds.) Elsevier (1996) 205-208 . 5 Bohner M., Lemaitre J. & Ring T. A. J. Amer. Ceram. Soc. 79 (1996) 1427-1434 . ^ Montgomery D.C., DesignandAnalysis of Experiments, (3rdEd) Wiley & Sons New York, 1991.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
INFLUENC E OF THE PARTICL E SIZ E OF THE POWDE R PHAS E IN THE SETTIN G AND HARDENIN G BEHAVIOU R OF A CALCIU M PHOSPHAT E CEMEN T MP. Ginebra, E. Fernandez, F.C.M. Driessens, M.G. Boltong, J.A. Planell. Department of Materials Science and Metallurgical Engineering, Universitat Politecnica de Catalunya, Av. Diagonal 647, E-08028-Barcelona, SPAIN
ABSTRAC T The effect of the particle size of the powder phase of a calcium phosphate cement based on the hydrolysis of a-tricalcium phosphate on its setting parameters and mechanical behaviour is studied. This cement leads to the formation of a poorly crystalline calcium deficient hydroxyapatite, compositionally closer to bone mineral than stoichiometric hydroxyapatite. Two starting powders with different particle sizes were studied. The final setting was reduced from 45 to 8 minutes when the fineness of the cement powder increased. Also the cohesion time was significantly reduced. The hardening of the cement was accelerated when the particle size of the cement powder was reduced. The compressive strengthening rate increased by a factor of 5 in the first 24 h, whilst the final compressive strength at 360 h was similar for both cements. The evolution of the observed microstructure after different periods of time explained this behaviour. KEYWORD S Calcium phosphates, bone cement, hydroxyapatite INTRODUCTIO N Calcium phosphate cements are potentially very usefiil materials in a great variety of medical and dental applications; their major advantages are their in situmoulding ability and their excellent biocompatibility [1]. However, since their mechanical properties are poor, it is necessary to investigate some ways to improve them. It is known that, in general, the mechanical behaviour of a cement system is determined by factors affecting the powder phase, the liquid phase, the mixture of the powder and the liquid and the aging conditions. Specifically, the fineness of the powder phase achieved by milling is a key factor in determining the characteristics of the hardening process up to the final strength [2]. Taking this fact into account, the effect of the particle size of the powder phase of an apatitic calcium phosphate cement on its setting and hardening behaviour, in order to be able to design a material which complies with the clinical requirements has been investigated. In a previous work it was verified that the transformation which takes place during the setting and the hardening of this cement is the hydrolysis of the a-tricalcium phosphate (a-Ca3(P04)2 , a-TCP) to a calcium deficient hydroxyapatite [3]. 481
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MATERIAL S AND METHOD S The cement powder consisted of 85% a-TCP and 15% p-TCP, according to quantitative XRD analysis [4]. It was prepared by heating in air at 1300 C for 15 h and then quenching to room temperature in air an appropiate mixture of CaHP04 (Merck 02144) and CaCOa (Merck 2076). Two cement powders with different particle sizes were studied. To obtain them the TCP was milled for two different periods of time (30 and 360 minutes) in an agate ball mill (Pulverisette 6, Fritsch GmbH) at two different rotating speeds of the mill (460 and 640 rpm respectively). The particle size distribution of the two resulting powders was measured by means of laser diffraction. From now, on the cement with a finer powder will be coded as cement F, and the cement with a coarser powder as cement C. In both cements a 2 wt% of precipitated hydroxyapatite (Merck 2143) was added as a seed in the powder phase. The liquid phase consisted of an aqueous solution 2.5 wt% of disodium hydrogen phosphate (Na2HP04). The liquid-to-solid ratio used was 0.32 ml.g’V Powder and liquid were mixed in a mortar for about 1 minute, and the initial and final setting times of the cement paste were determined with Gillmore needles according to the C266ASTM standard. The cohesion time, defined as the time after which the cement did not suffer disintegration when immersed in Ringer’s solution was determined by visual inspection [5]. In order to evaluate the strength development cylindrical specimens with a diameter of 6 mm and a height of 12 mm were moulded. After 15 min, they were immersed in Ringer’s solution at 37T for periods of 1, 2, 4, 8, 16, 32, 64 and 360 h. For each period, 8 samples were prepared. After the immersion, the samples were polished, removed fi-om the mould and tested under compression at a cross-head speed of 1 mm/min. Microstructures of the fracture surfaces were investigated by scanning electron microscopy. RESULT S AND DISCUSSIO N Figure 1 shows the cumulative curves for the particle size distribution of the two cements studied. The median size was 10.88 fxm for the coarse powder and 2.22 ^m for thefineone. The values obtained for the initial setting time (I) and the final setting time (F) and for the cohesion time (CT) are given in Table 1. As expected, setting was much faster for the fine cement, since the specific surface of reaction was higher and the reaction could take place at a
;
/
:
f /
/
;
/ . ^^ / ^ 1
atlA^’
cemen t C A cemen t F i
1 « IIml
1
:
1 . 1 n i l.
size (^m)
Figure 1. Particle size distribution for the two cements studied
Cement C
1 F
I (min) 9.5 4
F (min) 45 7.5
TC (min) 1 7
3
1
Table 1. Initial (I) and final (F) setting times and cohesion time (CT) for the cements with coarse (cement C) and fine (cement F) particle size of the powder phase.
Influenceof Particle Size of Powder Phase in Setting and Hardening Behaviour:M.P. Ginebraet al.
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much higher rate in the initial stages. In addition, it has to be kept in mind that toghethe r with the chemical reaction that is taking place, the physical attractive forces betwee n the cement particles are stronger when they are smaller, contributing to the cement setting. The same arguments can explain also the fact that the cohesion time is reduced when the fineness of the cement powder t with the Kozenyincreases. On the other hand, the reduction of the cohesion time is in agreemen Carman equation, which relates the permeabilit y coefficien t of a porous solid with the inverse of the square of its specific surface [6]. The strength developmen t of both cement s is shown in Figure 2. In both cases a statistica l model could be fitted by a weighte d least-square s regression analysis to the evolution of the compressive strength with the reaction time. The compressiv e strength varied with time according to
c{ty c
/
1 - -e
-A r
V
(1)
)
The values for Coo (final compressiv e strength) and T at a 95% confidenc e level are shown in Table 2. It is clear that the hardening of the cement is accelerate d when the particle size of the cement powder is reduced. The compressiv e strengthenin g rate at the initial stages (CJT) increases by a factor of 5. This is reasonable , taking into consideratio n that chemically the mechanism of hardening consists first in the dissolution of the cement powder, and second in the precipitatio n of a product phase, in this case a calcium deficien t hydroxyapatite . Both mechanisms are accelerate d when the particle size of the powder increases . Since its specific surface increases , the dissolution is favoured and the liquid phase around the solid particles become s more supersaturate d with respect to the phosphate and calcium ions, favouring a more rapid precipitatio n of the product phase. The final compressive strength at 360 h is slightly higher for the fine cement than for the coarse one. Since physically the cement hardening is due to the entanglemen t of the crystals of the
Q.
[Cement A
C
cemen t F
1 F
cemen t C
50
100
150
200
250
300
350
C^ (MPa ) 34.6(1.0 ) 40.1(1.8 )
T(h)
28.2(2.2 ) 1 6.1(0.9 )
400
t(h)
Figure 2.- Compressive strength C of the two cements studied as a fimctionof the reaction time.
Table 2. Fitting parameters for the hardening curves of cement s C and F. (Standard deviation betwee n brackets).
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product phase, the morphology and size of these crystals can influence the final strength reached by the cement. In our case, the microstructural study showed that the crystals formed after the setting and hardening of the finer cement were much smaller (0.1- 1 ^im) than in the case of the coarse cement (1-10 i^m). AKNOWLEDGMENT S This investigation was supported in part by a grant from the Direccion Cientifica y Tecnica of Spain. The authors thank the CICYT for funding this work through project MAT940911. REFERENCE S 1. Driessens F.C.M, Planell J.A. and Gil F.J., Calcium phosphatebone cements.In: D.L.Wise, D.J. Trantolo, D.I. Altobelli, M.J. Yaszenski, J.D. Gresser and E.R. Schwartz, Encyclopedic Handbook of biomaterialsand bioengineering,Part B: Applications,Ed. Marcel Decker, NY, USA 1995, 855-877. 2. Lea F.M., The Chemistry of Cement and Concrete,3rd. Ed. Edward Arnold Publishers Ltd., London 1973 3. Ginebra M.P., Fernandez E., De Maeyer E.A.P., Verbeeck R.M.H., Boltong M.G., Ginebra J , Driessens F.C.M. Planell J.A., Setting reaction and hardening of an apatitic calcium phosphate cement, J. Dent. Res. 1997 76 (in press). 4. Fernandez E, Boltong MG, Ginebra MP, Planell JA, Driessens FCM (1995). "Effect of Ca2Si04 on the stability of a-TCP down to low temperatures". In: Proceedings of the 4th Euro Ceramics Conference, October 2-6 1995, Riccione, pp. 103-108. 5. Fernandez E, Boltong MG, Ginebra MP, Driessens FCM, Bermiidez O., Planell J.A., "Development of a method to measure the period of swelling of calcium phosphate cements", J. Mater. Sci. Letters1996 ,15,1004-1005. 6. Carman P.C, Flow of gases throughporous media.Academic Press, NY 1956.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
SUBCUTANEOU S TISSU E RESPONSE S TETRACALCIU M PHOSPHAT E CEMENT S
AND
KINETIC S
OF
CELL S
TO
M. Yoshikawa\ H. Oonishi^ Y. MandaP, F. Sugihar and T. Toda* ^Department of Endodontics, Osaka Dental University, 5-31, Otemae 1-chome, Chuo-ku, Osaka 540, Japan. ^Department of Orthopedic Surgery, Artificial Joint Section and Biomaterial Research Laboratory, Osaka-Minami National Hospital, 2-1, Kido-higasi, Kawachinagano 586, Japan. ^Nitta Gelatin Inc. Research Laboratory, 2-2, Futamata, Yao 281, Japan ABSTRAC T Calcium phosphate cements were experimentally developed. The powder phase of the cements comprised an equimolar mixture of tetracalcium phosphate and dicalcium phosphate dihydrate (TeDCPD). The liquid phase consisted of Mcllvam’s buffer solution of two-fold concentration (CM-1), the buffer solution containing calboxymethyl cellulose sodium salt (CM-2) or the buffer solution containing chondroitin-sulfate in place of the salt (CM-3). The biocompatibility of three cements, a compound of TeDCPD kneaded with glycerin (CM-4), and TeDCPD powder (PC) were estimated in rat subcutaneous tissue. Macrophages and multi-nuclear giant cells were present around CM-1, CM-2 and CM-3. CM-4 caused inflammation and necrosis in the adjacent tissue and induced numerous macrophages and fibroblasts. It was concluded that infiltration of cells to the border of implanted cements relates to the solubility and disintegration or setting tune of cements. KEYWORD S Tetracalcium phosphate cement. Macrophages, Multi-nuclear giant cells, Chondroitinsulfate, Glycerin, Subcutaneous tissue INTRODUCTIO N Three calcium phosphate cements (CM-1, CM-2, CM-3) using an equimolar mixture of tetracalcium phosphate (TeCP) and dicalcium phosphate (DCPD) were developed for use in endodontic treatment and orthopedic surgery. The cements will be used as a root canal sealer or direct pulp capping agent in endodontic clinics and as a bone cement m orthopedic surgery. Results of fundamental studies concerning periapical and subcutaneous tissue responses to CM-2 [1] and CM-3 [2] have been presented, and it was confirmed that these calcium phosphate cements have fine biocompatibility, causing relatively little acute mflammatory reaction in surrounding tissue. More detailed in vivo studies must be carried out before the cements can be applied clinically. The purpose of this study was to examine kinetics of cells in the boundary area of subcutaneous tissue after hnplantation of experimentally developed cements. MATERIAL S AND METHOD S Cement preparations Tetracalcium phosphate (TeCP) was obtained by the dry synthetic method of sintering a stoichiometric mixture of calcium phosphate dihydrate and calcium carbonate (mol ratio of Ca/P=2.0) at 1480 C for 8 hours, in the laboratory of Nitta Gelatin Inc.. The particles were 0.6 485
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to 44.0 |im, average of 12 to 13 jam, in size. DCPD was purchased from Wako Pure Chemical Inc.(Osaka, Japan) and was passed through a 32 jim-sieve. The powder phase of developed cements comprised 68% TeCP and 32% DCPD (TeDCPD). The liquid phase was Mcllvain’s buffer solution of two-fold concentration (CM-1, P/L=1.2g/g), the buffer solution contauimg 1.5% carboxymethyl cellulose sodium salt (CM-2, P/L=1.4g/g), or the buffer solution containing 5% chondroitin-sulfate A sodium salt (CS, Kanto Kasei, Tokyo, Japan) in place of the salt in CM2 (CM-3, P/L=1.5g/g). A compound of TeDCPD and glycerin (Nakarai Tesk Inc., Kyoto, Japan) was also produced (CM-4, P/L=1.2g/g). Experimental procedures The pH of cements was measured from 5 to 60 minutes after kneading at 5 or 10 minute intervals [3]. The results were represented as averages and standard deviations and significant differences about each group were examined with Student’s T-test. The responses of subcutaneous tissue and the kinetics of cells around the TeDCPD compounds were examined histopathologically. The TeDCPD compounds and TeDCPD powder (PC) were respectively placed m the dorsal subcutaneous tissue of 7-week old Sprague Dawley rats. At 1, 2, 3 and 4 weeks after preparation, the implanted materials were removed including the surrounding tissue. They were fixed, decalcified, embedded in paraffm, serially sectioned at 6 fim, stained with hematoxylin and eosin (H.E. staining), and histopathologically examined. RESULT S AND DISCUSSIO N Table 1. pH of TeDCPD compounds Compounds pH
CM-1
CM-2
CM-3
CM-4
8.6–0.2
8.4–0.3
8.5–0.4
*7.7–0.3
Figure 1. Four weeks after implantation of CM-1 A small number of Mcj) are seen around the implanted material. There are no cells inside the implant. C: Implanted material (CM-1) f: fibrous connective tissue (H.E. stain, Original Mag. x32.7)
Figure 2. Three weeks after implantation of CM-2 C: Implanted material (CM-2) Arrows: Proliferated cells and fiber in the cracks f: Fibrous connective tissue (H.E. stam. Original Mag. x30.0)
SubcutaneousTissue Responses to TetracalciumPhosphate Cements:M. Yoshikawa et al.
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Figure 3.
One week after implantation of CM-3 Numerous MGC (arrows) are seen around CM-3 (C). A small number of cells are present. (H.E. stain, Original Mag. x80.0)
Four weeks after implantation of CM-3 Numerous cells and fibrous connective tissue are present in the inner part of CM-3 (C). (H.E. stain, Original Mag. x25.0)
Figure 5.
Figure 6.
One week after implantation of CM-4 Tissue adjacent to CM-4 is necrotic (N). Numerous M(|), MGC and fibroblasts are present in the surrounding tissue. (H.E. stain. Original Mag. x43.2)
Figure 4.
One week after implantation of PC Only a small number of MGC are seen around PC (C). Many cells are infiltrated the implant. (H.E. stain. Original Mag. x43.2)
pH value of cement s None of the cements showed significant differences of pH with time. CM-4 significantly differed fi-om the other groups (Table 1). The liquid phase of CM-4 would not prevent buffer action of TeDCPD. So, the pH value of CM-4 was likely maintained at approximately 7.7. Kinetics of cells around the materials Few macrophages (Mcj)) and multi-nucleate giant cells (MGC) were present in the adjacent tissue of CM-1 at any observation period (Figure 1). As no other cells, such as fibroblasts, were
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present inside the implants, CM-1 was considered to have less biocompatibility than CM-2 and CM-3. Numerous M
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BIOLOGICA L BEHAVIOU R OF A BIOACTIV E BON E CEMEN T IMPLANTE D IN RABBI T TIBIA E A. Afonso(l), M. Vasconcelos(l), R. Branco(l), J. Cavalheiro(2) (1) Faculty of Dental Medicine, University of Porto, Rua Dr. Roberto Frias, 4200 Porto (2) INEB - Institute Engenharia Biomedica, Pr. Coronel Pacheco 1, 4000 Porto, Portugal ABSTRAC T A bioactive bone cement was prepared and adjusted to surgical needs. The bioactive cement and an acrylic cement (PMMA) were investigated in an in vivo study. The cements were placed in tibiae of rabbits. Five and ten weeks after the implantation, histological characterisation using light and scanning electron microscopy, was performed and histomorphometric evaluation was conducted. The bioactive bone cement specimens demonstrated significantly good performance when compared with the acrylic cement in respect to bone ingrowth around the implant and direct bone contact formation between the cement and the bone tissue. INTRODUCTIO N In orthopaedic and dental implantology, the immobilisation of artificial implants in bone tissues continues to be a major field of interest[l]. The usual cement used in orthopaedic surgery to attach and fix the metal implants to the bone is an acrylic cement (PMMA)[1,2]. There are several advantages and disadvantages in using PMMA[2]. The advantages of PMMA are a) it is surgically forgiving; b) fast fixation of implants; and c) good intrusion of the cement into the bone matrix, providing good initial fixation. The disadvantages are a) the cement is an order of magnitude, weaker than the bone or implant resulting in property mismatch at interfaces; b) poor cement distribution around the implant; c) shrinkage of the cement during polymerisation; d) chemical and thermal necrosis of surrounding tissue due to exothermic heat of polymerisation; and e) unreacted monomer release which can cause a large transient drop in blood pressure[2]. These several disadvantages in using PMMA as bone cement increase the needs to develop new cements with improved bioactive and mechanical properties. Our research group were together in the process of developing and testing, in biological conditions, a bone bioactive cement. The aim of this work is to study the in vivo behaviour of a bioactive bone cement using PMMA as a control material, after implantation in rabbit tibiae during 5 and 10 weeks. Histological characterisation using light and scanning electron microscopy was carried out and histomorphometric evaluation of the bone-to-cement contact was conducted. MATERIAL S AND METHOD S A powder of glass was prepared to produce a bioactive cement. Its properties and degradation behaviour was published before[3]. Briefly, a mixture of oxides was melted in a platinum crucible during 4 hours. Particles were then quenched in water, dried and grounded in agata ball mill. The nominal composition (wt%) of this glass was: Si02 35.6, CaO 42.4, P2O5 17.0 and NaOs 5.0. 489
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A paste was prepared by adding an aqueous solution containing NH4HPO4 and MH4H2PO4. The pH of this solution was 7.4. The setting time was measured by Gilmore’s needle tests. Initial setting time were modified by heat treatment of the glass powder, the size of the particles and the relationship of the powder/solution. The compressive strength of the bioactive cement is 41Mpa. The in vivo behaviour of both materials was assessed by implanting the cements in the right posterior tibiae of adult rabbits. All rabbits were operated on by a standard procedure in aseptic conditions. They were given intramuscular anaesthesia complemented with local anaesthesia. After performing a skin incision and obtaining a periosteal flap to expose the anteromedial face of the tibiae, the implant site was prepared with a spherical burr with continuous cooling. The hole was then well packed with de cements and the wound was closed in two layers. Rabbits used in this assay were sacrificed 5 and 10 weeks after the surgical procedure. The bone blocks were immersed in 4% neutral buffered formalin. Subsequently, specimens were dehydrated in a series of alcohol’s and embedded in a methyl-methacrylate resin. After polymerisation, specimens were sectioned with a diamond saw to a thickness of 200-250|Lim and ground down to about 30|Lim with a disc to prepare histologic slides for the light microscopy. Slides were then stained with hematoxylin and eosin. In the samples prepared for scanning electron microscopy, the samples were sectioned with a diamond saw to expose the bone surface with the implanted materials and polished with a series of SiC-paper discs using a successively finer grain size to obtain a smooth surface. Samples were gold-coated and examined in a JEOL JSM 6301F scanning electron microscope at an accelerating voltage of 15kV. Histological characterisation was performed under light and scanning electronic microscopy. Percentage of bone contact was measured using a curvimeter device. RESULTS AND DISCUSSION The reaction between the glass powder and ammonium phosphate solution allowed the initial fluid paste to be transformed into a solid. The product of this reaction is a crystalline calcium ammonium phosphate. Not all the particles of glass will react with the solution and original glass particles could be detected in the cement. The heat treatment of the material, the use of particles with different size, and the change of the relationship of the solid/liquid will modify the behaviour of the cement. 100
Particle size;jm \
^
Plastictim e
I n i t i al setting time
Figure 1. Influence of the powder particles on the plastic and setting times
Figure 2. Layer of fibrous tissue (arrows) PMMA cement, LM (XI00)
Biological Behaviour of a Bioactive Bone Cement Implantedin Rabbit Tibiae: A. Afonso et al.
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Figure 3. Light micrograph of the cements,5 weeks after implantation (X40). A - PMMA, B - Bioactive cement When the particles were heat treated above 500 C for 30 minutes, the initial amorphous structure of the glass was changed with the precipitation of a crystalline phase as it could be detected using x-ray diffraction. This phase is more stable than the original quenched structure and consequently, the reaction with the solution was slower. The time necessary to transform the viscous liquid into a paste was here called "plastic time". The plastic time and the time necessary to support the first Gilmore needle without a visible mark (the initial setting time) are represented in Figure 1. The large difference of the specific surface of the reaction of the particles, can explain the range of time achieved. By changing the relationship between the amount of the powder and the liquid used to prepare the paste, we can also provide large differences in the plastic time. By using particles with a medium size of 2.3|im, a decrease in the relationship of the solid/liquid from 2 to 1, will increase the time needed to consume the phosphate and the plastic time to be extended from 1.5 to 11 minutes. In all samples with PMMA, 5 and 10 weeks after implantation, no direct contact with bone was achieved and the cement was lost during histological preparation. The loosening of this cement in the first 10 weeks was reported in other studies[4]. A layer of connective tissue was observed in the bone surface (Fig.2). We can also, observe same signs of bone reabsorption and inflammatory cells (Fig.3a). This connective tissue layer persisted for long weeks in other studies[5]. In the samples with the bioactive bone cement, bone contact and regeneration is shown at light microscopy level, 5 and 10 weeks after implantation, without inflammatory or other adverse signals (Fig.3b). Bone deposition on the interface and cancelous bone formation in the external surface could be seen 5 weeks after the implantation. The cement was partially involved by this external cancelous bone. In the interface, the woven bone appears with large vessels and in direct contact with the cement surface without interposition of fibrous tissue. In the samples retrieved 10 weeks after the implantation, bone formation around the surface of the cement was complete and showed a more organised bone structure (mature bone). In the scanning electron microscopy, 5 weeks after the implantation (Fig.4), we can see the bone in direct contact with the cement. This bone presents a lamellae organisation and the presence of large sized vessels.
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PMMA Bioactive Cement
5 weeks 0 n=3 75.01–12.2% n=3
10 weeks 0 n=3 87.32–9.64% n=3
Table 1. Percentage of bone contact between the cements and the bone tissue
Figure 4. SEM view of the interface between the bioactive cement and the bone (5 weeks)
The histomorphometric evaluation permits us to obtain the values of the percentage of bone contact between the cements and the bone tissue. These values are presented in table 1. The acrylic cement didn’t present direct bone contact, because a fibrous tissue layer were interposed between the cement and the bone. The percentage of bone contact between the bioactive cement and the bone was increased with time. CONCLUSION S It’s possible to achieve a material that can be easily used in surgery with a careful control of the heat treatment, the size of the particles and the relationship of solid/liquid used in the preparation of the cement. The bioactive cement has a good bone integration and a high percentage of bone contact without signs of adverse reactions inversely to the PMMA cement. Moreover, the cement possesses an adhesion ability to the bone and direct bonding of the cement to the bone was observed in all the samples. The percentage of bone contact was increased with time. This bioactive cement can be used as a paste with a setting time of between 2-8 minutes, and changes into a rigid compound without any exothermic reaction, which is an excellent property for a bone cement REFERENCE S 1. Ishihara K, Aral H, Nakabayashi N, Morita S, Furuya K: J. Biomed.Mat. Res. 1992; 26: 937-945. 2. Henrich DE, Cram AE, Park JB, Liu YK, Reddi H: J. Biomed.Mat. Res. 1993; 27: 277280. 3. Carvalho B, Cavalheiro J: Proc. Int. Conf. on Medical Physics & BiomedicalEngineering in Nicosia - Cyprus, 1994:346-351. 4. Morberg P, Albrektsson T: J. Mater.Sci: Mater.Med 1992; 3: 170-174. 5. Taguchi Y, Yamamuro T, Nakamura T, et al.: J. Applied.Biomat.1990; 1: 217-223.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
HISTOLOGICA L STUD Y OF A DCPDBASE D CALCIU M PHOSPHAT E CEMENT . Patrick Frayssinet* , Laurent Gineste^, Philippe Conte**, Jacques Pages*, Nicole Rouquet*, Alain Lerch* * **
Bioland, 132 route d’Espagne, 31100 Toulouse, France Service de Chirurgie Orthopedique, CHU Toulouse Purpan, Toulouse, France Laboratoire du Tissu Osseux et des Pathologies Osteoarticulaires, Universite Paul Sabatier, Toulouse, France. ^ Service de Biologie Buccale. Ecole Dentaire, Universite Paul Sabatier, 31400 Toulouse ABSTRAC T A DCPD based calcium phosphate cement has been implanted and set within rabbit condyles. Histological sections were performed at 2, 6 and 18 weeks. After a mild foreign body reaction characterized by the presence of mononuclear cells at the implant surface, woven bone trabeculae formed from the cavity edges onto the material. The implant was fragmented in the early time of implantation and the fragments were invaded by cells before to be integrated within the newly formed bone matrix. At 18 weeks, the fragments had migrated in the whole condyle and the most of them were integrated while the implant volume had decreased. In conclusions : this calcium phosphate cement is a fast degradable material that can sustain bone apposition. INTRODUCTIO N Self-setting calcium phosphate cements are materials that can be handled by the surgeon in paste form and injected into bone cavities or defects. They then set to form a mineral matrix at the contact of which healing bone tissue can form. Replacement of the mineral by bone is supposed to occur as a result of the same process taking place at the calcium phosphate ceramic contact. We have developed a new calcium phosphate hydraulic cement that can set in a wet environment and be injected with a syringe into a bone defect or fracture site. The cement consists of a solid phase containing p-tricalcium phosphate and sodium pyrophosphate and a liquid phase composed of orthophosphoric and sulfuric acid. The setting reaction can be summarized as follows: Ca3(P04)2+H3P04+6H20-*3CaHP04,2H20.The physico-chemical properties of the cement can be modified by the purity and the granulometry of the tricalcium phosphate. Dicalcium phosphate is known to be one of the most soluble of calcium phosphate phases (1) and can be useful when material degradable within a few weeks is needed. The aim of the described experiment was to check that the material set when injected by syringe into a bone cavity, and to examine the short term effects of the material on the surrounding tissues. MATEIOAL S AND METHOD S Animals : The animals used for experiments were 9 female New Zealand white rabbits aged 6-7 months with body weights of 3700-5100 g and which had free access to food and water. The animals were observed daily for one week after implantation and three times a week for the remaining implantation period. The skin in the implanted region, animal mobility and behavior were noted. 493
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Material injected : Powder : 3 g (P-TCP : 2.954 g, anhydrous sodium pyrophosphate : 0.046 g) Liquid : 1.8 ml (phosphoric acid (4 M ) and sulfuric acid (0.1 M ) solution). In this study, the powder/liqui d ratio was 1.3 and the conversio n rate 95 % . The resulting solid material consiste d of 95% of dihydrated dicalcium phosphate , 4 % of |3-tricalciu m phosphate , and 1% of various mineral phases (sodium based and calcium sulfate). The density was 1.43, the porosity 45% and the mean pore diamete r 7. l^m. Injection of bonecement: The cement was inserted under anaesthesi a with intramuscular ketamine , through a lateral longitudinal skin incision over the knee, in holes drilled in the externa l condyles of rabbits. Holes were drilled in the externa l condyles of the right and left legs. The left hole was filled with the cement and the right hole was left empty as control. Each hole was 4 mm in diamete r and 5 mm in length. The cement was injecte d through a syringe in paste form. 3 animals were sacrified by nembutal injectio n after 2 weeks, 6 weeks and 18 weeks and the distal extremitie s of the femurs were collected . Histological processing: The samples were fixedin a 4 % formaldehyd e solution for 48 h, dehydrate d in increasing ethanol solution then embedde d in PMM A (polymethylmethacrylate) . Sections 5 |im thick were obtained by a hard tissue microtome (Reichert-Jun g Type E). They were then stained with Giemsa solution and by Von Kossa method. The remaining histologica l blocks were gold palladium-coate d and observed by SEM under a back-scattere d mode operate d at 10 Kv. RESULT S ^After two weeks: the implantation zone was occupied by a solid cylinder made of mineral grains that had been partly removed fi-om the section by the microtome blade. Thin newly-forme d trabeculae originating fi-omthe cortical zone were inserted perpendicularl y to the implant surface. These trabeculae were surrounded by an osteoid layer lined with active osteoblasts . Bone marrow cavities betwee n these trabeculae were filledwith connectiv e tissue containing many fibroblast-like cells and bone trabeculaefi-agments. No hematopoieti c cell could be found in these bone marrow cavities at this time. Some cell condensation s at the origin of connectiv e tissue trabeculae could be seen in the vicinity of the implant, some of which were sjmthesizin g an EC M (extracellula r matrix) undergoing mineralization . Resorption marks were apparent at the implant surface into which the tissue had grown. SEM showed that the material was constitute d of small white particles disseminate d within a grey matrix with some micropores. Several grey-level s were visible in the s setting. matrix suggesting an unhomogeneou The control bone defect was filledwith a loose connectiv e tissue in the form of a network with connectiv e tissue trabeculae several tens of ^m in thickness . Ossification had begun at the edges of the defect . After six weeks: the implants showed a mineral core which still appeared as a solid. The externa l cylinder region was invaded by bone trabeculae . The bone marrow cavities betwee n the trabeculae were completel y filledwith implantfi-agments. Thesefi-agmentsembedde d within the bone tissue contained cells betwee n the crystals. The bone trabeculae structure was immature with many
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remodelling marks. An organic matrix was visible in some sites at the implant surface and had mvaded the space available betwee n the material grains. Many grains detache d from the material were phagocytose d by macrophages and giant cells. SEM reveale d the presence of materialfragmentsfar from the implantation zone in the condyle A dissolution zone 100 to 300^m thick under the material surface could be observed in one implant. Mineralizing bone trabeculae were ingrowing in this zone. The material core lined by this dissolution zone was edged by a densifie d layer less than 100 ^m thick. The control defect s were occupied by low density trabecular bone in which the bone marrow cavities containe d cytologicall y normal marrow. ^After eightee n weeks: small grains emitte d by the material were less numerous in the bone marrow cavities surrounding the fragmentedmaterials. The remaining implants were fragmented with two or three big fragmentstotally lined by bone trabeculae parallel to thefragmentsurfaces. This hmng bone layer was linked to the cortical bone by bone trabeculae . The SEM showed that the material was a solid with pores ranging from a few to one hundred inm. White grains of different sizes ranging from one or two to several tens of ^m were dispersed within a grey matrix. Bone matrix showing different grey levels was in contact with the material Some white particles could be found in the Havers canals and were probably grouped within the histiocytes . Some such particles were included within the bone matrix. The control zones were difficuh to differentiat e from the rest of the bone tissue. However, the low density of trabeculae in this region was characteristi c of the control zone. Figure 1 : Back-scattere d SEM of a two weeks implanted cement ( C). Very immature bone trabeculae (->) and bonefragmentsare in contact with the implant surface. The implant shows different grey levels due to different densitie s suggesting that the cement did not set homogeneousl y after implantation. Some implant fragments are detache d fromthe surface. Bar : 1 mm.
Figure 2 : Interface of a six week implanted cement (IC ) with bone tissue (BT) with osteocyte s (os) at the cement contact. The cement contains two phases. The white particles are made of TCP. Bar: lO^m.
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DISCUSSIO N AND CONCLUSIO N This study showed that : 1 / The cement set when injected into a bone marrow cavity and formed a porous calcium phosphate structure; 2/ Two different calcium phosphate phases with diflferent solubility rates could be identified by SEM observation.; 3/ A mild foreign body reaction occurred in the vicinity of the material with a majority of mononucleated cells. Calcium phosphate particles were phagocytosed in these cells ; 4/ The degradation rate of the material was relatively high when compared to the usual HA and TCP ceramics but compatible with the ingrov^h of bone formingtrabeculae within the resorbing material. Different fast-setting calcium phosphates have been described that all show osteoconductive properties and can be injected (2, 3, 4, 5). In this experiment, the dissolution products induced a low foreign body reaction. However, the presence of macrophages and giant cells in the implantation site did not impair bone formation within the same site. Some differentiating osteoblasts synthesizing an osteoid matrix were identified among the macrophages. No increase in the bone resorption could be detected at any time. The ossification process occurring at the material contact was not exactly identical to that occurring at the contact of calcium phosphate ceramics (6). It is frequent that an organic matrix deposit is found at the material surface ingrowing between the crystallites of the mineral matrix. This matrix is not lined by osteoblasts like osteoid and does not precede the formation of bone at the material surface like it is at the ceramic surface. On the opposite, this matrix could follow the formation of bone in the close proximity to the mineral. Osteoblasts are not, or only very rarely, found at the material surface as is observed at the surface of calcium phosphate ceramics. Most of the osteoblasts differentiate at some distance from the material and there is a neat ingrowth of the forming bone trabeculae from the cavity edges into the spaces made available by the material resorption. There is no material degradation resulting from a creeping substitution process as has been shown at the ceramic contact. REFERENCE S 1- NancoUas, H, In vitro studies of calcium phosphate crystallisation. In : Mann, S, Webb, J, Williams, RJP, (eds) Biomineralization. Chemical and Biochemical Perspectives. VCH New York, 1989,ppl57-182 2- Driessens, FCM, Boltong, MG, Bermudez, O, Planell, JA, Formulation and setting times of some calcium orthophosphate cements: a pilot study. Journal of Material Science:Material in Medicine 1993 ; 4: 503-508 3- Driessens, FCM, Boltong, MG, Bermudez, 0, Planell, JA, Ginebra, MP, Fernandez, E, Effective formulations for the preparation of calcium phosphate bone cements. Journal of Material Science: Material in Medicine 1994 ; 5 : 164-170 4- Ishikawa, K, Takagi, S, Chow, LC, Ishikawa, Y, Properties and mechanisms of fast-setting calcium phosphate cements. Journal of Material Science:Material in Medicine 1996 ; 17 : 14291432 5- Kurashina, K, Kurita, H, Hirano, M, de Blieck, JMA, Klein, CPAT, de Groot, K, Calcium phosphate cement: in vitroand in vivo studies of the a-tricalcium phosphate-dicalcium phosphate dibasic-tetracalcium phosphate monoxide system. Journal of Material Science: Material in Medicine 1995 ; 6 : 343-347 6- Frayssinet, P, Trouillet, XL, Rouquet, N, Azimus, E, Autefage, A, Osseointegration of macroporous calcium phosphate ceramics having a different chemical composition Biomaterials 1993 ; 14: 423-429
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BIOACTIV E BON E CEMEN T STUDIE D IN CANIN E TOTA L HI P ARTHROPLASTY , 2 YEAR S FOLLO W UP STUD Y H. Fujitai, T. Nakamurai, K. Idol, y. Matsudai, H. lidai, M. Kobayashii, M. Oka2, and Y. Kitamura^ 1 Department of Orthopaedic Surgery, Faculty of Medicine, Kyoto University. Kawahara-cho 54, Shogoin, Sakyo-ku, Kyoto 606, Japan, 2 Research Center for Biomedical Engineering, Kyoto University, Kyoto, Japan, 3 Nippon Electric Glass Co. Ltd., Ohtu, Japan ABSTRAC T Polymethybnethacrylate (PMMA) bone cement which has been widely used in total hip arthroplasty (THA) has several problems. To improve these problems, the authors have developed a new bioactive bone cement, which has a capability of bonding directly with bone, and has greater mechanical strength than PMMA bone cement. In this study, THAs were performed in dogs using bioactive bone cement (BABC) consisting of AW glass- ceramic powder and Si02 powder as the filler and bisphenol-a-glycidyl methacrylate (Bis-GMA) based resin as the organic matrix, and the outcomes were compared with the results of PMMA bone cement. The bonding strength of BABC to bone in dogs’ femora was 3.4 Mpa at 2 years, whereas that of PMMA bone cement was 1.7 MPa (p<0.05). Histological exammation showed dh-ect bonding between BABC and bone, whereas in PMMA bone cement, a fibrous tissue layer was evident at the bone cement interface. No degdadation inside of the cement was observed. Under weight bearing condition, such as THA, this bioactive bone cement revealed high bonding strength and direct bondmg to bone. KEYWORDS ; Bioactive Bone Cement, AW-glass-ceramic, Bis-GMA, Canine, Total Hip Arthroplasty INTRODUCTIO N PMMA bone cement which has been widely used in THA has several problems such as nonbonding behavior with bone, high heat generation during polymerization, and relatively low mechanical strength. To improve these problems, we have developed a new bioactive bone cement. In previous study, it has been revealed that this cement forms a chemical bond with living bone through a Ca-P-rich layer and has a high mechanical strength [1,2,3] . Matsuda et al. reported that bioactive bone cement consisting of bioactive glass powder and Bis-GMA-based resin showed a higher bonding strength than PMMA cement in canine THA up to 6 months [4] , In this study, THAs were performed in beagle dogs using bioactive bone cement consisting of AW glass- ceramic powder and Si02 powder as the filler and Bis-GMA-based resin as the organic matrix. In Bioceramics 9, follow up period we reported was up to 6 months. In this study, we followed up to 2 years and the outcomes were compared with the results of PMMA bone cement. MATERIAL S AND METHOD S Cement Preparation BABC consists of AW glass-ceramic powder and Si02 powder as thefiller(AW glass ceramic 497
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: Si02=72 : 28, by weight ratio) and Bis-GMA-based resin as the organic matrix. AW glass ceramic powder (average particle diameter ; 4.0 tim) has a nominal compositio n of CaO AA.1%, Mg O 4.6% ,Si02 34.0%, P2O5 16.2%, and CaF2 0.5% by weight ratio and has a crystal phase of apatite and woUastonite . An average particle diamete r of Si02 powder is 3.5 iim. Both powders were treate d with r -methacrylox y propyl trimethox y silane. Bis-GMA-based resin was prepared from equal weights of Bis-GM A and triethylene-glyco l dimethacrylat e (TEGDMA) . The authors prepared two different composition s of BAB C (filler : resin= 85:15 for the acetabula r side, filler : resin= 79:21 for the femoral side, by weight ratio) because the manual cementin g technique for the acetabular side and the cement injectio n technique for the femoral side are now generally acknowledge d in human THAs . Increased ratio of powder for the fixation of the acetabula r component allowed manual handling of the cement , and decrease d ratio for the femoral side allowed the cement injectio n into the femoral canal. Commercially available PMM A bone cement ( CM W 1 for the acetabula r component , CM W 3 for the femoral component ) was chosen for the control. Mechanical Properties of Bone Cements The ultimate compressive strength, elastic modulus, tensile strength, bending strength and fracture toughness of the bioactive bone cement s and PMM A bone cement (CMW 3 ) were measured. Cylindrical specimen s (6 mm diameter , 12 mm long) were prepared for the compression test and specimen s measuring 2 X 2 mm in cross section m the middle portion and 4 X 2 mm at both ends were made for the tensile test. Rectangular beams (20 X 4X 3 mm) were prepared for the bending strength test and rectangular beams (20 X4X 3 mm) with a notch were n was soaked made for thefracturetoughness test. After the composite had hardened, each specime in simulated body fluid at 37 C for 1 day and then subjecte d to a test using an Instron-type testing machine (Autograph model AG-10TB, Shimadzu, Kyoto, Japan) at a cross-hea d speed of 0.5 mm/min. The bending andfracturetoughness tests were carried out by the three-poin t method. The values of elastic modulus were calculate d from the stress-strai n curve of the bending test. Hip prosthesis The hip prosthesis consiste d of an acetabula r component , 18 mm in outer diameter, made of ultra high molecular weight polyethylene , a 12 mm femoral head made of stainless steel (SUS316L), and a straight stemmed , collared femoral componen t made of stainless steel (SUS-316L). The femoral shaft was 4 mm in diamete r and 50 mm in length. Animal Experiment Adult beagle dogs weighing 9.5 to 10.5 kg were anesthetize d by intramuscular injectio n of ketamine HC L (10 mg/kg body weight) and atropine sulfate (1.0 mg/dog). By the transtrochanteri c approach, the hip was dislocate d and a femoral osteotom y was performed proximal Table 1 Mechanical Properties of Bone Cements (mean –S.D.) BAB C Type of Cements BAB C (Injection type) (Dough type) Compressive Streng;th(MPa) 2 2 1 – 1 3* 269–14* Elastic Modulus(GPa) 9.7–0.4 * 14.4–0.9 * Tensile Strength(MPa) 5 1 – 1# 5 2 – 4# Bending Strength(MPa) 1 1 8 – 8* 1 3 9 – 6* Fracture Toughness(Mpami/2 ) 1.61–0.04 # 1.63–0.04 # * All pairs were significant (ANOVA , p<0.05). # Significant versus PMM A (ANOVA , p<0.05).
PMM A (CMW3 )
9 5 – 1* 2.7–0.1 * 25–1 9 5 – 2* 1.14–0.0 9
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to the lesser trochanter . The articular surface of the acetabulu m was reamed with 2 reamers to 20 mm in diamete r to expose the subchondral bone and 7 anchoring holes were made with a 2.5 mm drill. The acetabula r componen t was fixedto the pelvis with bone cement manually. The femoral canal was serially reamed to remove the spongiosa. Bone cement was injecte d into the femoral canal from the bottom to the top. The femoral componen t was held securely by hand until polymerizatio n of the cement had completed . After the hip joint was relocated , greater trochante r was reattache d using two SUSS 16 stainless wires. The hip was then checked for stability and range of motion. Finally,the wound was closed in layers. 8 THA s were performed, 4 hip joints in each group. All dogs were sacrificed at 2 years after surgery by intravenou s overdose of pentobarbital . Push Out Test The femur was cut into 5 mm sections , axial to the long axis. The sections were numbered from proximal to distal in sequence . Odd numbered sections were used for the histological analysis and even numbered sections were subjecte d to the push out test. The cement and femoral component were pushed togethe r with a metal rod 6 mm in diamete r at a cross head speed of 0.5 mm per minute, using an Instron type testing machine (Autograph Model AG-IOTG , Shimazu, Kyoto, Japan). The bonding strength at the interface was calculate d by dividing the failure load by the total area of bone cement interface . Histological Examination The odd numbered 5-mm section from the femur and the pelvis was fixedin 10% phosphate buffered formalin and was completel y potted in methacrylate , dental bone cement . Alcohol or organic solvents were not used to avoid damaging the PMM A bone cement [5] . Undecalcifie d 150 to 200 ixmthick slides were made using a cutting machine (BS-3000, EXAKT , Norderstedt , Germany) and a grinding machine (MG-4000, EXAKT , Norderstedt , Germany). Interface of the cement and the bone was evaluate d by Giemsa surface staining using light microscope and by scanning electro n microscope (Hitachi S-800, Hitachi, Tokyo, Japan) equipped with an energy
(a)
(b)
Figure 1. Radiographs taken 2 years after surgery using BABC , (a) A-P view, (b) lateral view of the femur.
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dispersive Xray microanalyzer (EMAX-3000, Horiba, Tokyo, Japan). RESULTS Mechanical Properties pf Bpng Cementg Mechanical properties of bone cements are shown in Table 1. BABC showed higher mechanical properties than PMMA (p<0.05, ANOVA). Clinical Results All dogs were able to bear their body weight within 1 week and walk without limping by 3 weeks. At the time of sacrifice, neither infection nor abnormal inflammatory reaction of the hip was shown on gross examination. In 1 dog in the PMMA group, at the time of harvest, the femoral component could be moved manually due to stem fracture following cementfracture.In the remaining dogs, all the implants appeared to be securely fixed to the skeletons at the time of sacrifice (Figure 1). Push Out Test Bonding strength of BABC was 3.4 – 0.7 MPa (mean – S.D.) at 2 years, however, that in the PMMA was 1.7 – 0.7 MPa. There was a significant difference in the results between the 2 groups (p<0.05, ANOVA). Histological Examination (a) Giemsa Surface Staining Femoral side In the BABC, the surface of the cement was well covered with regenerated bone and direct bonding between BABC and the bone was observed. In the PMMA, a thick fibrous tissue layer was evident at the bone cement interface. Acetabular side In the BABC, cement fractures were observed in some specimens. In the PMMA, a thick fibrous tissue layer was evident at the bone cement interface. ( b ) SEMEPMA Femoral side In the BABC, the cement showed dfa-ect bonding to bone without any intervening soft tissue layer. In the PMMA, an intervening soft tissue layer consistently existed at the bone cement interface. By electron probe microanalysis, the silicon level decreased while calcium and phosphorus level increased slightly across the bone cement interface. Ca-P-rich layer approximately 30 Mm in thickness was present on the surface of the bioactive bone cement, and the cement bonded directly to host bone through this layer. DISCUSSIO N We have been trying to make bioactive bone cements using bioactive glass or glass ceramic powder combined with Bis-GMA resin [1,2,3,4] . These cements showed good bioactivity and high mechanical strength. In the present experiment the authors attempted to evaluate new cement consisting of AW glass ceramic powder and Si02 powder as a filler and Bis-GMA resin under weight bearing condition up to 2 years. This cement showed good bioactivity and ability to bond directly with bone in femoral side. Cementfracturesof the acetabular side may be caused by thin cement mantle in this experiment. REFERENCES 1. Kawanabe, K., et al., J. Appl. Biomater., 1993, 4, 135-141. 2. Tamura, J., et al., J. Biomed. Mater. Res., 1995, 29, 551-559. 3. Tamura, J., et al., J. Biomed. Mater. Res., 1996, 30, 85-94. 4. Matsuda, Y., et al., Clin. Orthop., in press. 5. Jasty, M., et al, J.Bone Joint Surg., 72-A, 8, 1220-1229.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EXPERIMENTA L STUD Y ON CHITOSA N FILLER S
HYDROXYAPATIT E / N-CARBOXYMETHY
L
R. Martinetti^, L. Dolcini ^, A. Ravaglioli ^, A. Krajewski ^ and C. Mangano ^ ipIN-CERAMICA FAENZA s.r.l. Via Ravegnana, 186 48018 Faenza, Italy ^IRTEC-CNR, Faenza, Via Granarolo n. 64, 48018 Faenza, Italy 3 Dental Center , Via Trento n. 16, 22015 Gravedona, Como, Italy
ABSTRAC T Hydroxyapatite / N-Carboxymethyl chitosan (HA / N-CMC) fillers were prepared and studied as a new biomaterial for medical use. N-CMC could either be a suitable structure where calcium and phosphate ions could arrange to begin mineralisation process or to influence the metabolism of the cells, inducing the synthesis of specific protein able to control the natural mechanism of hydroxyapatite crystals growth and consequently the mineralisation of new osteoid tissue. Starting materials and final samples were characterised, morphology and microstructure were evaluated by SEM and IR spectroscopy was used to detect the typical bands. N-CMC solution used for the preparation of thefillerscoating was also characterised from the reological point of view. This experimental work underlines the important role of N-CMC for biomedical application, when in association with hydroxyapatite. KEYWORD S N-Carboxymethyl chitosan coating, porous hydroxyapatite, bioactive fillers. INTRODUCTIO N Chitosan is particularly considered for its properties to stimulate restorative processes of the tissues. Literature on m v//ro studies with osteoblasts [1] reports that, when chitosan is present, the Gla-protein, a protein able to control the hydroxyapatite crystal deposition is detected. Chitosan is an important interface between bone and hydroxyapatite because it promotes an osteoconductive reaction during the filling of irregular bone defects [2]. It stimulates bone tissue regeneration and helps the wound healing, in association with calcium phosphate compounds it can reduce the healing of bone defects. Chitosan allows Van der Waals bonds [3], strong enough with phosphate calcium salts. It is known, e.g.that the mixture chitosan-hydroxyapatite, calcium oxide and zinc oxide brings to the formation of a body reliable for different shapes as that of the spongy sheet usefiil for the regeneration of the bone tissue. In N-CMC the repeating units are N-Acetylglucosamine and N-Carboxymethyl glucosamine, its polimer chain is linear. 503
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The carbonate group of N-CMC may weakely integrate itself in the hydroxyapatite structure lattice on the interface planes trying to constitute the correspondent carbonate-apatite. In some way the Van der Waals bond increases constituting a kind of chelating for epithetical deposition of the N-CMC molecules. N-CMC could so promote enlargement of specific crystallographic planes of HA. The promoted planes will be those on which PO4 ^’ can be easily substituted by the carbonate groups pf N-CMC. This is particularly allowed when the distance between a carbonate group and the other along the N-CMC molecula is egual to a lattice distance between a suitable P04^" lattice site and another on the interfacial plane exhibited by HA. The presence of glycine units on the biopolymer is a further reason of affinity to the human tissues which contain collagen mainly constituted by glycine MATERIAL S AND METHOD S N-CMC used in this study is a glycosaminoglycan obtained from chitosan. It carries glycine units and its molecular weight is 700 K Dalton. Its nature and its general chemical behaviour are those of an amphoteric polysaccharide, where the cationic nature is prevailing. It combines in itself the main characteristics of chitosan and glycine. Fully miscible with water and with water solutions; it is included in the field of biopolymers, which exhibits positive actions on the human body. It is an excellentfilm-formingagent. It is proved that it stimulated the ordered growth of damaged tissues.
O
g
3500 3000 2500 2000 1500 1000 Wavenumber [cm’^] Figure 1. IR spectrum: a = HA/N-CMC, b= HA, c= N-CMC
ExperimentalStudy on Hydroxyapatite/N-CarboxymethylChitosan Fillers: R. Martinettiet al.
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The HA granules, used to prepared the proposed fillers in associatio n with N-CMC , have low density, were made by granulation method (patent n. B093A 000435 ) [4] starting from hydroxyapatite powder. A solution of N-CM C with 1.0 wt% concentratio n was prepared and by aerosol method, used to obtain a thin coating of about 1.0 |im on hydroxyapatit e granules. The samples prepared were dried at 37 C for 6 hours and then sterilise d by y - ray irradiation at 25 kGy. The N-CM C solution was characterise d from the reological point of view using a Rotovisco RV2 0 (Haake device). The IR spectra in the solid state, were collecte d for N-CMC , HA and HA / N-CM C with a Bruker 113v FT IR spectromete r on KB r pellets as support in the spectral range 4000 - 400 cm" ^. The pellets were prepared with 0.7, 1, 1.5 and 2.0 mgr of sample and 200 mg of KBr . SE M (Leica Cambridge mod. 360 Stereoscan , Englan ) was used to study the microstructur e and the morphology of the starting materials and of the final sample. Preliminary invivo tests were performed on rabbit femurs. RESULT S AND DISCUSSIO N The pH of N-CM C solution was controlle d during the coating process: it was 4.5; the : Eta = 27.6 viscosity was measured at D (1 / s) = 1000, the following data was determined (mPas). The infrared spectrum shown in the figure Ic, identify the chemical nature of N-CMC : it shows adsorption bands at 1620 - 1630 cm"^ correspondin g to the substitute d amines and imines while the typical band of carboxyl group was detecte d at 1730 cm"^ while the infrared spectrum shown in the figure lb, identify the chemical nature of hydroxyapatite . The sample studied is identifie d by the infrared spectrum shown in the figure la: a general overlap of spectra lb and Ic is deducible with intensificatio n of the main bands, in particular in the range betwee n 1700 and 1200 cm"^ it is not possible to determine if an interferenc e exists betwee n the materials. The microstructur e of starting materials and of experimenta l sample are shown in figure 2: in figure 2a the morphology of N-CM C layer; in figure 2b the microstructur e of HA granules (crystals size 0.05 - 0.1 \\m)\in figure 2c the surface of N-CM C coating on HA granule, while, in figure 2d the interface HA/N-CMC .
I A .... . if; #im
. . ’ ^ fe’’-....-
% ^
H ^»’’ieijm
J
’’’’^i.. J r ’IM. (2a)
(2b)
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(2c )
(2d)
Figure 2. SEM morphology: 2a = N-CMC layer, 2b = HA microstructure, 2c = N-CMC coating on HA granule, 2d = interface HA / N-CMC Preliminar hystological results relatively specimens retrieved after in vivo tests on rabbit femurs (2 weeks) proved that HA / N-CMC fillers have good biocompatibility with bone and shown a direct contact with newly formed bone. Further tests are in progress. AKNOWLEDGMEN T We would like to express our thank to Dr. Valentina Biasini (IRTEC-CNR, Faenza) for her assistance with SEM observation and Dr. Gigliola Lusvardi (Universita degli studi di Modena, Dipartimento di Chimica) for her kind suggestions on ER spectra. REFERENCE S 1. F. Lo Bianco, C. Parodi, S. Spinato, L. Lo Bianco: DENTAL CADMOS , 1993, 10, 40-46. 2. M. Mattioli, G. Biagini, R. Muzzarelli, C. Castaldini, M.G. Gandolfi, A. Krajewski, A. Ravaglioli, M. Fini, R. Giardino: J. of Bioactive and Compatible Polymers , 1995, 10, 249-257. 3. M. Takechi, Y. Miyamoto, K. Ishikawa, M. Yuasa, M. Nagayama, M.Kon, K. Asaoka: J. ofMaterial Science: Material in Medicine, 1996, 7, 317-322. 4. A. Ravaglioli, A. Krajewski, A. Piattelli, C. Mangano, R. Martinetti: Tadashi Kokubo, Takashi Nakamura, Fumiaki Miyaji, Japan, Bioceramics Vol. 9,1996, 185-188.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
INJECTABL E CHITOSAMIN E HYDROXYLAPATIT
E BON E PAST E
J.J.Railhac\ P.Sharrock^, D.Galy-Fourcade\ CZahraouP & N.Sans^ ^ Service Central de Radiologie, Centre Hospitaller Universitaire de Toulouse-P^iipan ^ Laboratoire de Chimie Inorganique, Universite Paul Sabatier, Toulouse, France. ABSTRAC T We have studied a hydroxylapatite (HA) bone cement composed of a liquid containing lactic acid and chitosamine in water - glycerol and of a solid containing HA and calcium hydroxide. The mixture of both components provides a paste which settles as a mass, stabilized by the chitosamine. In vitro and in vivo results show that dissolution of calcium lactate occurs with some loss of particulate matter. A specially prepared injectable form, with zinc oxide as a radioopacifier, may be placed in epiphyseal bone by catheterization with a diaphyseal percutaneous approach. We feel new biomaterials should be developed together with advanced surgical techniques. KEYWORD S Hydroxylapatite cement, chitosamine, lactic acid, catheterization. INTRODUCTIO N Recent developments in calcium phosphate cements stem from interest in having a biocompatible filler in injectable form. Coral granules and calcium phosphate powders have been dispersed in water-based gels which can be handled in syringes [1]. Gels, however, do not have sufficient strength to stablize particles in a bleeding area. Cements can be made to harden by an acid-base reaction which produces a salt which settles as a crystalline hydrated mass [2]. Such cements are very sensitive to parameters such as particle size and reactivity which in turn depend on the thermal history of the particles and the amount of water used in the reaction [3] as well as the initial pH. Several calcium phosphate based mixtures have been elaborated with short or long hardening times [4,5,6,7]. The resulting cements, once hardened, have been implanted and shown to be osteoconductive, or resorbable and replaced by new bone in animal experiments [8]. Their in vivo use, however, seems to be limited to small fractures or the filling of tightly closed cavities [9]. Alternatively, the cements may be mixed, modelled into the desired shape, and then implanted. This avoids problems related to aqueous dilution of the components during in vivo hardening. We now report preliminary results concerning an HA based cement stabilized with chitosamine [10]. MATERIAL S AN D METHOD S The calcium hydroxide, lactic acid, zinc oxide and glycerol were from Fluka, Buchs, Switzerland. The hydroxylapatite used was Ossatite made by MCP, Toulouse, France. Chitosamine was isolated from a Rhizopus strain provided by Gist-Brocades and also obtained by hydrolysis of crab chitin by heating in alcoholic potasium hydroxide. Catheterization experiments were realized on fresh human amputation specimens and also on live sheep. Standard orthopedic equipment was used to test various transcortical drilling angles by a percutaneous approach. Soft tissues were protected by a drill sheath. Several probes and straight or curved catheters were used to test shaft penetration. The range of sizes went from 5 to 16 Fr. The internal canule of a Van Sonnenberg 507
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drain from Meditech was used to attach syringes via the Luer Lock mechanism to test bicomponent paste injection. RESULTS AND DISCUSSION We initially developped an HA cement with calcium lactate as the hardening reagent. The product was marketed in France in 1966 and may be considered as a self hardening paste with HA as a « non-reactive » filler dispersed in a calcium lactate solidifying system. The results are good, particularly in endosseous situtations such as total hip replacement with bone reconstruction, but in nearly 50% of the cases followed, problems related to HA particle migration are reported. When the HA paste is in contact with healing soft tissues or heamatomae, the matrix dissolves, liberating HA into the surrounding environment. White, non-infectious, milk-like losses have been observed through the scars accompanied by a cutaneous inflammatory reaction, which subsides without harmfuU effects. These results, which were not observed previously during animal trials carried out in the bones of rats, rabbits and dogs, could be reproduced by implanting HA paste in a rat subcutaneous model. Previous experiments with pure HA granules implanted in soft tissues led to the conclusion that HA is osteoinductive because new bone tissue could be observed in a nonosseous site. We feel this should be explained as showing the need for osteoclasts in resorbing larger HA particles. HA thus triggers a cell differentiation mechanism which helps to remove HA from vascularized non-osseous tissue. Macrophages involved in wound clean-up can also carry away small HA particles. Polybetaglucosamine (chitosamine) was used as an organic binding agent in an effort to stabilize the self hardening paste. Chitosamine is a natural polymer [11] present in fungi or present in N-acetylated form in crab shells. The chitosamine had a high molecular weight and high viscosity in solution because we used short hydrolysis times. Chitosamine is dissolved in waterlactic acid solution at a 4% concentration and then mixed with the HA powder containing 5% basic Ca(0H)2. The mixture yields a paste which hardens in 3 minutes. The solidification reaction occurs even underwater. The use of glycerol and slow reactive mixtures allows injection of the soft viscous mass through a 2 mm i.d. canule of 30 cm length. Figure 1 shows the pH of a dilute chitosamine suspension in distilled water as a function of the mole fraction of various added acids to glucosamine monomer. To obtain these curves, the chitosamine was freshly precipitated from acid solution and rinsed to a constant neutral pH. Strong acid such as HCl protonates the free amino groups and dissolves chitosamine at a one to one acidbase ratio. Complete dissolution occurs at the inflection point near pH 2.5. Weak organic acids dissolve chitosan at higher pH values (3.5 for lactic acid and 4.5 for acetic acid). Lactic acid consumes a larger amount of acid to do this with a two to one acid-base ratio. Figure 2 illustrates the titration curve of a lactic acid solution with or without chitosamine, by adding Ca(0H)2 containing HA. The presence of chitosamine effectively lowers the pH of the buffer zone corresponding to calcium lactate formation from pH 6.5 to pH 5.5. Deprotonation of soluble chitosamine salts modifies the pH slightly and precipitates the insoluble chitosamine binder. This reaction occurs simultaneously with the neutrahzation of lactic acid by calcium hydroxide. Further addition of base reveals the equivalence point, which is not modified significantly by the presence of chitosamine, but is influenced by the presence of zinc oxide. Zinc oxide, used as a radioopacifier, acts as a weak base consuming some lactic acid and also binding chitosamine in complex formation. The X-ray diffraction patterns of the HA paste before and after ageing in water for one month reveal the presence of hexagonal zincite. The absence of crystalUne compounds other than HA and zincite is to be noted as illustrated in figure 3. Thermogravimetric analysis shows a 7.72 weight% content of combustible organic matter lost at 600’’C. This weight loss decreases to 4.78%
InjectableChitosamineHydroxyapatiteBone Paste: JJ. Railhac et al.
509
PH 10
without £hitosamine ^t without ZnO
6
1
2
(acid/base ) molar ratio
gHA
Figure 1. Titration of chitosamine with a) HCl, b) lactic acid, c) acetic acid.
Figure 2. Titration of 2.8 ml of 4% lactic acid with HA containing 5% Ca(0H)2.
following ageing in water. These results may be compared with the 12% total weight loss observed for oven-dried samples submerged in water one week and dried again. Thus only part of the organic matter dissolves in water but most of the soluble components are in fact mineral salts. Thus chitosamine in low concentration effectively binds HA particles together but does not prevent dissolution of soluble calcium lactates. A 40’’ oblique drilling of the cortex was found to be optimal to perforate long bones in a diaphyseal percutaneous approach. The pasage in melaphyso-epiphyseal spongy bone was possible with straight probes. The oblique angle then allows the introduction of catheters through a coaxial guiding canule. CT guidance prooved usefull in confirming an epiphyseal goal which could be reached by selecting the proper approach on the cortical perimeter. Bone maiTow could be aspirated, and epiphyseal bone inigated and drained. The HA naste containing zinc oxide could be visualized
u 15
16
THETA
18
Figure 3. X-ray diffraclion pallcrn of lx)ne pasle after ageing in water stiowing /incile(*) in HA.
Figure 4. X-ray picture of lx)ne shaft with injected HA lx)nc paste.
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properly with X-rays as shown in figure 4. The standard bi-component formula had to be adapted to enable syringe injection. Smaller particle sizes and slower reaction rates were used to lower viscosity and facilitate flow. Partial cement softening was observed, indicating a stronger cohesive formulation is needed if stabiUzation of injected cement by total hardening is required CONCLUSION HA containing bone paste can be delivered into a bony site by catheterization or classical surgical techniques. A distant diaphyseal percutaneous approach a minima to treat various longbone sites is interesting when a direct approach would increase bone fragility or cross sensitive anatomical areas. This technique may help develop the use of injectable biomaterials and prevent the loss of particulate HA from bicomponent mixtures during the hardening reaction. Organic polymers may serve as binders and possibly as hosts for therapeutic agents such as growth factors loaded on ceramic supports.
REFERENCES 1. 2. 3. 4.
5. 6. 7. 8. 9. 10. 11.
A.Dupraz, T.P.Nguyen, M.Richard, J.Delecrin and G.Daculsi, « Microcharacterization of a new injectable polymer/ceramic composite as bone substitute in spine surgery», JMater.Sci,Mat.in Med.,1 (1996) 52-55. W.Brown and L.Chow, « Dental restoration cement pastes » US patentW A 518 430 (1985). L.Chow, M.Markovic, S.Takagi and M.Cherng, « Injectable calcium phosphate cements: Effects of cement liquid on the physical properties of the cement », Innov. Techn.Biol. Mec/., 18(1) (1997) 11-14. K.Kurashina, H.Kurita, M.Hirano, J.M.A.de Blieck, C.P.A.T.Klein and K. de Groot, « Calcium phosphate cement: in vitro and in vivo studies of the tricalcium phosphatedicalcium phosphate dibasic-tetracalcium phosphate monoxide system », J.Mater.Sci.Mat. in Med.,6 (1995)340-341. M.Takeshi, Y.Miyamoto, K.Ishikawa, M.Yuasa, M.Nagayama, M.Kon and K.Asoaka, « Non-decay type fast-setting calcium phosphate cement using chitosan », J.Mater.Sci. Mat.in Med.,7 (1996) 317-322. F.C.M.Driessens, M.G.Boltong, O.Bermudez, J.A.Planell, M.P.Ginebra and E.Femandez, « Effective formulations for the preparation of calcium phosphate bone cements », J.Mater.Sci.Mat. in Med.,5 (1994) 164-170. K.Kurashina, H.Kurita, A.Kotani, H.Takeuchi and M.Hirano, « In vivo study of a calcium phosphate cement consisting of tricalcium phosphate/dicalcium phosphate dibasic/tetracalcium phosphate monoxide », Biomaterials,18 (1997) 147-151. A.Mirtchi, LLemaitre and E.Muntig, «Calcium phosphate cements: Study of the tricalcium phosphate-dicalcium phosphate-calcite cements » Biomaterials,11 (1990) 83-88. B.Constantz, I.Ison, M.Fulmer, R.Poser, S.Smith, M.VanWagoner, J.Ross, S.Goldstein, LJupiter and D.Rosenthal, « Skeletal repair by in situ formation of the mineral phase of bone », Science.267 (1995) 1796-1798. M.Nagase, R.B.Chen, Y.Araya and T.Nakajima, « Evaluation of a bone substitute prepared from tricalcium phosphate and an acid polysaccharide solution », J.Oral Maxillofac.Surg., 49(1991)1305-1309. S.Bhaskara and Chandra P. Sharma, « Use of chitosan as a biomaterial: Studies on its safety and hemostatic potential », J. Biomed.Mater.Res.,34 (1977) 21-28.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MANUFACTUR E OF A HYDROXYAPATIT
E -CHITI N COMPOSIT E
A.C.A. Wan’, E. Khor’*, G.W. Haslings"’^ Department of Chemistry’, Biomat Center, Department of Mechanical and Production Engineering’’ and Biomedical Materials Program, Institute of Materials Research & Engineering , The National University of Singapore, Kent Ridge, Singapore 119260.
ABSTRAC T We have developed a combination of chitin, a natural polymer found in the exoskeleton of crustaceans and insects with calcium hydroxyapatite (HA) to produce a new "composite" material. HA powder was incorporated into chitin via a solution processing method to form an intimate mi.xture. Upon casting of this mixture into molds of fixed dimensions and subsequent removal of solvent. HA containing chitin flexible plates are obtained. The amount of HA was varied from 10% to 50% by mass of HA. The tensile modulus, yield stress and elongation to fracture, measured at a cross head speed of 5 mm/min, of these HA containing chitin flexible plates were evaluated. Results indicated the inability of HA to reinforce chitin, however the plasticity of chitin was retained at high loading of filler. KEYWORD S hydroxyapatite, chitin, composites, solution processing INTRODUCTIO N Chitin exists in nature as the reinforcing fiber for naturally occurring composites such as the mollusk and cnistacean shells, in conjunction with an inorganic component like calcium carbonate. In the natural state, chitin possesses good mechanical properties due to hydrogen bonding between the chitin polymer chains, leading to high crystallinity. In analogy to its role in these biocomposites, a chitin-HA material can be conceptualized for potential application in the orthopedic field, i.e. as a bone substitute. The chitin matrix plays a dual role- as a binder to prevent migration of HA from the implant site and to act as a scaffold pending the regeneration of host bone. The latter is possible due to chitin’s biodegradability, its (3-1,4 glycosidic linkages being susceptible to the lysozyme present in humans.]!] The HA component, besides being osteoconductive, may modify the mechanical properties of the chitin matrix as demonstrated by a similar bone analog comprising polyethylene and HA. [2] 511
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The incorporation of HA into chitin presents a technical difficulty. Chitin, like many other cellulosic materials decomposes before melting can be achieved, due to strong intermolecular hydrogen-bonding. This decomposition occurs at approximately 210 C. (unpublished TGA data) The impossibility of obtaining a chitin melt implies that the common melt processing routes (e.g. melt extnision and injection molding) are inaccessible for the blending of chitin with HA and the subsequent molding step. Fortunately, the discovery of non-acidic, non-degrading organic solvent systems for chitin in the 70s[3] has presented a route for the blending of chitin with HA. The solution processing route is explored and successfully applied in this work towards the manufacture of chitin-HA composites. MATERIAL S AND METHOD S Chitin was obtained from Polysciences and purified according to standard methods. N,N-dimethylacetamide (DMAc) (JT Baker) and anhydrous LiCl (BDH or Merck) were of analytical grade. Blending of chitinwithhydroxyapatite To incorporate 40% by mass of hydroxyapatite (HA) into chitin, 2.4 g of anhydrous LiCl in a 100 ml boiling tube was dried at HO^’C in a vacuum oven for at least l/2h and allowed to cool to ambient temperature. 50 ml of DMAc was added to the LiCl and the salt was dissolved with the aid of magnetic stirring. 0.24 g of HA powder was added to the resultant DMAc/ 5% LiCl solution and dispersed by stirring for 15 min. This was followed by addition of 0.24 g of purified chitin flakes and stirring was continued for a further 2 days to dissolve the chitin completely. In this way, a fine dispersion of HA particles in a 0,5% chitin solution was obtained. HA-chitin compositeplates The chitin-HA mixture was poured into molds of dimensions 22.5 cm X 5.5 cm and spread uniformly with the aid of a glass rod. Evaporation of DMAc overnight gave a chitin gel in which the HA powder was uniformly dispersed. This HA-chitin gel was washed in deionized water to remove residual DMAc and LiCl. Finally the gel was immersed in acetone, stretched and air-dried beneath glass plates to give a hydroxyapatite-chitin dry plate of dimensions approximately 12 cm x 2 cm and of cross sectional areas ranging between 0.32 and 0.51 mm^. Tensiletesting Using a blade, the specimens were trimmed to regular rectangular shape for tensile testing. The tensile tests were performed on a Shimadzu-AA series tensile tester, using a gauge length of 50 mm, cross-head speed of 5 mm/min and full scale load of 50 kgf. Abrasive paper was used to prevent breaking of the specimen at the grips. Scanning electronmicroscopy Specimens for scanning electron microscopy (SEM) were dried under vacuum overnight and gold-coated using a JFC-1100 fine coat ion sputterer. SEM micrographs were obtained using a JEOL JSM-T220A at a voltage of 15 kV.
Manufacture of a Hydroxyapatite-ChitinComposite:A.C.A. Wan et al.
513
RESULT S AND DISCUSSIO N A chilin-HA composite was successfully prepared in this work, employing a 2-stage molding process involving a solution of chitin in N,N-dimethylacetamide/ 5% LiCI. Intimate blending of the chitin solution with HA was achieved by simultaneously dispersing HA powder and chitin flakes in the solvent system. The gradually increasing viscosity of the solution due to the dissolution of chitin is believed to provide the shear stresses required to disaggregate the HA agglomerates and thus arrive at a fine dispersion of the latter. In the first stage of molding, the chitin-HA blend was cast into suitable molds in a fume cupboard. Solvent evaporation with concomitant entr>’ of water vapor led to the process of chitin gelation. As the gelation process commenced within an hour or two, the fine HA particulates remained suspended in the chitin gel to give a uniform distribution of HA in the chitin matrix. In preliminary experiments where the chitin solution was first prepared, followed by introduction of HA, inefficient dispersion of the HA particulates resulted due to the viscous nature of the chitin solution. The heavier HA agglomerates sedimented during the molding process, resulting in an asymmetric distribution of HA in the chitin gel. The chitin gel as such does not possess the mechanical properties to match that of bone, its elastic modulus being approximately 20 Mpa, in contrast to at least 50 Mpa for trabecular bone. Therefore a second stage of molding was required, which involved the conversion of the chitin-HA gel to a dried chitin-HA plate. Development of crystallinity upon drying enabled the chitin-HA material to achieve mechanical properties approaching that of hard tissue. The partially crystalline nature of the chitin plate was indicated by broad peaks at 29=8 and 19 in the XRD difl^raction spectrum. Chitin-HA gels that were allowed to air-dr\’ undenvent non-uniform shrinkage, due to different regions of the material dr>’ing at different rates. In addition, a period of at least 3 days was required for complete dr\ing. To overcome this difficulty, solvent-drying was employed. The HA. By the law of mixtures, inclusion of a brittle inorganic (HA) phase into a lower modulus polymer phase would be expected to increase the composite modulus in proportion to the volume fractions of each component. This observation can be explained by the fact that the strong hydrogen bonding interactions that exist in the chitin polymer are disrupted by inteiposing HA
1
-1 E
I 6
»
Maximu m tensile stress Tensile modulu s
|
’ ^
1 ’^.
1
\^
-\
I \ ,^
% by mas s of HA Figure 1 : Tensile propertie s of hydroxyapatite-chiti n composite s
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particles. The fact that the matrix becomes less compact with inclusion of hydro\7apatite is evident from the minimal difference in the measured densities for the chitin plate(0.12 gem") and 50% by mass HA composite plate (0.14 gcm"^), the density of hydroxyapatite (3.16 gcm’^) being much higher than that of the chitin plate. The presence of voids around the HA particulates also leads to the decreasing trend of stress at break with increase in % by mass of HA. These voids initiate crack formation during the application of stress, allowing fracture to occur at a lower load. In parallel to the observations where the density and modulus of the brittle HA filler was not proportionally represented in the overall composite properties, the % elongation at yield of the highly filled composite (50% by mass of HA) was not significantly reduced from that of unfilled chitin, both measuring at appro.ximately 14%. SUMMAR Y Solution processing is a potentially useful method for blending chitin with inorganic fillers or other polymer phases. A hydro.x7apatite-chitin composite has been achieved by this route, with physical properties that are adequate for bone substitution. No reinforcement of the chitin matrix by HA was observed, due to the natural crystallinity of unmodified chitin. However, the plasticity of chitin at high loading of HA was retained, which indicates a high fracture toughness for this composite material. ACKNOWLEDGMENT S The authors are grateful to the European Union for financial sponsorship (Contract CII*-CT94-0142). A.C.A. Wan would like to thank the National University of Singapore for a research scholarship. REFERENCE S 1. Muzzarelli, R.A.A., Biochemical significance of exogenous chitins and chitosans in animals and patients, Carhohyd Polyni, 1993, 20, 7-16. 2. Bonfield, W., Behiri, J.C, Doyle, C , Bowman, J., Abram, J. In: Biomaterials and Biomechanics^Elscw’iQr, Amsicrdmw1984, 421-426. 3. Austin, P.R In: Chitin,Chitosan and RelatedEnzymes,Academic Press Inc., Orlando FL 1984, 227-237.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
LOAD-BEARIN G AND DUCTIL E HYDROXYLAPATITE/POLYETHYLEN COMPOSITE S FOR BON E REPLACEMEN T
E
R. L. Reis^’^ A. M. Cunha^ M. J. Bevis^ ^ Dept. of Metallurgical Eng., Univ. Porto, FEUP, Rua dos Bragas, 4099 Porto Codex, PORTUGAL ^INEB - Institute for Biomedical Engineering, Pra9a Coronel Pacheco 1, 4050 Porto, PORTUGAL Dept. of Polymer Eng., U. Minho, Campus de Azurem, 4800 Guimaraes, PORTUGAL The Wolfson Centre for Materials Processing., Brunei U., Middlesex UBS 3PH, UK.
ABSTRAC T This work reports the application of non-conventional processing techniques to produce hydroxylapatite/polyethylene (HA/PE) composites. HA/PE composites were prepared with different amounts (10 to 50% by weight) of the reinforcement under two preparation procedures: bi-axial rotating drum and twin screw extrusion (TSE). These compounds were processed by conventional injection moulding and by Scorim {shear controlled orientation in injection moulding) under a wide range of processing windows. Results of tensile tests are discussed in terms of the morphology of the mouldings evaluated by polarised light microscopy and scanning electron microscopy. By optimising the processing routes and parameters it was possible to develop composites with a stiffiiess and strength in the bounds of human cortical bone, but exhibiting a very high ductility. For composites, compounded by TSE and then Scorim processed, a maximum modulus of 7.5 GPa and a strain at break as high as 19% could be obtained, for a HA weight fraction of 30%. The obtained mechanical properties (matching those of bone) combined with the bioactive behaviour of the HA phase, widen the field of application of these type of composites to orthopaedic load-bearing implants. Keywords: Hydroxylapatite; polyethylene; composites; processing; bone-analogue; mechanical properties; L INTRODUCTIO N Hydroxylapatite (HA)/polyethylene (PE) composites, which were first developed by W. Bonfield (1), are a well established bone replacement material. For the last 15 years HA/PE composites have been proposed for a range of modest load bearing applications such as cranial facial, maxillo facial and middle ear implants and orbital floor reconstruction (2-3). Following FDA approval in 1994 the material is now commercially available under the trade-name HAPEXfi (2). Recently, also the possibility of reinforcing PE with other bioactive ceramics and glasses has been assessed (4-5). However, for all these systems the reported mechanical properties, specially stiffiiess (1-7), are still much lower than typical values of human cortical bone and there remain 515
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areas of orthopaedic surgery which require bone substitute materials with considerable higher mechanical properties. Several approaches (2-8) have been used in order to improve the mechanical properties of HA/PE composites. The orientation of a semi-crystalline polymer structure can lead to a significant enhancement of the mechanical properties in specific directions, providing a mean for tailoring the anisotropy of the part (8-9). Some examples are the approaches of the groups of I. Ward (8) and M. J. Bevis (9) that have been successfully applied to several materials including polyethylene. Hydrostatic extrusion (8), has been recently used to produce load-bearing HA/PE composites with strength and stiffness in the desired range of human cortical bone and with the added benefit of superior ductility (8). The obtained results are very promising and may allow, for the first time, the use of these implants in permanent load-bearing applications. This work reports the application of non-conventional processing techniques to produce HA/PE ductile composites with a stiffiiess and strength in the bounds of human cortical bone, using a technique based on injection moulding, allowing for the production of parts with complex geometries and thick cross-sections. 2. MATERIAL S & METHOD S The studied materials were a high molecular weight polyethylene (HMWPE) grade (Hostalen GM9255F from Hoechst), with a melt flow index of 4.12g/600s (98 N, 240 ^C), which has been reinforced with hydroxylapatite (HA). The HA (from Plasma Biotal, Tideswell, UK), was sintered at 1200 C for 12 hours. After sintering and crashing in a ball mill the diameter of the HA particles was in the range of 3 to 7 fim (average of 4.5 |im), as determined by laser granulometric analysis. Composites were obtained for different HA amounts (10 to 50% by weight) under two compounding procedures: rotating drum (RD) and twin-screw extrusion (TSE). These compounds were processed by standard injection moulding and by a non-conventional injection moulding technology, into circular cross-section tensile samples (axysymmetric specimens - (^5 mm of diameter, 40 mm long). The later technique referred to as Scorim (shear controlledorientationin injectionmoulding)allows for the deliberated mimic of bone anisotropic structure by means of controlling the molecular orientation within the moulding. Two machines were used: a KlocknerFerromatik FM-20 and a Demag D-150 NCIII-K (fitted with a Scorim head), respectively. The processing conditions of the moulding program used are presented in Table I. For producing the Scorim mouldings only TSE compounded materials were used. Two optimised Scorim conditions (S2 and S3), presented in Table I, were used for processing both the polymer matrix and the developed composites. Table I - Processing conditions used in the conventional and Scorim Pinj* Holding ts^orim Condition Tjni (W [MPa] pressure [gj time (s) conventional 230^ 5 15 S2 230"" 10 35 S3 230’’ 10 35 all with the Scorim pistons oscillating out of phase Temperature profile - 180, 200, 220, 230 C
40 40
injection moulding. Packing Pj^^x. (Uni/Bi) [MPa] (Assyn/ Sync) 150 Bi/Assyn 210 Uni/Assyn 300
* pressures in the hydraulic system
Load-Bearing and Ductile HAjPE Compositesfor Bone Replacement:R.L. Reis et al.
517
The moulded samples were tensile tested in order to determine the ultimate tensile strength (UTS), the secant modulus at 1% strain (Eio/^), and the strain at break (Sr), on an Instron 4505 universal mechanical testing machine, fitted with a resistive extensometer (gauge length, 10 mm). The cross head speed was 5 mm/min until 1% strain, and then increased to 50 mm/min until fracture. The mouldings were frirther characterised optical polarised light transmission microscopy (PLM) in a Olympus BH-A microscope. Complementary characterisation of the materials included also: scanning electron microscopy (SEM) analysis of the fracture surfaces using a Leica Cambridge S360 and Jeol JSM 35-C equipments and energy dispersive spectroscopy (EDS) in a Noran Instruments device. 3. RESULT S & DISCUSSIO N Figs, la and lb summarise the most important results obtained in this study. It is noticeable the effect of the Scorim processing over the materials mechanical behaviour. In general terms, the mouldings produced with the higher compactation pressures (condition S3) evidence the best compromise between stiffness and ductility. For instance, for the HMWPE matrix it was possible to increase the stiffiiess from 1.2 (conventional moulding) to 4.5 GPa, and the UTS from 25 to 84 MPa, by means of optimising the Scorim parameters. Fig.2 presents the typical fracture surfaces of the HMWPE matrix when conventionally (Fig. 2a) or Scorim (Fig. 2b) moulded. The orientation rings, responsible for the attained mechanical performance, typical of the out of phase operation of the Scorim equipment are clear. The results obtained for composites compounded by TSE were much better, in terms of modulus than those resulting from a simple mixing in the biaxial rotating drum. The results accomplished with TSE are owe to the very good distribution of the HA reinforcement particles within the polymeric matrix, combined with the intermeshing of the polymer with the ceramic. 30
-sr 6 tt-
^
^
m^
A
.
%
"^
"
1 0 ^ ^> 0
0 0
20 HA(%) 40 A (%) ^u
60 t)U
0
. ^Sr^ ^j
_^
A
-^
^
*
20 HA(%) 40
i 60
a) Fig. 1 - a) Secant modulus at 1% strain and b) strain at break vs. HA weightfractionfor: ( ) conventionally; ( ) conventionally (TSE compounded); (A) Scorim S2 (TSE); and ( ) Scorim S3 (TSE); HA/PE moulded samples. For Scorim processed samples, a maximum modulus of 7.5 GPa (UTS= 74 MPa) could be obtained, with a ductility as high as 18.7%. This was possible for an HA weight fraction of 30%, when samples were compounded by TSE and then Scorim S3 injection moulded. The ductility of the composites is strongly increased when Scorim processing is used (Fig. lb). For all the processing conditions, and HA amounts up to 30% by weight, the modulus is favoured by the increment of the filler amount. However, higher HA fractions lead to a deterioration of this property. This should be related to the smaller degree of orientation of the HA phase when compared with the polymer matrix. Figs. 2c (10% HA) and 2d (30% HA) shows typical fracture
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surfaces of HA reinforced samples, being clear the smaller degree of orientation in Fig. 2d. This fact is the main responsible for the simultaneous decrease in the modulus and strain at break detected for 50% HA S2 and S3 processed composites. For the Scorim processed composites it has been possible to achieve a modulus matching the target value of human cortical bone with a simultaneous enhancement of sample ductility. For instance, it was possible to process PE+50% HA composites with a strain at break as high as 1518%. This performance may allow for a all new range of applications of this type of materials, on which their traditional brittle behaviour was a major constraint. The developed processing routes allow for the production of thick parts with complex geometry, that combine stiffness and strength in the range of cortical bone with an unusually high ductility (important for making these materials surgeon friendly).
Fig.2 - SE M micrograph of fracture surfaces of: a) conventionall y moulded HMWPE ; b) Scorim S3 moulded HMWPE ; c) Scorim S3 moulded 10% HA composite ; d) Scorim S3 moulded 30%H A composite.
4. CONCLUSIONS For the purpose of processing HA/PE composites the combination of Scorim technology with TSE compounding can be considered, so far, a unique way of inducing anisotropy to thick sections, with almost any desired geometry, and to produce very stiff but ductile composites that may be used in biomedical applications that have to bear important mechanical loads.
REFERENCES 1. 2. 3. 4. 5. 6. 7.
8. 9.
W. Bonfield, Biomaterials, 2, (1981), 185 M. Wang, D. Porter, W. Bonfield, Brit. Ceram. Trans., 93, (1994), 104 W. Bonfield, in Bioceramics 9, T. Kokubo, T. Nakamura, F. Miyaji, eds., Elsevier Science Ltd, Oxford, (1996), 11 M. Wang, T. Kokubo, W. Bonfield, in Bioceramics 9, T. Kokubo, T. Nakamura, F. Miyaji, eds., Elsevier Science Ltd, Oxford, (1996), 387 J. Huang, M. Wang, I. Rehman, W. Bonfield, in Bioceramics 9, T. Kokubo, T. Nakamura, F. Miyaji, eds., Elsevier Science Ltd, Oxford, (1996), 431 R. L. Reis, A.M. Cunha, S. R. Lacerda, M. H. Femandes, R. N. Correia, in Bioceramics 9, T. Kokubo, T. Nakamura, F. Miyaji, eds., Elsevier Science Ltd, Oxford, (1996), 435 A. M. Cunha, R. L. Reis, F. G. Ferreira, P. L. Granja, in Advances in Materials Science and Orthopaedic Surgery, R. Kossowsky, N. Kossowsky, eds., NATO/AS I Series (E: Applied Sciences) , Kluwer Acad. Publ., Dordrecht, (1995), 163 I. M. Ward, W. Bonfield, N. H. Ladizesky, Partnership in Polymers, Cambridge, Oct., (1996), 44 C. I. Ogbonna, G. Kalay, P. S. Allan, M. J. Bevis, J. Appl. Polym. Sci., 58, (1995), 2131
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
IN VITRO ASSESSMEN T OF HYDROX Y APATITE REINFORCE D POLYETHYLEN E COMPOSITE S
AND
BIOGLASSfi -
J. Huangl, L. Di Silvio^, M. Wangl, K.E. Tanner^ and W. Bonfieldl 1 IRC in Biomedical Materials, Queen Mary and Westfield College, Mile End Road, London El 4NS,UK 2 Institute of Orthopaedics, IRC in Biomedical Materials, Royal National Orthopaedic Hospital, Brockley HiU, Stanmore, Middlesex HA7 4LP, UK
ABSTRAC T The biocompatibility of hydroxyapatite (HA)- and Bioglassfi- particle reinforced polyethylene (HAPEX and Bioglassfi/HDPE) composites has been assessed using an in vitroprimary human osteoblast-like (HOB) cell culture model. Neither the non-surface reactive materials, Le, unfilled polyethylene and HAPEX , nor the surface reactive materials, i.e. Bioglassfi and Bioglassfi/HDPE composite, released any ’toxic leachables’ detrimental to cell viability. A ’stimulation effect* was observed in those HOB cells cultured in the extract from Bioglassfi/HDPE composite. When considering the HOB cells cultured directly on the composites, HAPEX^ showed a higher proliferation rate and better osteoblast phenotype expression. The HOB cells retained their osteoblast morphology, with cell processes embedded within the particles on both composite surfaces, indicating favourable microenvironments for cell proliferation. KEYWORD S Hydroxy^atite, Bioglassfi, composite, cell culture, cytotoxicity and SEM INTRODUCTIO N Hydroxyapatite (HA)- and Bioglassfi- particle reinforced polyethylene (HAPEX^ and Bioglassfi/HDPE) composites are being developed as second generation biomaterials. HAPEX , a bone analogue material [1], has already achieved clinical success as an orbitalfloorprosthesis [2,3] and has been launched as a middle ear implant [4]. A tough Bioglassfi/HDPE composite [5], which retains in vitrobioactivity [6,7], has been developed. In vitrocell culture models, which allow the biological assessment of materials at a cellular level, are becoming increasingly useful in the development of new biomaterials [8,9,10]. In this study, the in vitrobiocompatibility of two composites are investigated usmg a primary human osteoblast-like (HOB) cell culture model. MATERIAL S AND METHOD S The specimens of unfilled polyethylene (PE), 40 vol.% HAPEX (HAPEX) and small and large particle sized 40 vol.% Bioglassfi/HDPE composites (40S and 40L) were sectioned to squares (10 X 10 nun) from their comfffession moulded plates. The surfaces of the specimens were polished down to 1 \im (diamond paste) using an Abramin polisher (Struer, UK). The specimens were cleaned in ethanol, using an ultrasonic bath, and sterilised by gamma irradiation at a dose of 2.5 Mrad (Isotron, UK), using standard procedures for medical devices. Bioglassfi 45S5 (BG) discs 519
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(OlO mm) supplied by US Biomaterials were sterilised by dry heat at 160T for 3 hours. Thermanox (TMX) was used as a control. Primary human osteoblast-like cells (HOB) were isolated from trabecular bone fragments, obtained frcMn the femoral heads of patients undergoing hip replacement surgery [8]. The cells were cultured in Dulbecco’s Modified Eagle’s Medium QDMEM) supplemented with 10% foetal calf serum (PCS). The cytotoxicity of the test materials was determined by exposure of the cells to the aqueous extract of the materials and by direct contact of the HOB cells with materials for various time periods. Aqueous extracts were obtained from HAPEX , PE, 40L, 40S and BG by elution from test specimens in a sterile universal containing 10 ml DMEM medium, which was placed on a roller at 37"C for 24 hours. Tissue culture plastic (Tcp) and polyvinylchloride (PVC) were used as negative, non-toxic, and positive, toxic controls, respectively. Cell viability in the extracts was measured by uptake of vital dye, such as Neutral Red, by viable cells. HOB cells (approximately 1 x 10^ cells ml"^) were carefully seeded on each specimen and allowed to attach for one hour prior to flooding with 1 ml DMEM medium. The cultures were incubated at 37"C in humidified atmosphere in the presence of 5% CO2 for various time. The culture medium was carefully changed, in order to minimise disturbance of the culture conditions, at appropriate time intervals. At each time point, the metabolic activity of the cells on the materials was assessed with the aid of a REEJOX indicator (employed in Alamar blue), which measures the response to chemical reduction of growth medium resultingfromcell growth. The cell growth and proliferation on test materials were measured by total DNA content and tritiated thymidine [ H]-TdR incwporation. The HOB cells were incubated in the presence of 1 ^iCi ml"l of [ H]-TdR (Amersham International, UK) for the final 16 hours of culture. The cells were enzymatically lysed in a papain digest solution. Total cellular DNA was measured using the Hoechst method [11]. [ H]-TdR incorporation was measured by trichloroacetic acid precipitation of the cell digest and the amount ofradiolabelincorporated was measured using a scintillation counter. Production of alkaline phosphatase (ALP), an osteoblast phenotype marker, was determined biochemically using a COBAS-BIO (Roche, UK) analyser. P-nitro phenol phosphate in diethanolamine buffer (Merck, UK) was used as the substrate. Test specimens for examination of cell-material interaction were seeded at a density of approximately 8 x 10^ HOB cells ml-^ and incubated at 37T in a humidified air with 5% CO2. After 24 hours incubation, the cultures were fixed with 2.5% glutaraldehyde buffered in 0.1 M sodium cacodylate, stained with 1% osmium tetroxide and 1% tannic acid buffer, dehydrated in ascending grades of alcohol solutions (20 to 100%) and then in hexamethyl-disilazane (Sigma, UK), and finally air dried overnight. The cultures were coated with a thin layer of gold before examination under a JOEL scanning electron microscope. RESULT S AND DISCUSSIO N No evidence of cytotoxicity was observed following incubation of cells in the medium eluted from any test material over a period of 24 hours, indicating that no toxic teachable was released. An increase in cellular activity was observed with Bioglassfi/HDPE composites 408 and 40L (Figure la) when compared to Tcp. The stimulation effect observed was probably related to the release of Si from the Bioglassfi into the extracts [12]. The metabolic activity of the cells was not affected when cultured in direct contact with the materials. An increase in activity with time was observed (Figure lb). The proliferation of HOB cells expressed as the total DNA content increased with time on BG, PE and HAPEX . A higher rate of [^H]-TdR incorporation/DNA was seen on day 1, where the greatest proliferation of HOB cells was found on HAPEX (Figure 2a). The total ALP activity increased with time over the period studied for test materials, with the exception of composites 40S and 40L. The ALP activity increased considerably on HAPEX at day 7, which is indicative of osteoblast differentiation (Figure 2b).
In Vitro Assessmentof Hydroxyapatite-and Bioglass^: J. Huang et al.
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(b) ^
12t. t
i 1 Test Materials
4 7 Incubation time (day)
Figure 1. (a) Viability of HOB cells following 24 hours exposure to 24 hour eluted media from test and control materials, statisticall y significant difference s from Tcp using Student’s t-test indicated by * p < 0.05 or ** p < 0.01. (b) Metabolic activity of HOB cells on test materials from day 1 to in Alamar Blue assay. 14 measured fromfluorescence HO B cells were able to attach to all the test materials and generally maintained their osteoblas t morphology. Numerous HOB cells with processe s were seen covering the surface of TM X control after 24 hours culture. A complete cell covering was observed on BG , with cells having visible filapodia, while relatively fewer cells were observed on the surface of PE, where the cells appeared in ’patches’. On HAPEX , a compact layer of cells was observed covering the entire surface and maintained the polygonal morphology of osteoblasts , (Figure 3a), HOB cells spearedflattened with filapodia attache d to HA particles (Figure 3b). Cell processe s embedde d within Bioglassfi particles were observed on Bioglassfi/HDPE composite 40S (Figure 3c), while cells at various stages of division were often observed on 40L (Figure 3d). CONCLUSION S From the results of the cytotoxicit y test, biochemica l evaluation and morphological assessment , both composite s were biocompatibl e with the bone cells and encourage d cell attachmen t and growth. Therefore , it can be conclude d that these composite s are promising implant materials for a variety of applications . ()
TM X HAPE X
1
0 S
408 40L
jte^tii^Lj
4 Incubation time (day)
Incubation time (day)
Figure 2. (a) [^H]-TdR incorporatio n by HOB cells and (b) AL P activity of HOB cells cultured on test materials: TM X control, HAPEX^^ , Bioglassfi/HDPE composite s 40S and 40L from day 1 to day?.
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(a)
(0
(d)
Figure 3. Morphology of HOB cells on (a) and (b) HAPEX 40Sand(d)40L.
and Bioglassfi/HDPE composites (c)
ACKNOWLEDGEMENT S The financial support from US Biomaterials for this project and the continuing support of EPSRC for the IRC core programme are gratefully acknowledged, together with help from Mr M. Kayser, Mrs C. Clifford and Mrs N. Gurav. REFFRENC E 1. Bonfield, W., J. Biomed.Eng. 1988,10, 522-536. 2. Downes, R.N., Vardy, S., Tanner, K.E. and Bonfield, W, Bioceramics4, 1991, 239-246. 3. Tanner, K.E., Downes, R.N., Bonfield, W., British Ceram. Trans.,1994,93, 104-107. 4. Bonfield, W., Bioceramics9,1996,11-13. 5. Wang, M., Bonfield, W. and Hench, L. L., Bioceramics8,1995. 383-388. 6. Huang, J., Wang, M., Rehman, I., Knowles, J. and Bonfield, W., Bioceramics8,1995, 389395. 7. Huang, J., Wang, M., Rehman, I., and Bonfield, W., Bioceramics9, 1996,431-434. 8. Di Silvio, L., PhD thesis. University of London, 1995 9. Vrouwenvelder, W.C.A., Groot, C.G and de Groot, K., /. Biomed.Mater. Res.,1993, 27, 465-475. 10. Bagambisa, F.B. and Joos, U., Biomaterials,1990,11, 50-56. 11. Rago, R., Mitchen, J. and Wilding, G., Anal.Biochem..1990,191, 31-34. 12. Hench, L.L., Bioceramics7, 1994, 3-14.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
OSTEOCONDUCTIV E HYDROX Y APATITE S
PROPERTIE S
OF
PUR E
AND
TYPE- A
CARBONATE D
S.A. Redey^ D. Bemache-Assolant^ C. Rey\ P.J. Marie^ M. Nardin^ and L. Sedel^ ^Institut de Chimie des Surfaces et Interfaces, 15 Rue Jean Starcky, 68057 Mulhouse, France ^LMCTS, CNRS UPRESA 6015, 123 Avenue A. Thomas, 87060 Limoges Cedex, France ’UPRESA 5071, E.N.S. de Chimie, INPT, 38 Rue des 36 Ponts, 31400 Toulouse, France ’^INSERM U349, Hopital Lariboisiere, 6 Rue Guy Patin, 75475 Paris Cedex 10, France ^Faculte de Medecine Lariboisiere-St-Louis, 10 Avenue de Verdun, 75010 Paris, France
ABSTRAC T Adhesion, proliferation and differentiation of human trabecular (HT) osteoblastic cells, as well as changes in wettability and surface chemistry, were investigated on synthetic dense hydroxyapatite (HA) and carbonated HA surfaces (CHA). HT cell behaviour was evaluated on HA and CHA by determining cell attachment after 18 hours, cell proliferation up to 28 days as well as extracellular matrix (ECM) production. Whereas cell proliferation was similar on the two substrates, HT cells showed poor adherence to CHA where they gathered in clusters since, due to an important drop of polar surface energy of HA after carbonation, polar interactions with the surface were not favoured. However, newly formed apatite production per cell was found to be higher on CHA using Fourier-Transform Infi-aRed microscopy. Full chemical analysis of the surfaces were completed by X-ray Photoelectron Spectroscopy. These results show that HA surface carbonation induces significant alterations in physico-chemical properties, osteoblast fiinction and ECM formation on HA. KEYWORD S Adhesion - Carbonation - Extracellular matrix - Hydroxyapatite - Osteoblastic cells
INTRODUCTIO N When unplanted, a biomaterial can be considered as osteoconductive when it provides rapid proliferation and differentiation of osteoblastic cells at its surface, leading to the synthesis of an interphase of mineralised collagenous bone matrix between bone tissue and bulk implant. This can only be achieved if cells adhere properly to the substrate; thus, interfacial interactions between proteins (including cell integrins) and bioactive material appear to be key parameters of osteoconductivity. Since synthetic resorbable ceramics were presented as possible bone substitutes [1], there has been a growing interest in the potential use of hydroxyapatite (HA) in orthopaedic surgery and dentistry. Carbonated hydroxyapatite (CHA), with a chemical composition closer to bone or dental enamel than HA [2], is thought to be a promising alternative since it is well-established that all bioactive materials develop a layer of biologically active CHA [3]. 523
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In this work, after the initial characterisation of synthetic dense HA and CHA surfaces, human trabecular (HT) osteoblastic cell adhesion, proliferation and differentiation were studied on the two substrates with standard culture plastic as a reference. Mineralisation of the surfaces was evaluated by both biological and physico-chemical techniques. MATERIAL S AND METHOD S HA powder was synthesised with solutions of ammonium phosphate and calcium nitrate using a double decomposition method [4]. The powder was compressed into small pellets in a steel die under a pressure of 150 MPa with a subsequent sintering at 1200 C during 2 hours. HA pellets, with a relative density of 0.94, were finally polished with SiC paper (grade 4000). Carbonation of HA surfaces was performed by heating HA for 3 days at 900 C in a dry COj-saturated atmosphere. Type-A carbonation was identified by diffiise refiectance infrared spectroscopy. Surface energy determinatio n Surface energies were determined before cell culture by wettability [5] using a Kriiss G2 Contact Angle Measuring System. Measurements with tricresylphosphate and a-bromonaphthalene allowed the calculation of the dispersive component of the surface energy, whereas the polar interaction energy was evaluated with bi-distilled water. H T osteoblasti c cell culture Normal HT cells were isolated from adult trabecular bone as previously described [6]. Cells were cultured in DMEM supplemented with 100 Ul/ml penicillin, 100 |Lig/ml streptomycin and 10% foetal calf serum (FCS) and incubated at 37 C in a humidified atmosphere [6]. Cells were detached by trypsinisation then plated at 25000 cells/cm^ in a minimal volume (200 |LI1) on HA or CHA pellets or culture plastic dishes. The culture medium was DMEM with 10% FCS supplemented with 50 fig/ml ascorbic acid and 1.5 mM p-glycerol phosphate. Cell attachment: After 18 hours incubation at 37 C in a humidified atmosphere, culture surfaces were rinsed with PBS, fixed with paraformaldehyde (PEA) and stained with toluidine blue. Number of attached cells was determined by optical microscopy. CeJl proliferation: DNA synthesis was evaluated after addition of (^H)-thymidine to the medium 4 hours before the end of the culture [6]. Furthermore, cells were counted with an haematocytometer after they were detached by trypsinisation. Cell differentiation: Matrix synthesis by HT cells was determined after addition of (^H)-proline to the medium 24 hours before the end of the culture. Type I carboxyterminal peptide (PICP) release into the medium was evaluated using a specific radioimmunoassay, enabling the determination of collagen synthesis. Histology was performed after sample embedding in methyl methacrylate. Surface characterisatio n X-ray Photoelectron Spectroscopy (XPS) and Fourier-Transform Infrared spectroscopy (FTIR) were performed on HA and CHA surfaces before and after cell culture, cells being fixed with PEA and stained with toluidine blue. XPS analysis was performed with a Leybold LHSl 1 spectrometer. FTIR analysis was performed using a Bruker IFS66 analyser coupled to an Infrared Microscope. Data were collected on apatite surfaces with objective aperture ranging from lOOjim to 400 jim. Samples, previously coated by cathodic gold-sputtering were also examined by scanning electron microscopy.
OsteoconductivePropertiesof Pure and Type-A CarbonatedHydroxy apatites:S.A. Redey et al.
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RESULTS AND DISCUSSION Surface energies of HA and CHA before cell culture are reported in Table 1. Since the polar interaction energy of HA with water is significantly decreased after carbonation, interactions between CHA and biological species are not favoured according to wettability criteria. This was confirmed by cell attachment measurements, showing that the number of adherent cells was less than twice lower on CHA compared to HA (Fig.l). If similar cell proliferation were observed on the two substrates, HT cells spread over the whole HA surface (one to a few layers) whereas they gather in localised clusters on CHA, as confirmed by optical microscopy observation. (^H)-proline incorporation showed that synthesis of ECM was much higher on HA than on CHA (Fig.l). This was confirmed by histological analysis. Newly formed apatite on HA and CHA after cell culture was evaluated by XPS and FTIR microscopy. The surface analysed was large (6.8 mm^) with XPS and small (<0.13mm^) with FTIR. XPS results show that C/P, 0/P and N/P ratios increased with cell density on each substrate and reached higher values for HA (Table 2), which confirms HT cell culture results previously mentioned. Mineralisation after cell culture was observed on HA and CHA FTIR spectra with relative intensities of carbonate peaks at 1540cm"’, 1455cm"’ and 1420cm"’ [7] (see Fig. 2). Apatite formation increased with cell density and was locally higher on CHA than on HA at comparable cell density. However, protein adsorption was apparently higher on HA than on CHA (peaks at 1608cm"’ and 1578cm"’). This suggests that osteoblasts are more active on CHA than on HA, although the lower initial cell attachment on CHA make HA a more osteoconductive support in these experunental conditions. Cell number is hence a predominant factor in bone formation, as previously reported [8].
rs" (mJ/m^) |HA |CHA
34.5 – 3 28 –5
/^(mJ/m=’) 44 – 2
9– 5 1
Table 1: Surface energies: dispersive compo› nents and polar interactions with water.
Number of adherent cells after 18 hours 10000 8000 I
LJI _
0/P
N/P
13.2 7.1
5.3 3.4
0.7 1
1.6
Table 2: XPS analysis of HA and CHA after HT cell culture.
Differentiation after 5 weeks: (3H)-proline incorporation in the matrix 20000
15000 I 10000 I 5000 I
6000 I 4000 2000 i 0
HA
IHA |CHA
C/P
0 CH A
IKIiiij HA
CH A
Figure 1: HT cell adherence and differentiation on HA and CHA. (* = significant results)
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^
0 . 8. 0.7 - mmmmmmmmm 0.6- iiii!ii!iiiiiiili!iiB^^^^^ ^ 0 .5
CH A
, . TT A
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1 0.4 -
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Figure 2: FTIR spectra of HA and CHA after HT cell culture (internal references)
Figure 3: SEM photograph of newly formed apatite on HA after HT cell culture
Newly formed apatite crystals exhibit a specific "rosace"-shape as shown on Fig. 3; they suggest a crystallisation along an axis perpendicular to the initial hydroxyapatite surface.
CONCLUSIO N Surface carbonation considerably modifies physico-chemical and biological properties of hydroxyapatite pellets. Whereas human osteoblastic cell proliferation was similar on both HA and CHA substrates, much less extracellular matrix was produced by the HT cells on CHA, although mineral production per cell was found to be greater. The critical parameter appears to be the initial cell attachment, the latter being driven not only by specific chemical recognition sites at the surface, but also by surface thermodynamics. ACKNOWLEDGMENT S This work was supported by the Centre National de la Recherche Scientifique (DIMAT) and the Assistance Publique / Hopitaux de Paris ("Biomateriaux Inorganiques Osteoconducteurs"). REFERENCE S 1. Graves, G.A. Jr, Hentrich, R.L. Jr, Stein, H.G. and Bajpai, P.K. In: Engineering in Medicine, part I, Interscience Publishers, New York 1972, 91-115. 2. Legros, R., Bahnain, N.and Bonel, G. J. Chem.Res. Synop. 1986,1, 8-9. 3. Hench, L.L. Chemistryand Industry1995, 7/17, 547-550. 4. Bonel, G., Heughebaert, J.C, Heughebaert, M., Lacout, J.L. and Lebugle, Ann. N. Y. Acad. Sci. 1988,523,115-130. 5. Schultz, J. and Carre, A. In: Macromolecules, 27^^ International Symposium on Macromolecules,Pergamon Press, Oxford 1981, 289. 6. Marie, P.J., Lomri, A., Sabbagh, A. and Basle, M. In VitroCell.Develop.Biol. 1989, 25, 373380 7. Regnier, P., Lasaga, A.C., Bemer, R.A., Han, O.H. and Zilm K.W. American Mineralogist 1994,79,809-818 8. Marie, P.J. Calcif. TissueInt. 1995, 56, S13-16
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
COAGULATIO N TIME S OF BLOO D IN CONTAC T WIT H GEL-DERIVE D SILICA ALUMIN A COMPOSIT E POWDERS Seisuke Takashimal, Chikara Ohtsuki^, Satoshi Hayakawa^ and Akiyoshi Osaka^ 1 Co-operative Research Center, Okayama University, 5,302, Haga, Okayama 701-12, Japan 2 Biomaterials Lab., Faculty of Engineering, Okayama University, 3-1-1, Tsushima-Naka, Okayama, 700, Japan ABSTRAC T Blood compatibility of gel-derived silica-alumina composite powders in varied mixing molar ratios were examined by measuring 3 coagulation parameters (PTT, ¥Y and Fib) when contacted with fresh healthy serum. It was observed that the compatibility depended on the mixing ratio and the calcining temperature, although silica-alumina generally showed poor blood compatibihty. It was recognized that blend of caldimi phosphate into the silica-alumina composite (Si02/Al203=5/5, calcium phosphate concentration: 5-10 wt%, Ca/P=1.68. calcined temp.: 600 C) was improved the parameters of blood compatibility to the practical level. INTRODUCTIO N In present time, the extracorporeal blood purification therapies by membranes and adsorbents are effective for the incurable diseases[l-5]. In these therapies, it is recognized that the adsorption therapy is more effective than that by membrane because of more specific removal of the pathogenic substances in patient serum. It is essential that the adsorbents should have not only specific activity against the pathogenic substances but also the blood compatibility. In this study, the authors examined the blood compatibihty of gel-derived siUca-alumina(SiAl) composite powders with varied mixing ratios. In addition, the effect of blend of calcium phosphate(CaP) in the Si-Al composite gels on the blood compatibihty was investigated in order to find optimum conditions for the blood compatibility of this type of composite. MATERIAL S AND METHOD S 1) Preparation of gel derived Si-Al composite powders The mixtures of silica and alumina colloidal gels (Nissan Chemicals, Tokyo) were prepared in varied molar ratios(Si02/Al203=R: 1/9 to 9/1) and subsequent by calcined at 400 to 1,000 C 2) Preparation of Si-Al gels containing CaP(Si-Al-CaP) An (NH4)2HP04 aqueous solution of pH 10 and the alumina colloidal gel were mixed into the silica colloidal gel containing Ca(N03)2*4H20 of pH 10 with a composition of R=5/5 and Ca/P=1.68 by the peristaltic pump at the same time. The hquid phase of the mixture was removed and the powders obtained were subsequendy calcined at 400 to 1,000 C. 3) Measurement of PTT, PT and Fib of serum Blood coagulation parameters, i.e. active partial thromboplastin time(PTT(sec.)), prothrombin time(PT(sec.)) and amounts of Fibrin (Fib(mg)) were measured when the composite powders prepared contacted with healthy blood serum, by using KoaguLab MJ. Ortho Chnical 527
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Diagonotic Co. Ltd. After the prepared samples were pulverized under 300 [Am, 0.3g of the obtained powder (i.e. absorbent) was immersed in 1.0ml of saline and degased at 120 C for 20 mins. After it was cooled, 2.0ml a mixture of fresh serum extracted from 10 adult volunteers was added to the adsorbent and stirred in the glass tube. After 3hrs, the serum was separated from adsorbent by cellulose acetate membrane with pore size of 0.45^m. PTT, PT and Fib of the treated serum were measured by KoaguLab MJ. In the measur^nent, fresh mixed serum(2.0ml of serum + 1.0ml of saUne) was used as a negative control, and 0.79% poly hydroxy methyl methacrylate coated active carbon(PHEMA-AC, DHP-1, Kuraray Co., Osaka) as a positive control. 4) Characterization of the composite powders (adsorbents) The composite powders prepared were characterized by an X-ray diffractometer and infrared spectrometer. Add/Base titration using O.OlN-diethylamine aqueous solution and ’-0.02N-Ha aqueous solution with phenolphthalein(as indicator) was carried out to determine electric charges of adsorbents surfaces. RESULT S AND DISCUSSIO N 1) Blood compatibility of Si-Al composites Figure 1 shows the blood compatibility parameters ( PTT(a), PT(b) and Fib(c)) plotted against the molar ratio for the Si-Al composites calcined at 400, 600, 1,000 C as well as asprepared one. On PTT, all observed values were largely deviated from the control. The clotting time abnormally prolonged is attributed to decrease in amounts of thromboplastin in the serum. On PT, it was observed that PT affected by the calcining temperature in the lower Si02/Al203(=R) range, and that PT seemed to be gathered to the restrict zone of 40’-60s in the higher R range. Trend of change in Fib values was similar in line shape to that of PT in higher R range. These experimental results showed that the Si-Al composites showed poor blood compatibility, although the parameters varied with R ratios. Then, an attempt has been carried out to improve the blood compatibiUty of the Si-Al composites. The authors previously reported that calcium phosphate(CaP) powder shows high blood compatibility and that it had the specific adsorption activity against P2~ ^oglo^^Jliti[^l» which was a pathogenic substance to amyloidosis of
3/7
5/5
7/3
SI02/AI203 ratio
3/7
5/5
7/3
Si02//M203 ratio
3/7
5/5
7/3
SI02/AI203 ratio
Figure 1 Clotting factors(PTT, PT and Fib) of the gel-derived Si-Al composites with various Si02/Al203 ratios heat-treated at various temperatures. ; As-prepared V ; 400 C D ; 600 C O ; 1000 C Blank PHEMA-AC
Coagulation Times of Blood With Gel-DerivedSilica-Alumina CompositePowders: S. Takashima et al. 170,
529
(c ) [Fib|
160
^^ |,150
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t T
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7
CaP conten t (wt%)
CaP conten t (wt%)
.
1 . 3
1 . 5
1 . 7
1 .
i
Cap conten t (wt%)
Figure 2 Clottiiig factors (FrT(a), PT(b) and Fib(c)) of Si-Al-CaP heat-treate d at 600C containing various CaP content (wt%) in fresh human blood plasma Blank PHEMA-A C chronical renal failure patients [7]. According to this finding, the highly dispersed Si-AlCaP(R=5/5) adsorbents in varied concentration s were prepared, and the blood compatibiUty of these adsorbents was measured. Figure 2 shows the parameter s of blood compatibilit y in this experimenta l series. PTT values were close to the control at 7-’10wt% er"f CaP (60’-80s). PT was also close to that of the controls at 5-7 wt% (18-’24s) and Fib considerabl y ^proached that of the controls in 5-10 wt% (152’-16 2 mg). However, the serums were coagulate d when contacte d with the adsorbents at higher concentration s of CaP(15, 30 wt%). For references , the blood compatibiUty of absorbents prepared by mechanica l blend of hydoryxapatite(HAp ) powder to SiAl were examined and the results were shown in Figure 3. It was confirmed that the blood compatibility of the mechanica l blended adsorbents was less than that of Si-Al-CaP which might have highly dispersed calcium phosphate . It was suggeste d that the blood compatiWHty of the adsorbent was depende d on the preparation conditicms including R, the calcining temperature , the CaP concentration , Ca/P ratio and the dispersive degree of CaP in the Si-Al composite powder.
0
3
5
7
HAp conten t (wt%)
3
5
7
HAp conten t (wt%)
3
5
7
HAp conten t (wl%)
Figure 3 Qotting factors (PTT(a), PT(b) and Fib(c)) of mechanical y blend containing various HA p content (wt%) in fresh human blood plasma Blank PHEMA-A C
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10 0
100
90 801 70 E
60E -
O X
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Si02/A|203 ratio
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30
Cap conten t (wt%)
Fig. 4 Amounts of dropped 0.02N-Ha to diethylamine versus varied Si02/Al203 ratio and varied content of calcium phosphate (wt%). The gel-derived composites. ; As-prepared V ; Treated at 400 C D ; Treated at 600 C O ; Treated at 1000 C Blank PHEMA-AC 2) Characterization of the adsorbent It is important to understand the reascm why the Si-Al composite powders with highly dispersed CaP have the good blood compatibility. Therefore some attempts to clarify the relationship of the blood compatibility and the characteristics on these adsorbents have been carried out The calcined Si-Al(R=5/5) sample did not show any significant change in the crystal phase determined by X-ray diffraction pattern, except for the slight change of line sh^)e contributed to OH (ca. l,100cm-l) on the infrared spectra. No remari^able change was also detected cm both spectra of Si-Al-CaP. On the other hand, the line shq)e of the titration curves for the Si-Al composites with varied concentrations shows similar traids to the Fib curves in Fig. 1, whereas the electric charges on the surface of Si-Al-CaP composites had a constant value even if the CaP concentration is varied, as shown on Figure 4 (a) and (b). It was guessed that Fib behavior was related the electric charge of the adsorbent surface. CONCLUSION S 1. The blood compatibiUty of gel derived Si-Al composite powder was improved by adding CaP synthesized in lower ccmcentration in Si-Al gel phase. 2. It is guessed from the line shq)es of titration curves that Fib behavior is well correlated the electric charge on the surface of Si-Al absorbent. REFERENC E 1. Saibara, T. etal,Japanese J. Artific.Organs, 1989,18,378-389. 2. Shiozaki, S. etaL, JapaneseJ. Artific.Organs. 1987,16,1037-1(>40. 3. Tani, N. JapaneseJ. Medic. Instrument,1988,58,266-273. 4. Tanihara, M. etal.,JapaneseJ. Artific.Organs, 1989,18,15-18. 5. Agishi, T. etai, JapaneseJ. Artific.Organs. 1991,20,318-323. 6. Takashima, S., etal..In: Bioceramics.,Volume 9, Pergamon, Oxford, 217-220. 7. Kitano, Y. etai, JapaneseJ. Artific.Organs, 1988,20,1491-1496.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
PREPARATIO N OF P^-IMPLANTE D YjOj-AljOj-SiO j GLAS S FOR RADIOTHERAP Y OF CANCE R M. Kawashita\ F. Miyaji\ T. Kokubo\ G. H. Takaoka^ I. Yamada^ Y. Suzuki^ and M. Inoue^ ^Department of Material Chemistry, Faculty of Engineering, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan, ^lon Beam Engineering Experimental Laboratory, Faculty of Engineering, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan, ^lon Engineering Research Institute Corporation, Hirakata, Osaka 573-01, Japan ABSTRAC T P^ ion was implanted into a Y203-Al203-Si02 glass, which has already been used for radiotherapy, at 200 keV with a dose of 1x10^^ cm’^ in order to increase its treatment effect. The phosphorus was distributed up to the glass surface and a part of it near the surface was oxidized. The P^-implanted Y203-Al203-Si02 glass released appreciable amounts of P, Y and Si into water at 95 C for 7 d, although the P^-implanted silica glass hardly released these elements under the same condition. This result indicates that P^ ion must be implanted with lower doses than Ix 10^^ cm"^ or at higher implantation energies than 200 keV in order to obtain highly chemically durable P-containing Y203-Al203-Si02 glass, preventing the distribution of phosphorus to the surface layer. KEYWORDS : Phosphorus, Ion implantation, Y203-Al203-Si02 glass. Radiotherapy INTRODUCTIO N Radiotherapy is one of the effective treatments of cancers. External irradiation, however, often causes damages to healthy tissues. It has been reported that a 17Y203-19Al203-64Si02 (mol%) glass is usefiil for in situirradiation of cancers [1]. Y-89 in the glass can be activated to P-emitter Y-90 with 64.1 h half-life by neutron bombardment. Microspheres of the activated glass can give large local irradiation of the short-ranged highly ionizing p-ray to the tumors with little radiation dose of neighboring organs, when they are injected to the tumors. Y-90, however, may result in the substantial decay before the treatment owing to short half-life of 64.1 h. P-31 with 100% natural abundance similar to Y-89 can be activated to pemitter P-32 with 14.3 d half-life by neutron bombardment. But highly phosphorus-containing glasses prepared by the conventional melting method are usually less chemically durable. It can be expected that a glass, which is more effective for radiotherapy, can be obtained by P^ ion implantation into Y203-Al203-Si02 glass. The present authors previously showed that P^ ion can be successfiilly implanted into a silica glass under 200 keV with a dose as high as 1x10^^ cm"^ without giving adverse effect on its high chemical durability [2-4], In the present study, P^ ion was implanted into 17Y203-19Al203-64Si02 (mol%) glass with a dose of 1x10^^ cm"^ under 200 keV in order to obtain the glass with higher treatment effect of radiotherapy. The state of the implanted phosphorus, and structural change of the glass surface due to ion implantation were examined. Chemical durability of the P^-implanted 531
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glass was discussed in terms of its surface structure. MATERIAL S AND METHOD S A glass batch of 17Y203-19Al203-64Si02 (mol%) composition consisting of reagent-grade chemicals of Y2O3, AI2O3 and Si02 was melted in a platinum crucible at 1600 C for 2 h, poured on a stainless steel plate to be formed into plates 1 mm thick. The obtained glass was cut into rectangular specimens 10x10x1 mm^ in size, polished with 3-4 |im diamond paste, washed with pure acetone in an ultrasonic cleaner. The glass plate was annealed at 850 C for 1 h in order to eliminate the strains. The glass was implanted with P^ ion at 200 keV with a dose of 1x10^^ cm’^. The distribution of phosphorus was measured by a Rutherford backscattering spectrometry (RBS) using 2 MeV "^He^ ions with 170 incident angle. The state of phosphorus in the glass was investigated by measuring survey spectrum with an X-ray photoelectron spectroscope (XPS) (MT-5500, ULVAC-PHI Co. Ltd., Chigasaki, Japan), using MgKa X-ray as the source. The P^-implanted glass was soaked in 20 ml of distilled water at 95 C for 7 d in a polypropylene bottle, shaken at a rate of 120 strokesmin’^ with a stroke length of 3 cm. The concentrations of the phosphorus and silicon released from the glass were measured by an inductively coupled plasma atomic emission spectrometer (SPS-1500 VR, Seiko Instruments Inc., Tokyo, Japan). RESULT S AND DISCUSSIO N Figure 1 shows the RBS spectrum of P^-implanted Y203"Al203-Si02 glass, in comparison with that of unimplanted original glass. Judging from a broad peak at 200-300 in channel number, it is assiuned that P^ ion was successfully implanted into the glass, although the peak of phosphorus overlapped those of aluminum and silicon. A rising was observed at about 400 in channel number for PMmplanted Y203-Al203-Si02 glass. This indicates that yttrium ions moved to the glass surface during the ion implantation. Figure 2 shows the P2p XPS spectra of P^-implanted Y203-Al203-Si02 glass. Two peaks.
P’^-implanted
Unimplanted Y
1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1
200
300
400
Channel nunnbe r
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Figure 1 RBS spectrum of P^-implanted Y203-Al203-Si02 glass, in comparison with that of unimplanted original glass.
134 130 Binding energy /eV
Figure 2 P2p XPS spectra of P^-implanted Y203-Al203-Si02 glass.
P^-Implanted Y2O3 -AI2O3 -Si02 Glass for Radiotherapyof Cancer: M. Kawashita et al.
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assigned to elementa l phosphorus (130 eV) and oxidized phosphorus (134 eV), were observed at the surface. With increasing depth, the intensity of the peak ascribed to the elementa l phosphorus increase d while that ascribed to the oxidized phosphorus decreased . This means that a part of phosphorus exists as oxidized one near the surface, although most of it exists as elementa l colloids in the glass. Figure 3 shows the concentration s of phosphorus, silicon and yttrium release d from the PMmplanted Y203-Al203-Si0 2 glass into water at 95 C for 7 d. The glass release d appreciable amounts of P, Y and Si into the hot water. The appreciable release of yttrium and phosphorus may be attribute d to the surface localizatio n of yttrium (Fig. 1) and the formation of chemically less durable phosphorus oxide near the glass surface (Fig. 2), respectively . The release s of larger amount of the yttrium and the phosphorus might enhance the release of silicon. Figure 4 shows the RBS spectra of PMmplanted Y203-Al203-Si0 2 glasses before and after soaking in water at 95 C for 7 d. Peak area ascribed to phosphorus remarkably decrease d after the soaking in hot water. This result indicates that the implanted phosphorus was release d into water by the soaking, which is consisten t with the result shown in Fig. 3. For silica glass implanted with P^ ion at 200 keV with a dose of 1x10^^ cm’^ [4], the implanted phosphorus was localized only in deepe r regions, and hence the glass showed high chemical durability. For P^-implanted Y203-Al203-Si0 2 glass, the implanted phosphorus can not penetrat e into the deep region of the glass, because it contains heavy Y-89. Consequently , the phosphorus, which was implanted into Y203-Al203-Si0 2 glass, was oxidized to form chemically less durable phosphorus oxide at the glass surface and the glass showed lower chemical durability compared with P^-implanted silica glass. This means that the implanted phosphorus must be localized in deep region for keeping high chemical durability of the glass even after the P^ ion implantation . It is therefor e suggeste d that P^ ion must be implanted with lower doses than Ix 10^^ cm"^ or at higher implantation energies than 200 keV in order to obtain highly chemically durable P-containing Y203-Al203-Si0 2 glass, preventin g the accumulatio n of phosphorus in the surface layer.
Unim planted ^ P -implante d Unlmplante d d ^ ^ ^^ P’^-implante Unimplante d ^ P’^-implante d 0
0.5
1
1.5
Concentratio n /pp m
2
Figure 3 Concentration s of P, Y and Si release d from P^-implanted Y203-Al203-Si0 2 glass soaked in water at 95 C for 7 d.
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200
300
400
Channel number
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Figure 4 RBS spectra of P^-implanted Y203-Al203-Si02 glasses before and after soaking in water at 95 C for 7 d. CONCLUSION S Y203-Al203-Si02 glass was implanted with P^ ion at 200 keV with a dose of 1x10^^ cm ^ The P^-implanted Y203-Al203-Si02 glass released appreciable amounts of P, Y and Si, although the P^-implanted silica glass hardly released these elements under the same condition. This is because the implanted phosphorus was widely distributed up to the surface and a part of it near the surface was oxidized in the former glass, while it was localized in deeper region in the latter glass. P^ ion must be implanted with lower doses than 1x10^^ cm"^ or at higher implantation energies than 200 keV in order to obtain highly chemically durable P-containing Y2O3-AI2O3Si02 glass, preventing the distribution of phosphorus to the surface layer. ACKNOWLEDGMENT S We thank Radiation Laboratory of Nuclear Engineering, Kyoto University, for RBS measurement. This work was supported by Grant-in-Aid for Young Scientists from Ministry of Education, Science, Sports and Culture, Japan.
REFERENCES L
Ehrhardt, G.J. and Day, D.E., NucL Med Biol.1987,14, 233-242.
2. Kawashita, M., Miyaji, F., Kokubo, T., Takaoka, G.H. and Yamada, I., J.Ceram.Soc, Jpn. 1996, 104, 710-714. Kawashita, M., Miyaji, F., Kokubo, T., Takaoka, G.H., Yamada, I., Suzuki, Y. and Kajiyama, K., Nucl Inst Meth.in Phys. Res. B, 1997, 121, 323-327. Kawashita, M., Miyaji, F., Kokubo, T., Takaoka, G.H. and Yamada, I., In: Proc. 2ndlntl Meet.PacificRim Ceram. Soc, The Austalasian Ceram. Soc., 1996, in press.
Bioceramics, Volume 10 Edited by L, Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
NEW FERROMAGNETIC BONE CEMENT FOR LOCAL HYPERTHERMIA K. Takegami*, T. Sano*, H. Wakabayashi*, J. Sonoda*, T. Yamazaki*, S. Morita**, T. Shibuya**, A. Uchida* *Department of Orthopedic Surgery, Mie University Faculty of Medicine, Edobashi 2-174, Tsu, Mie 514, Japan **Nippon Electric Glass Co., Ltd. ABSTRAC T We have developed a ferromagnetic bone cement as thermoseed to generate heat by hysteresis loss. This material resembles bioactive bone cement in composition, with a portion of the bioactive glass ceramic component replaced by magnetite powder. The temperature of this thermoseed rises in proportion to the weight ratio of magnetite powder, the volume of the thermoseed, and the intensity of the magnetic field. The heat-generating ability of this thermoseed implanted into rabbit tibiae was investigated by applying an alternate magnetic field In this system, it is very easy to increase the temperature of the thermoseed in bone beyond 50X1 by adjusting the above-mentioned control factors. Localized hyperthermia in an experimental bone tumor model was induced and inhibition of the tumor growth by this hyperthermic therapy was confirmed radiologically and histologically. These results demonstrate that ferromagnetic bone cement is useful for the treatment of musculoskeletal tumors. KE Y WORD S ferromagnetic thermoseed, hyperthermia, bone cement INTRODUCTIO N The efficacy of hyperthermia for tumors heated above 42X has been confirmed by many researchers [1,2]. In the deep region, however, it is difficult to heat the tumor selectively. As a result of this problem, various methods have been developed for localized hyperthermia by magnetic induction heating [3,4]. We have developed a new ferromagnetic thermoseed of the cement type for local hyperthermia of the tumors in the deep regions, especially in bone. This material generates heat by applying an alternate magnetic field. It resembles bioactive bone cement [5] in composition, with a portion of the bioactive glass ceramic component replaced by magnetite (Fe304) powder and thus can be molded into various shapes. In this stucfy, we investigated this cement’s heat-generating ability under various conditions for its efficacy in an experimental model of bone tumors. MATERIAL S AN D METHOD S Ferromagnetic bone cement was supplied by Nippon Electric Glass Co., Ltd.( Ohtsu, Japan). This cement was composed of the two-paste type. In one paste, benzoyl peroxide (0.1 wt%) was dissolved and, in the other paste, N, N-dimethyl-p-toluidine (0.1 wt% ) was dissolved. The paste consisted of magnerite ( Fe304) and silica glass (SiOj) powders as filler and a BIS-GMA-based resin. The ratio of the filler to the resin was 9:1 (weight: weight). The average particle size of the magnetite powders was 535
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13 |Lim and that of the siHca glass powder 3 |Lim. The BIS-GMA- based resin was composed of bis-a-glycidylmethacrylate and triethylene-glycol dimethacrylate. Two pastes of the same weight were mixed and kneaded by hand for 1 minute. The mixture was molded into various shapes andsohdified in about? minutes. There was little rise in temperature of the cement during polymerizing. Investigation of heat-generatin g ability under various conditions The heat-generating ability of the ferromagnetic bone cement was investigated as a function of the content of magnetite, the volume of the cement, and the intensity of an appHed magnetic field. Four kinds of cement (containing this powder at 10%, 20%, 40%, or 80% weight ratio) were molded into a cubic block of 20 x 20 x 20 mm for studying magnetite content. The cement-containing magnetite at 40% weight ratio was molded into rectangular blocks of 20 x 20 x 20 mm, 20 x 20 x 10 mm, and 20 X 20 X 5 mm relative to cement volume. Each of these blocks was subjected to an alternate magnetic field to generate heat. The rectangular block (20 x 20 x 5 mm) was prepared relative to the intensity of the applied magnetic field and then subjected to altemate magnetic fields of various intensities. After the blocks were heated for ten minutes, the temperature of the surface on the blocks was measured with a fluoropdc thermometer (model 3000; Laxtron, Mountain View, CA, U.S.A.). Investigation of distribution of heat in the bone Japanese white rabbits weighing 2.0 to 2.5 kg were anesthetized with 25 mg/kg intravenous pentobarbital sodium. The cement (containing magnedteat 50% weight ratio), a molded pillar with a diameter of 4 mm and a length of 25 mm, was inserted into the medullary canal of the rabbit tibiae. After closure of the wound, the lower leg containing the cement was placed in an altemate magnetic field and heated Five sensors from the thermometer were used for thermometry in the leg. In the portion containing the implanted cement, one of the sensors was inserted at the surface of the medullary canal and two were at the medial and lateral interfaces between bone and muscle. In the distal portion, one was at the medullary canal at a distance 10 mm from the cement and one was at medial interface between bone and muscle at a distance 9 mm from the cement (Fig. 1 A). The temperature of the cement was adjusted to 50- 6 0 ^ . Local hyperthermia for experimenta l bone tumor Japanese white rabbits weighing 2.0 to 2.5 kg were implanted with VX2 tumor blocks, 2 x 2 x 2 mm, in the right tibiae under local anesthesia. Two weeks later, a second operation was performed. At this time, the rabbits were divided into two groups: the no treatment group and the hyperthermia
d e 16 24 32 4 0 time (min.) Figure 1 A: the position of thermometry. Surface on the cement (a), medial (b) and lateral (c) interfaces between bone and muscle, cortical surface(d) and medullary canal (e) in distal portion. B: time/temperature curve.
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New FerromagneticBone Cementfor Local Hyperthermia:K.^Takegamiet al.
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group. In the hyperthermia group, each rabbit was anesthetized with intravenous pentobarbital sodium. The medullary canal of the right tibia was curretted and the bone defect was filled with the ferromagnetic cement. An alternate magnetic field was applied to the right leg for 50 minutes. During heating, the temperature on the cortical surface of the tibia was maintained at 42-45"C. All of the rabbits were killed 5 weeks after transplantation of the VX2 tumor and radiographs were taken of the right leg. The right tibia was removed and evaluated histologically. RESULT S AN D DISCUSSIO N Heat-generatin g ability under various conditions An increase of magnetite content from 20% to 80% caused an increase in temperature of the material (Fig. 2a). Maximum temperature of 80% magnetite was 6IX! in the magnetic field of 80-Oe, 100 kHz. In regard to the cement volume, the temperature of the cement rose at afixedrate in proportion to the volume (Fig. 2b). The temperature of the cement kept rising as the intensity of the alternate magnetic field was increased (Fig. 2c). These results were similar to other ferromagnetic thermoseeds that produced heat by hysteresis loss [6]. As this material is of the cement type, an accurate estimate of the volume of the implanted cement can be determined prior to implantation. Therefore, it is important that the intensity of the magneticfieldcan be easily adjusted by the power output of the AC magnetic field generator and that the temperature of the cement also can be controlled. Distribution of heat in the bone In the rabbit tibiae, the cement was heated to 50X! within four minutes and then the temperature of the cement was maintained within a range of 50-60X! for approximately one hour. The temperature at the interface between bone and muscle increased to 43-45X and was maintained for the duration of this experiment (Fig. IB). Cell destruction specifically occurred in tumor cells without any apparent damage to normal cells at the latter temperature [ 1 ]. This demonstrated that heat damage to soft tissue around the bone is prevented by this method. On the other hand, the tissue surrounding the cement was heated by conduction and the temperature of the bone adjacent to the cement rose easily above 45X1. A temperature above 45X! is better for obtaining rehable tumoricidal effects to any tumor cells remaining after surgical resection. However, in the distal portion implanted with the cement, the temperature at the cortical surface did not increase above 42X! and the temperature at the medullary canal was maintained at 37X!. At some distance from the thermoseed, however, bone marrow was insufficientiy heated for hyperthermia. Local hyperthermia for experimenta l bone tumor In all rabbits, the radiolucency within the area of implantation was seen on radiographs taken 2 weeks post tumor transplantation.
0 5 10 15 20 60 10 0 14 0 18 0 20 40 60 b height of block (mm) content (96) Figure 2 The temperature change of this material a: relationship between temperature and volume, (base of block, 20 x 20 mm. Content of magnetite, 40%. Magnetic field, 70 Oe.) b: relationship between temperature and content of magnetite, (block, 20 X 20 X 20 mm. Magnetic field, 110 Oe.) c: relationship between temperature and intensity of magnetic field, (block, 20 x 20 x 5 mm. Content of magnetite, 40%.)
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Figure 3 Radiological findings of the no treatment group (A) and the hypertheraiia group (B), Pathological finding of the hyperthermia group (C). Figure 3 A andB show radiological findings of each group 5 weeks after transplantation. Although marked destruction of the tibia was seen and pathologic fractures occurred in the no treatment group, radiolucency, cortical destruction, and pathologic fractures were not seen in the hyperthermia group. Pathologic evaluation revealed viable peripheral and necrotic central portions in the no treatment group. In the hyperthermia group, necrotic portions expanded from the cement (Fig. 3C). The hind legs of rabbits were temporarily swollen after heat production, but no irreversible complications, such as skin necrosis, occurred These results demonstrate that the bone tumor was sufficiently and selectively heated in this system. CONCLUSIO N This study demonstrated that ferromagnetic bone cement is useful for local hyperthermia in bone. REFERENC E [1] Overgaard, J. Cancer, 1977, 39, 2637-2646. [2] Dichson, J.A. and Ellis, H.A. Nature, 1974, 248, 354-358. [3] Meijer, J. G., van Wieringen,N., Koedooder,C., Nieuwenhuys,G.J. and van Diji, J.D.P. Med. Phys. 1995, 22, 101-104. [4] Ohura, K., Kenaga, M., Nakamura, T., Yamamuro, T., Ebisawa, Y., Kokubo, T., Kotoura, Y. and Oka, M. J. Applied Biomaterials, 1991, 2, 153-159. [5] Kawanabe, K., Tamura, J., Yamamuro, T., Nakamura, T., Kokubo, T. and Yoshihara, S. J. Applied Biomaterials, 1993, 4, 135-141. [6] Jordan, A., Wust, P., Fahling, H., John, W. and Hinz, A., Felix, R. INT. J. Hyperthermia, 1993, 9, 51-68.
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
ADSORPTIO N OF L- LYSIN E ONT O SILIC A GLASS : A SYNERGISTI C APPROAC H COMBININ G MOLECULA R MODELIN G WIT H EXPERIMENTA L ANALYSI S Robert A. Latour Jr.^, Jon K. West^, Larry L. Hench^, Sharon D. Trembley ^, Yuan Tian^, Gary C. Lickfield"^, Alfred P. Wheeler^ 1 Department of Bioengineering, Clemson University, Clemson, SC, USA 29634 2 Department of Materials Sci. & Engineering, University of Florida, Gainesville, FL, USA 32611 3 Department of Materials, Imperial College, University of London, London, UK SW7-2BP 4 School of Textiles, Fiber and Polymer Science, Clemson University, Clemson, SC, USA 29634 5 Department of Biological Sciences, Clemson University, Clemson, SC, USA 29634 ABSTRAC T As a basic science approach to the problem of protein adsorption on implant surfaces, experimental and molecular modeling studies were conducted to study the adsorption of a protein residue (lysine) onto silica glass. Experimentally, 4 molecular weights of poly-L-lysine were adsorbed onto glass microspheres at 4 temperature levels. Adsorption enthalpy for each molecular weight was determined from the adsorption isotherms and plotted versus the degree of polymerization to estimate the enthalpy per adsorbed residue. Molecular modeling was also performed using a CAChe Worksystem. The geometries of the silica rings and lysine were optimized using the AMI semi-empirical method and the secondary structure of poly-L-lysine was evaluated using the MM2 molecular mechanics method. The modeling results yielded reasonably close correlation with the experimentally measured adsorption energy (-2.6 kcal/mol), and provided insights into likely adsorption mechanisms. KEYWORDS : adsorption, lysine, proteins, molecular modeling, computational chemistry INTRODUCTIO N Although numerous investigations have been conducted to investigate the adsorption of whole proteins to surfaces [1-5], because of the large numbers of variables involved, these studies have yielded relatively little information regarding protein conformational changes following adsorption. It is proposed that alternative approaches to protein adsorption are necessary to develop an understanding of protein adsorption behavior. Proteins are complex macromolecules with a minimum of 3 levels of hierarchical organization termed primary, secondary, and tertiary structure. It is proposed that to begin to understand, and thus predict the behavior of adsorbed proteins, we must first develop an understanding of the adsorption process at the most fundamental level (i.e. primary structural level). This understanding should then provide a basis to begin to sequentially address the adsorption process involving the higher levels of structure. This paper presents the results of a study of the adsorption behavior of proteins at the primary structural level. A combined experimental [6] and molecular modeling approach was taken which first involved the analysis of adsorption isotherm data to experimentally determine the adsorption enthalpy for mid-chain amino-acid adsorption onto a model glass surface. This was then followed by molecular modeling of the same system in order to provide insights into the potential molecular mechanisms involved in the adsorption process. 541
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MATEMAL S AND METHOD S The adsorption enthalpy for mid-chain L-lysine adsorption onto a model silica glass surface was first measured by a Langmuir isotherm adsorption process. Poly-L-lysine hydrobromide salts with poly-L-lysine molecular weights (MW) of 1000, 1000-3000, 5000-15000, and 15000-30000 g/mol (Sigma Chemical, St. Louis, MO, USA) were used as the adsorbate and borosilicate glass microspheres with diameter between 105-125 |Lim (Cataphote Inc., Jackson, MS, USA) were used as the adsorbent. Phosphate buffered saline (pH=7.3) was used as the solvent. Poly-L-lysinesaline solutions were mixed having initial concentrations (CQ) of 0.005, 0.010, 0.030, and 0.100 mg/ml. The microbeads were methanol washed, dried, and weighed into Eppendorf vials. 1.5 ml of each concentration of poly-L-lysine/saline solution was added to each vial and rotated endover-end for 2 h in a convection oven at constant temperatures of 25, 37, 45, and 55’’C to equilibrate. Control (saline without poly-L-lysine) and experimental vials were prepared for each concentration-temperature combination in duplicate and triplicate, respectively. A ninhydrin assay (Sigma Chemical Co., St. Louis, MO, USA) was used to measure the equilibrium concentration (Ce) of poly-L-lysine in solution. The amount of poly-L-lysine adsorbed on the glass at equilibrium (qe) was determined via mass balance by comparing CQ to Ce- The adsorption isotherm data was then plotted as qe vs. Ce for each temperature and MW of poly-L-lysine, and the Langmuir isotherm equation:
QC
was best fit to the data by nonlinear regression, with Q representing the amount of poly-L-lysine adsorbed at surface saturation and a = 55.0 exp(AG/RT), where 55.0 represents the activity of water in physiologic saline, AG = change in Gibb’s free energy/mole, R = ideal gas constant, and T = absolute temperature; and where AG = AH - T AS, with AH = enthalpy change/mol and AS = entropy change/mol. The mean – 95% confidence interval of AH was directly determined for each MW of poly-L-lysine by non-linear regression using SAS (Statistical Analysis Software, SAS Institute, Cary, NC). The values of AH were obtained for each MW range of poly-L-lysine and plotted versus the degree of polymerization (DP) of poly-L-lysine (lysine residue MW = 1 2 9 g/mol), with the slope of the initial linear portion of this plot providing an estimate of the average AH value per residue of poly-L-lysine adsorbed. Following the experimental work, molecular modeling of lysine adsorption to silica glass was conducted using a CAChe Worksystem (CAChe, Beaverton, OR, USA) through a Macintosh computing environment. The geometries for silicate rings and the lysine molecule were optimized using the AMI Precise unrestricted Hartree-Fock level of molecular theory [7]. The optimized structures included the interaction of 3 to 6 member silicate rings with lysine. In these silicate rings, there are generally 2 bridging Si-O-Si bonds between adjacent tetrahedra and 2 non-bridging Si-OH bonds, or silanols. Lysine has a peptide structure [-NH-CH(R)-CO-] with the R side-group being [-(CH2)4-NH3"*"]. The lysine molecule was terminated by -H and -OH groups to satisfy its molecular structure as required for the molecular models. In the modeling process, the lysine and silica rings were brought together and the geometry of the combined structure optimized to determine its minimum energy conformation. The energy released upon lysine adsorption to the silica ring was calculated as the difference in energy between the optimized silica ring and lysine molecule taken separately compared with the energy of the optimized silica ring - lysine complex. Because the experimental portion of this study involved poly-L-lysine (and not just lysine), it was recognized that secondary structural effects of the poly-L-lysine would likely play a role in the measured adsorption process. For this reason, a second modeling study was conducted to predict the likely secondary structure of poly-L-lysine, with this data then being used to assess the
Adsorption of L-Lysine Onto Silica Glass: L.A. Latour Jr. et al.
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Figure 1. Adsorption Isotherm. 1000 MW, 25*’C
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120
Figure 2. Enthalpy vs. Degree Polymerization
likely availability of the amine-terminated side-chains of poly-L-lysine to the silica surface. For this study, a 16-peptide chain of lysine (MW=2,082 g/mol) was modeled and optimized using an MM2 based molecular mechanics method [8]. RESULT S From the experimental adsorption studies, the adsorption isotherms for each poly-L-lysine MW force field were found to follow typical Langmuir behavior and were used to calculate the values of the adsorption enthalpy for each system. An example isotherm plot of qe versus Ce for 1000 g/mol poly-L-lysine adsorption onto the glass at 25 C is shown in Figure 1. The plot of AH vs. DP is shown in Figure 2. This reveals an initial linear portion, followed by an apparent plateau beginning around the DP=40 data point. The plateau evident for large DP is believed to reflect retained tertiary conformation of the poly-L-lysine molecular structure following adsorption. This effect has also been observed in polymer adsorption studies [9]. The best fit slope of the linear portion of Figure 2 is -0.23 –0.13 kcal/mol/DP (mean – 95% confidence interval), with this representing an estimate of the energy/mol released per average lysine residue adsorbed onto the silica glass surface. The molecular modeling AMI results predicted hydrogen bond formation between the sidegroup amine of lysine with the silanols of the 3-, 4-, 5-, and 6-member silicate rings with interaction energies (AH) of -4.6, +1.5,-1.1, and -6.8 kcal/mol, respectively. Figure 3 shows the results of the optimized 5-member ring-lysine structure. The relative mole fractions of 3-, 4-, 5-, and 6-member rings in silica glass have been estimated to be 0.026, 0.181, 0.275, and 0.518, respectively [10] and the Si02 concentration in the glass microspheres used in the experimental study was 72% [6]. Based upon these values, the net interaction energy of lysine with silica was calculated to be -2.6 kcal/mol. Although this value is higher than the experimentally measured adsorption enthalpy of -0.23 kcal/mol value, secondary molecular structure must also be considered. The MM2 optimization of the secondary structure of poly-L-lysine is presented in Figure 4. As shown, the analysis predicted that the molecule gently rotates with about 1 full rotation for every 8 residues. Thus, if this secondary structure is^maintained after adsorption, only about 1 or 2 out of every 8 residues may actually be accessible for interaction with the silica surface. Assuming this, the experimentally derived value may be artificially low by a factor of 4x to 8x, for a revised estimated experimental energy/mer value of -0.92 – 0.52 to -1.84 – 1.04 kcal/mole, which is reasonably close to the AMI result of -2.6 kcal/mol.
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Bioceramics Volume10
’ \ SILIC A X
Cten
53.5 Afi Silicon
P ^’2.511 A
^Carbon O Oxygen
Q^^^JTQ
i
^ 2 . 3 79 A
^
LYSIN E
0 Nitrogen O Hydrogen
Figure 3. 5-Member Ring-Lysine Structure
\ 19.5 A ^ Figure 4. Secondary Structure of Poly-L-lysine
SUMMAR Y The adsorption energy of poly-L-lysine onto silica glass was determined both experimentally and by using the AMI semi-empirical and MM2 molecular mechanics methods. The modeling results were found to agree reasonably well with the experimentally based value after considering possible secondary structural effects which could not be directly appreciated from the experimental results alone. This study illustrates the synergistic effect of combining molecular modeling with experimental studies in the investigation of adsorption phenomena in order to develop a better understanding of likely molecular-level behavior involved in adsorption processes. Further work is planned to expand the scope of this effort to address the adsorption of a wider range of amino acids onto silica glass and other surfaces, with the long range goal of using this information to develop more complex models for predicting the adsorption behavior of secondary and finally tertiary protein structure. ACKNOWLEDGMEN T The authors wish to acknowledge and thank Prof. Peter M. Burrows, Department of Experimental Statistics, Clemson University for assistance with the nonlinear statistical analysis of the experimental data; and the US Air Force for grant #F49620-95-1-0382 which provided funding for the molecular modeling portion of this research.
REFERENCES
Grinnel, F. and Feld, M.K., 7. Biol Chem.,1982, 257,4888-4893. Underwood, P.A., Steele, J.G. and Dalton, B.A., J. Cell ScL, 1993,104, 793-803. Underwood, P.A. and Bennett F.A., J. Cell ScL, 1989, 93, 641-649. Fabrizius-Homan D.J. and Cooper S.L., / Biomed.Mater.Res.,1991, 25, 953-971. Norde W., Pure Appl Chem.,1994, 66, 491-496. Trembley, S.D., M.S. Thesis, Dept. of Bioengineering, Clemson University, Clemson, SC. Dewar, M.J.S., Zoebisch, E.G., Healy, E.F. and Stewart, J.P., J. Am. Chem.Soc, 1985, 107, 3902-3909. 8. Allinger, N., J. Amer.Chem.Soc, 1977, 99, 8127-8134. 9. Korn, M. and Killmann, E., J. Colloid InterfaceSci., 1980, 76, 119-131. 10. Bell, R. and Dean, P., Phil. Mag., 1972, 25, 1381.
1. 2. 3. 4. 5. 6. 7.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EFFECT S OF DIVALEN T CATION S ON CALCIU M PHOSPHATE S PRECIPITATIO N ON A LANGMUIR-BLODGETT E MONOLAYE R S. B. Cho’, Y. Suetsugu\ J. Tanaka\ R. Azumi^ and M. Matsumoto^ ^National Institute for Research in Inorganic Materials, Namiki 1-1, Tsukuba, Ibaraki 305, Japan. ^National Institute of Materials and Chemical Research, Higashi 1-1, Tsukuba, Ibaraki 305, Japan.
ABSTRAC T Various kinds of cation-coupled Langmuir-Blodgett (LB) monolayer were prepared using arachidic acid and calcium phosphate formation on their surfeces was investigated in simulated body fluid (SBF). The results showed that the zeta potential of LB monolayer decreased by coupling with various divalent cations in the order of pure LB monolayer < Ca-coupled LB monolayer < Mgcoupled LB monolayer. Calcium phosphate precipitates were formed on the surfeces of pure LB monolayer, Ca- and Mg-coupledLB monolayers within 2 weeks, whereas a hydrophobic substrate without arachidic acid monolayer formed no precipitate at all. Optical microscopic observations showed that the morphology ofcalcium phosphate precipitated on the various substrates was affected by the kinds ofcation; especially, Mg-coupled LB monolayer gave dendrite-like calcium phosphate. KEYWORD S Langmuir-Blodgett monolayer, Calcium phosphate. Crystal nucleation, Zeta potential. INTRODUCTIO N Natural bone is a nanocomposite material, in which an assembly of small apatite particles is efectively reinforced by collagen fibers. This feet suggests that load-bearable bone-repairing materials can be obtained through organic-inorganic composite materials. In order to develop new kinds of bioactive organic-inorganic composite material, it is important to understand the intrinsic properties ofboth organic and inorganic materials. Further, knowledge on biochemical behaviors of organic-inorganic composite materials under body environment will allow the successfiil development of the bioactive composite materials. On the basis of these considerations, we have revealed interactions between polar groups of organic materials and inorganic ions using a Langmuir-Blodgett monolayer under the simulated body fluid whose ionic concentrations are nearly equal to those of human blood plasma [ 1, 2]. In the present study, in order to elucidate the effects of divalent cations on the precipitation of calcium phosphate, apatite forming ability of various kinds of cation-coupled LB monolayer was investigated using simulated body fluid. 545
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MATERIAL S AND METHO D (1) Preparation of Langmuir Blodgett Monolayer A slide glass substrate of 13 x 26 x 0.9 mm^ in size was beforehand hydrophobized with 1, 1, 1, 3, 3, 3 - hexamethyldisilazane and then dried at 105 C fori day. Ion-exchanged distilled pure water was used as an aqueous subphase. Arachidic acid (C19H39COOH, Kodak Co.) was dissolved in chloroform to obtain a spreading solution ofconcentration 5 x 10’ M. Arachidic acid of3 00 ^il was spread as the chloroform solution onto the aqueous subphase. After spreading, the film was left to equilibrate with subphase forabout 5 min priorto compression. Subsequently, the film was slowly compressed until the surface pressure needed for transfer to a substrate and kept at constant pressure for 5 min in order to fecilitate relaxation. The monolayer of arachidic acid was transferred onto the substrate by a vertical dipping method under a surface pressure of 25 mN-m"’ at a dipping speed of 4 mm min"\ Afterthe dipping, the substrate was dropped into the subphase and then removed from the subphase, following by soaking in aqueous solution containing difeent concentrations of divalent cations. (2) Soaking in Simulated Body Fluid A simulated body fluid(SBF) which had Na’ 142.0, K^ 5.0, Mg’^1.5, Ca’^ 2.5, Cf 147.8, HCO3" 4.2, HPO/’ 1.0, SO4’"0.5 mmol-dm’^ was prepared by dissolving reagents NaCl,NaHC03, KCl, K2HP04-3H20, MgCl2, CaCb and Na2S04 in distilled water after Kokubo et al. [3,4]. The fluid was buffered at pH 7.4 (at36.5 C) with 50 mmol*dm"^ oftris(-hydroxymethyl)-aminomethane ((CH20H)3CNH2) and 45 mmol*dm’^ hydrochloric acid (HCl). At least, three pieces of substrates were prepared for each soaking. The substrates were withdrawn from the aqueous solution containing different concentrations of divalent cations and resoaked in 30 ml of simulated body fluid. (3) Analysis of surface structure After the substrates were soaked in SBF forvarious periods, they were removedfromthe fluid and gently washed with ion-exchanged distilled water. The substrates were dried at room temperature. Their surfeces were analyzed by Fourier transformed infrared (FT-IR) reflection spectroscopy (Spectrum 2000, Perkin Elmer, U. S. A.). In the FT-IR reflection spectroscopy, the reflection angle was set at 75 . The specimen surface was observed by an optical microscope and an atomic force microscope (AFM: Nanoscope Ilia, Digital Instruments Inc. U. S. A.). The zeta potential of various substrates was measured by an electrophoretic light scattering method (LEZA-600, Otsuka Electronics Co., Japan). The solution used for the zeta potential measurement was bufered at pH 7.4 at 36.5 C with tris(hydroxymethyl)-aminomethane and hydrochloric acid. The concentration change of calcium and phosphorus ions in SBF before and afterthe soaking was measured with inductively coupled plasma (ICP) atomic emission spectroscopy (SPS1700VR, Seiko Instruments Inc., Tokyo, Japan). For the ICP measurements, 1 ml offluid was drawnfromthe bottle and added into 10 ml of ion-exchanged distilled water. RESULT S AND DISCUSSIO N Figure 1 shows the FT-IR reflection spectra of the surfece of the pure LB monolayer and hydrophobised slideglass soaked in SBF for 2 weeks. The IR peaks observed at 570, 610, about 1050 and 1120 cm’ can be ascribed to crystalline apatite [6]. This apatite contains carbonate ions, since peaks at about 875, 1420 and 1450 cm"’ areascribed to COs’^ ions [6]. On the other hand, the results ofthe FT-IR reflection spectra showed that no apatite formation occurred on the surfece of hydrophobised slideglass without LB monolayer even after2 weeks soaking in SBF. Two IR peaks
Effects of Divalent Cations on Calcium PhosphatesPrecipitationon LB Monolayer: S.B. Cho et al.
O apatite COg^’ions V slide glas s
9 CD O C
B
LB monolaye r
\ji^^
rt
\
V
"o DC
J
Tj
Y\
547
^ \
V
hydrophobise d glas s
1 200 0
1
1 160 0
1
L 120 0
1
80 0
40 0
Wave number/cm’ ^ Figure 1. FT-IR reflection spectr a of the surface s of pure LB monolaye r and hydrophobise d glas s soake d in SB F for 2 weeks .
pure LB monolayer
Ca-couple d LB monolayer
Mg-coupled LB monolayer
Figure 2. Optical microscopi c photograph s of the surface s of the LB monolayer , Ca-couple d and Mg-coupled LB monolaye r soake d in SB F for 2 weeks .
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Bioceramics Volume10
at 470 and 810 cm’^ are ascribed to Si-O-Si bending vibration, and one peak at 1100 cm’’ to Si-0 stretching vibration [5]. Figure 2 shows the optical microscopic photographs of the surface of the pure LB monolayer, Ca- and Mg-coupled LB monolayers soaked in SBF for 2 weeks. It can be seenfromFig. 2 that there is no remarkable difference in morphology ofprecipitates on the pure LB monolayer and Ca-coupled LB monolayer, whereas dendrite-like precipitates were formed on the Mg-coupled LB monolayer. Pure, Ca- and Mg-coupled LB monolayers showed the appreciable decrease in calcium and phosphorus concentrations in SBF during the soaking, but a little decrease in magnesium concentration. The hydrophobised slide glass hardly showed any change in element concentrations. The decreases in calcium and phosphorus concentrations for the pure, Ca- and Mg-coupled LB monolayer are attributed to the formation of calcium phosphates on the substrates surfaces of the by consuming calcium and phosphate ions from SBF. The zeta potential of pure LB monolayer, Ca- and Mg-coupled LB monolayers were-60.8, -47.2 and -52.9 mV before soaking in SBF, respectively. It is thought that the difeence in zeta potential between Ca- and Mg-coupled LB monolayers is attributed to the diftrence in bonding energy between the carboxyl groups of arachidic acid and divalent cations in aqueous solutions. The feet that zeta potential for Ca-coupled LB monolayer was larger than that for Mg-coupled LB monolayer suggests that Caions are more reactive to polar groups ofarachidic acid than Mgions in aqueous environment. After2 weeks soaking in SBF, the zeta potentials decreased to -23.2 (pure LB monolayer), -24.6 (Ca-coupled LB monolayer) and -24.9 mV (Mg-coupled LB monolayer), respectively, and calcium phosphate precipitates formed on their surfaces. These results indicate that Ca and Mg ion did not inhibit calcium phosphate precipitation but control the morphology of precipitates on their surfeces in SBF. CONCLUSIO N In order to elucidate the effects of divalent cations on calcium phosphate precipitation, various kinds of cation-coupled LB monolayer were prepared and then soaked in SBF for various periods. Calcium phosphate precipitates formed on the surfaces of pure LB monolayer, Ca- and Mg-coupled LB monolayers within 2 weeks, whereas the hydrophobic substrate without arachidic acid monolayer did not form precipitate at all. The morphology ofcalcium phosphate precipitated on the various substrates was affected by the kinds of cations; especially, Mg-coupled LB monolayer gave dendrite-like calcium phosphate. From these results, it was indicated that divalent cations coupled with polar groups ofLB monolayer mainly control the morphology ofprecipitates on their surfaces. REFERENC E 1. S. B. Cho, M. Kikuchi, J. Tanaka, R. Azumi, M. Matsumoto, in Proceeding of 5th World Biomaterials Congress, 1996, 240. 2. S. B. Cho, M. Kikuchi, J. Tanaka, R. Azumi, M. Matsumoto, In Bioceramics Vol. 9, Pergamon, Oxford, 1996, 403-406. 3. T. Kokubo, H. Kushitani, S. Sakka, T. Kitsugi and T. Yamamuro, J. Biomed. Mater. Res. 1990, 24, 721-734. 4. S. B. Cho, K. Nakanishi, T. Kokubo, N. Soga, C. Ohtsuki, T. Nakamura, T. Kitsugi, and T. Yamamuro, J. Am. Ceram. Soc, 1995, 78, 1769-1774. 5. K. H. Karlsson, K. FrobergandT. Ringbom, J. Non-Cryst. Solids, 1989, 112, 69-72. 6. J. C. Elliott, In: Structure and Chemistry of the apatite and other calcium orthophosphates, Elsevier, Amsterdam., 1994.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) 01997 Elsevier Science Ltd
EFFECT OF PROCESSIN G ON THE CHARACTERISTIC S OF A 20 vol.% PLATELET-REINFORCE D HYDROXYAPATIT E COMPOSIT E
AI2O3
S. Gautier, E. Champion and D. Bemache-Assollant Laboratoire de Materiaux Ceramiques et Traitements de Surface, ESA CNRS 6015, 123 avenue Albert Thomas, 87060 Limoges, France
ABSTRAC T Hydroxyapatite matrix composites containing 20 vol.% of alumina platelets were produced by slip casting and hot pressing. The influence of elaboration parameters on the microstructure and mechanical characteristics of the composites was investigated. The obtention of an homogeneous distribution of alumina platelets within the matrix requires the realisation of slurries containing 65 wt% of powder and 3.1 wt% of deflocculant. A significant toughening of HAP can be reached providing a good control of production parameters is obtained. In comparison with the monolithic hydroxyapatite (Kic = 0.75 MPa. vm ), the fracture toughness of composites increases up to Kic = 2.9MPa.Vnr. KEYWORD S : Composites, Hydroxyapatite, Alumina, Slurries, Toughness. INTRODUCTIO N Dense hydroxyapatite (Caio(P04)6(OH)2; HAP) exhibits an important brittleness which restricts its potential applications as biological implants. HAP ceramics are characterised by a very low fracture toughness, about 1 MPa vm [1-4], associated with subcritical crack growth under mechanical stresses when placed in a liquid environment [4,5]. Composite technology is known to be one way to improve the mechanical reliability of brittle materials [6]. We have demonstrated in a previous paper that the incorporation of monocrystalline alumina platelets could be efficient to toughen HAP ceramics [7]. Nevertheless, the presence of small unbroken aggregates of platelets in the materials constituted detrimental deffects for the mechanical reliability. This work is concerned with the determination of the influence of elaboration conditions on the characteristics of a 20 vol.% AI2O3 platelet-hydroxyapatite matrix composite. It consisted in evaluating the effects of slurries composition and hot pressing on the microstructure and toughness of the material in order to define the most appropriate production parameters. MATERIALS AND METHOD S Hydroxyapatite used in this study is a commercial powder (Bioland). It has a stoichiometric atomic ratio Ca/P = 1.667 and a specific surface area of 21.2 mV^- Alumina platelets (grade T’O, Elf Atochem) are disk-shaped monocrystals. They have an average diameter of 5 |im and a thickness of about 0.3 jim. 549
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Composite mixtures containing 20 vol.% of alumina were homogenised in demineralized water with ammonium polymethacrylate as deflocculant. The suspensions were ball milled during 5 hours in an alumina container. Green compacts were produced by slip casting of slurries in plaster moulds. Then, they were densified by hot pressing in an argon atmsophere. The relative density of composites was measured by the Archimedean method in water. Scanning electron microscopy (SEM) on chemically etched surfaces (lactic acid 0.15 M) was used for microstructural observations. Two types of experiments were performed to evaluate the influence of production parameters. (i) Experimental design [8] was used to investigate and quantify the effects of powder and deflocculant contents of the slurries. The rheological characteristics of composites suspensions were measured at 25 C using a rotational viscosimeter. The relative viscosity was determined at the constant shear rate of 350 s ^ Slip casted samples were hot pressed at 1200 C for 30 min, under a compressive stress of 30 MPa. (ii) To investigate the effects of thermal treatments on the microstructure and toughness, composites were hot pressed in the 1000 C to 1250 C temperature range for various times under a compressive stress of 10 MPa. In both experiments, the fracture toughness of hot pressed composites was determined by Vickers indentation, under a 4.9 load, from the measurement of the length of induced cracks. Kic was calculated using the relationship proposed by Evans and Charles [9]. Kic was measured either in the direction perpendicular (Kic -1-) or parallel (Kic //) to the hot pressing axis. RESULT S AND DISCUSSIO N Experimental design coordinates (i.e. (P)owder and (D)eflocculant contents of the slurries) and characteristics measured on composites materials are summarized in table 1. Whatever the composition, the behaviour of the slurries remains quasi-Newtonian and is associated with an appropriate relative viscosity for slip casting. Empirical relationships Ycai = f(P,D), given in a second degree polynomial form, were calculated to estimate the measured characteristics Yexp of hot pressed composites. A graphic representation of equations associated with relative density and Kic// are given in figures 1 and 2, respectively. Microstructural observations of composites sho\y important heterogeneities of platelets distribution in materials issued from slurries containing less than 60 wt% of powder (figure 3a). On the opposite, from 60 wt% an homogeneous distribution of alumina within the matrix is observed (fig. 3b). Table 1. Influence of slurries characteristics on densification and toughness of composite materials.
Powder (wt%) 70 50 65 55 65 55 60
Slurries characteristics Deflocculant Viscosity (wt%) (inPa.s) 4 4 4.8 3.1 3.1 4.8 4
70 22 40 25 35 25 30
Hot pressed materials (1200 C-30 min-30 MPa) Relative density K,ol K,c// (MPa. ViiT) T(% ) 2.3 1.5 98.0 1.6 1.9 90.9 1.0 98.2 2.3 96.4 2.1 1.3 96.4 2.9 2.0 1.7 86.9 2.9 98.6 2.0 1.5
Characteristicsof a 20 Vol.% AI2O3 Platelet-ReinforcedHA Composite:S. Gautieret al.
70
Figure 1. Relative density of composites .
551
P (% )
Figure 2. Fracture toughness (Kjc //) of composites .
Figure 3. Micrographs of hot pressed composites , (a) Heterogeneou s distribution of platelets , (b) homogeneou s distribution of platelets . 4 Deflocculan t (\vt%)
(5%)+
(70%) Powder (wt%)
Figure 4. Schematic representatio n of optimal slurry compositio n for composite production (in grey). (O): heterogeneou s distribution of platelets , (A): homogeneou s distribution of platelets .
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Table 2. Influence of hot pressing parameter s on composite s characteristics . Hot pressing conditions 1100 1200C 1250 1200 1250C
C-30min-10MP a - 30 min - 10 MP a C-30min-10MP a C-30min-30MP a - 4 hours - 30 MP a
Relative density (T%at–0.1)
Grain size (umatiO.l)
95.7 96.0 99.0 96.4 95.7
1.0 1.0 1.3 0.8 2.1
Kic//
Kiel
24«.3 2j«).3
1.4"-^
2.6^-=
1.5^’’ 2.0"’ 1.5"-’
(MFa-ViiT ) (MPa.>/m")
29–o.4 j7–0.3
j3«).3
From this experimenta l design, it is possible to define a domain for which the characteristic s of composite s are optimal (figure 4). Finally, it appears that the most appropriate slurry compositio n allowing elaboratio n of homogeneou s composite s with high relative density and toughness is 65 wt% of powder and 3.1 wt% of deflocculant . From this slurry compostion , composite s were heated for various hot pressing conditions , as summarized in table 2. The fracturetoughness of composite s decrease s with increasing average grain size of HAP matrix. Therefore , it is necessar y to prevent grain growth during densification . Kic values depends on the measuremen t direction, it is much more important in the parallel direction than in the perpendicula r one. This can be explaine d by a preferre d orientatio n of platelet s which tend to lie in plane perpendicula r to the direction of applied load during hot pressing. In comparison with the monolithic matrix (Kic = 0.75"^^^ MPa. vm ), the incorporatio n of alumina platelet s induces a strong toughening , with maximum values of Kic = 2.9 MPa. vm . The disk-shaped morphology of alumina allows to reach higher fracturetoughness than that obtained with the use of particles (Kic« 2 MPa. Vm" )[10]. CONCLUSIO N The incorporatio n of alumina platelet s in HAP matrix improves the mechanica l reliability. Thefracturetoughness can be more than three times that of monolithic HAP . The optimisatio n of the different characteristic s of composite s requires a good control of production parameters . To this end, experimenta l design appeared as a very useful technique of investigation . REFERENCE S 1. Jarcho, M., Bolen, C.H., Thomas, M.B., Bobick, J., Kay, J.F. and Doremus, R.H. J. Mat.ScL 1976, 11, 2027-2035 . 2. Akao, M., Aoki, H. and Kato, K. 7. Mat.Sci.1981,16 , 809-812 . 3. Halouani, R., Bernache-Assollant , D., Champion, E. and Ababou, A. J. Mat.Sci.Mater.Med. 1994, 5, 563-568 . 4. De With, G., Van Dijk, H.J.A., Hattu, N. and Prijs, K. J. Mat.Sci. 1981,16 , 1592-1598 . 5. De Groot, K. In: Bioceramics, Annals New-York Acad. Sci. 1988, 227-233 . 6. Rice, R.W. Ceram. EngngSci.Proc.1981, 2, 661-701 . 7. Champion, E., Gautier, S. and Bernache-Assollant , D. J. Mat.Sci.Mat.Med.1996, 7, 125-130 . 8. Box, G.E.P., Hunter, W.G. and Hunter J.S. Statistics for experimenters, John Wiley and Sons, New York, 1978. 9. Evans, A.G. and Charles, E. A. J. Am.Ceram.Soc.1976, 59 (8-9), 371-372 . 10. Noma, T., Shoji, N., Wada, S. and Suzuky, T. J. Ceram.Soc.Jap.1993,10 1 (8), 923-927 .
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
MECHANICAL
EVALUATIO N OF PHOSPHAT E BIODEGRADABL MEAN S OF INDENTATIO N METHOD S
E GLASSE S BY
F.J.Gil, R.Terradas*, J.Clement, G.Avila*, S.Martinez*, and J.A.Planell. Dept. Ciencia dels Materials i Enginyeria Metal.lurgica. ETSEIB. Universitat Politecnica de Catalunya. Diagonal 647. 08028-Barcelona. Spain. * Dept. Cristal.lografia, Mineralogia i Recursos Minerals. Universitat Barcelona. Marti i Franques s/n. 08028-Barcelona. Spain. ABSTRAC T Biodegradable glasses may be of great interest in certain clinical applications. The aim of the present work is to assess the mechanical behaviour of three different biodegradable glasses obtained by melting in the system CaO-Na20-P205. Indentation methods have been used for the evaluation of mechanical properties, namely Vickers hardness, Young’s moduli and fracture toughness. KEYWORD S Biodegradable glasses, CaO-Na20-P205 system, mechanical properties. INTRODUCTIO N Physico-chemical studies related with the dissolution mechanisms of glasses in water lead to the development of glasses within a compositional interval which favours the gradual lixivation of their constitutents [1]. These glasses are formulated in the system (Na20-CaO-P205), where the unit which defines the network are [PO4] tetahedra which act in a similar way to [Si02] units in the silica glasses. The addition of Na20 breaks the continuity of the structure weakening it and reducing the durability of the glass. The addition of divalent cations like CaO acts in the opposite direction by increasing their durability. In fact, it is the ratio CaO/Na20 which will control the dissolution rate of the glass, leading to dissolution times ranging between a few days and a few months. The dissolution of such glasses take place without leaving any residues behind. The first studies aming to biomedical applications of such glasses were carried out with a molar content in P2O5 rangmg between 30 and 50% [2,3]. Dental and maxilofacial applications were proposed for the filling of cavities, and cytotoxicity tests as well as in vivo implantation tests [3] have shovm that such glasses are non cytotoxic and that a limited reaction, less severe than for a suture material used as control, takes place in vivo. The improvement of the bone bonding of hydroxyapatite granules treated with glass of this same system has been studied [4]. At present it has been noticed that glasses of the system P205-CaO-Na20 are soluble in acqueous media, biocompatible with a very low cytotoxicity in hard and soft tissues and which may be of great interest when combined with different types of biomaterials. The aim of the present work is to prepare three different formulations and to assess their mechanical properties in order to gain a broader insight about their possible applications. MATERIAL S AND METHOD S Three different glasses of the system P205-Na20-CaO have been prepared with molar percentages below 34% in Na20 and with different CaO/P205 ratios in order to control their 553
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Bioceramies Volume10
solubility [3]. The raw materials used were NH4H2PO4, Na2C03 and CaCOs. The glasses were obtained by melting in a platinum crucible at temperatures ranging between 1000 to 1200 C and by quenching on a metallic plate pre-heated at 300 C. The materials were finally annealed at temperatures between 300 to 450 C depending on their glass transition temperature. The glass transition temperatures were determined by means of dilatometer at a heating rate of 5 C/min. It is well known that the fracture toughness of brittle materials can be determined by means of a Vickers indentation if two conditions are fulfilled: the volume of the cavity is accommodated by the net radial flow of material which will produce a semi-spheric plastic zone surrounded by an elastic matrix, and the radial cracks of the indentation, perpendicular to the specimen surface, have to be at least of the same length that the diagonals of the Vickers indentation. Although different models have been proposed, in this case the Antsis, Chantikul, Lawn and Marshall model [5] will be used, where the fracture toughness Kjc can be written as:
Kic
a
1/2/
(1)
3/2
\c
where O is a constant independent of the material and taken as 0.016 in the present model, P is the applied load and c is the crack length measured in the optical microscope, H is the Vickers hardness and E is the Young’s modulus. Moreover, it has been shown [6] that E/H can be determined by mean of a simple Knoop indentation as:
a’
a
(2)
E
where b/a is the ratio of diagonal of the Knoop indenter (approximately 1/7), bVa’ is the ratio of diagonal of the Knoop indentation after elastic recovery on unloading, and a is a constant determined experimentally. This mean that by combining a Knoop indentation and a Vickers indentation, E can be evaluated. The calibration has been performed with different materials where E and H were well kown [7], giving a value a = 0.52. The indentation tests were carried out on glass samples which were ground and polished with 0.05 fim alumina particles. Care was taken to obtain parallel flat surfaces. The Knoop microindentations were performed by applying a 100 g load during 30 s and the Vickers microindentations were done by applying a load of 300 g during 30 s. A minimum of 20 microindentations were carried out in each case and for each material. RESULT S AND DISCUSSIO N The compositions of the glasses prepared are shown in Table I and their positions in the ternary compositional diagram are shown in Fig. 1. Table 1. Chemical compositions in molar content of different glasses studied.
1
BV-01
,1
BV-15
L
CaO
NajO
PA
31.0 11.0 8.0
48.0 44.5
L
21.0 44.5 30.5
BV-11
61.5
1 1
Mechanical Evaluation ofPhosphateBiodegradableGlasses by IndentationMethods: FJ. Gil et al. 555 N02O
|700-1050«>C t
Figure 1. Ternary compositiona l diagram where the composition s BV-01, BV-11 and BV-15 have been are shown. The glass transition temperature s obtained by dilatometri c methods can be seen in Table II . Such temperature s have been used to give the annealing treatmen t to the glasses. The results of the microhardnes s indentatio n tests carried out are summarized in Tables III and IV . It can be noticed that the Vickers hardness for these materials ranges betwee n 2.74 and 3.57 GPa and their Young’s moduli betwee n 35 and 93 GPa. By means of a Vickers indentatio n and by measuring the radial cracks produced, the fracture toughness was also evaluate d and it ranged betwee n 0.79 and 3.45 MP a m’’’^. In order to obtain a bette r understanding , the critical strain energy release rate was also evaluate d in plain strain as:
G.=%^(1-.’)
(3)
where v is the Poisson ratio which for most glasses is approximatel y 0.2. The values obtained are also shown in Table IV . Table II . Glass transition temperature s of the glasses studied.
1
1
Glass
I^ilQ
1L
BV-01 329. 1
BV-11 445.1
_
BV-15 367.1
Table III . Vickers Hardness (HV) and Young’s modulus (E) for the different glasses studied.
1 1
Glass BV-01 BV-11 BV-15
HVlGPa l 2.74 – 0.07 3.57 –0.08 3.06–0.0 5
E (GPa) 1 35.62 – 2.35 92.82 – 5.26 53.04 + 7.04 1
1 1
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Table IV. Fracture toughness of the different glasses studied. 1
Glass BV-01 BV-11
L BV-15
Kic (MPa m^^^) 0.8210.05 0.79 –0.03 3.45 –0.12
Gic(kJ.m-’)
1
1.81x10-^ 0.65 X 10-^
21.54x10’^
1
The present results can be interpreted in terms of the different CaO and Na20 contents and their bonding to the P2O5 network. In fact the structure of vitreous phosphoric oxide is described by chains as in the case of a polymer. Both Na and Ca ions have the ability to break the P2O5 chains, but whilst Na does not produce any strong bonding between lateral chains, the effect of Ca is to link them [8]. In other words, whilst Na breaks and weaken the P2O5 chains, Ca links the chains together. This interpretation gives ground to understand that with increasing Ca content the glasses will become stiffer, harder and with a higher glass transition temperature. In fact this is in complete agreement with the results obtained. It can be noticed that as the CaO content increases: BV-01, BV-15 and BV-11, the glass transition temperature, the Vickers hardness and the Young’s modulus increase. However, fracture toughness seem to correlate with the P2O5 content. In fact, it seems also that Ca may act as an embrittlement agent, since for a minor variation in P2O5 for BV01 and BV-11, and a large increase in Ca content, the strain energy release rate decreases by a factor of 3. However, further research should be required in order to provide a proper explanation, since fracture toughness is also very sensitive to the defect population in the glass. ACKNOWLEDGEMENT S The authors of the present study are grateful to the CICYT for providing financial support through the project MAT96-0981. REFERENCE S l.C.F. Drake and M. Graham, Inorganic glasses as slow release herbicides and fungicides, Chemical Society, Burlington House, London, (1976). 2. J.Bumie, T. Gilchrist, S.R.L Duff, C.F. Drake, N.G.L. Harding and A.J. Malcolm, "Controlled release glasses for biomedical uses", Biomaterials, 2 (4), (1981), 244-245. 3. J.Bumie and T.Gilchrist, "Controlled release glasses: A new biomaterial", in Ceramics in Surgery , Ed. P.Vicenzini, Elsevier, The Netherlands, (1983), 169-176. 4. A. Afonso, J.D. Santos, M. Vasconcelos, R. Branco and J. Cavalheiro, "Granules of osteoapatite and glass-reinforced hydroxyapatite implanted in rabbit tibiae", J. Mater. Science: Materials in Medicine, 7, (1996), 507-510. 5. G.R.Anstis, P.Chantikul, B.R.Lawn and D.B.Marshall, "A critical evaluation of indentation techniques for measuring toughness: I", J. Amer. Ceramic Soc, 64 (9), (1981), 533-538. 6. D.B. Marshall, T. Noma and A.G. Evans, "A simple method for determining elastic-modulus-tohardness ratios using Knoop indentations measurements", Communications of the Amer. Ceramic Soc, (1982), C-175-176. 7. M.Ontan6n,F.J.Gil,A.Casinos,F.Guiu and J.A.Planell, "Young’s modulus and fracture toughness of cortical bone evaluated by means of indentation techniques", in Biomaterial-Tissue Interfaces, Eds.: P.Doherty et al., Elsevier, Amsterdam (1992), 171-177. 8. J.E. Shelby, W.C. Lacourse and A.G. Clare, "Engineering Properties of Oxide Glasses and Other Inorganic Glasses", in Ceramics and Glasses, Engineered Materials Handbook, Vol. 4, ASM, USA, (1991), 845-857.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
SILICO N IN CONNECTIV E TISSUE : SEMI-EMPIRICA MODEL S
L MOLECULA R ORBITA L
Keith D. Lobel\ Jon K. West\ and Larry L. Hench’’^ Department of Materials Science and Engineering, University of Florida, Gainesville, FL 32611 ^Department of Materials, Imperial College, University of London, Prince Consort Rd.
ABSTRAC T Semi-empirical molecular orbital modelling is used to predict the chemistry of interactions between soluble silicon, Si(0H)4, and biological macromolecules of connective tissues which are known to possess high concentrations of silicon. Results show a particularly favorable low-energy reaction pathway which forms ester-like linkages between silanol groups and carbonyl groups, with a reaction barrier of only 0.2 kcal/mol and a systemic stabilization of 4.6 kcal/mol. The effect of organic templates on the nucleation and growth of hydrated silica is also demonstrated. The potential impact of these data on the 1) osteogenic properties of silicate-based bioactive glasses and 2) employment of silica-rich mineral exoskeletons by marine organisms is discussed. KEYWORD S
silicon, silica, molecular models, protein, adsorption
INTRODUCTIO N The structural role of hard connective tissues in higher order animals is met by the presence of rigid carbonated hydroxyapatite crystals embedded within a flexible matrix of type I collagen and ground substance (interwoven proteoglycans and noncollagenous glycoproteins). However, both hard and soft connective tissues have been shown since the work of Carlisle [1], and Schwarz and Milne [2] to contain relatively high quantities of silicon (Si), which is believed to play an important structural role in these tissues. Biological evidence suggests that elemental silicon exists as cross-linking chemical species between chains of glycosaminoglycans, the major structural polysaccharides of vertebrates [3]. Little is known, however, of the chemical binding sites, or chemical/metabolic pathways, which result in linking the organic and inorganic constituents. The prevalence of hydroxy 1 and carboxyl groups of both proteins and saccharides in these tissues make such sites potential candidates for reaction with elemental silicon [1,2,4], and warrantfiirtherinvestigation. Certain types of lower order plants, such as diatoms and radiolarians, clearly demonstrate the use of polymerized hydrated silica for growth of inorganic structural exoskeletons, which are also known to be intimately associated with an organic matrix [5,6]. Although the Si is present here as bulk-form silica, the interactions between organic and inorganic phases may be similar to those in connective tissues of higher animals. Other physiological roles of silicon are implicated by the osteogenic characteristics of silica-based zeolites, zeolite extracts, sodium silicate solutions and certain compositional ranges of bioactive glasses [7]. The mnate complexities of these biological systems makes it difficult to identify the salient modes of interaction between silicon and biomolecules of the host. While its limitations must be recognized, molecular orbital modelling offers an additional investigative tool for studying the basic science of biological interfaces which is able to circumvent the inherent problems of real-system analysis. 557
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METHOD S Several types of dehydrolysis reactions were simulated using the semi-empirical AMI computational model. The AMI method is a self consistent field method of calculation which is parameterized for molecules containing silicon, carbon, oxygen, nitrogen, and hydrogen. A recent review justifies the application of this model to the systems at hand [7]. In an extension of earlier MO modelling work by the authors [8,9] reactions between silanol groups and various functional groups of amino acids and sugars are investigated. Reaction energy pathways were calculated and transition states for the molecular mteractions were determined and verified with Unrestricted Hartree Fock (UHF)-level force calculations. Care was taken to insure that terminal ends cf opposite R-groups did not meet each other and contribute a false component of systemic stabilization. Silicon cross-links were modelled by reacting two similar organic molecules, for example serine (Ser), with two silanol groups of a silicic acid molecule: 2 Ser + Si(0H)4 > Ser-0-Si-O-Ser + 2H2O In addition, the energetics of organic template stabilization on the process of silicification was studied by modelling the sequential condensation of four silicic acid molecules to form a tetrasiloxane ring covalently bound to the template, as described elsewhere [9]. In this study, efforts were concentrated on hydroxyl and carboxylic acid groups of both sugars and amino acid side-chains, since these have been proposed as the reaction sites of the organic host material, in both connective tissue (bone and cartilage) and organic templates of siliceous frustules [1-4]. Figure 1 demonstrates the possible linking sites of Si(0H)4 between two hyaluronic acid monomers. Protein template models were constructed to mimic the p-sheet structure, while saccharide templates were given free mobility. (The distance between two cw-silanol groups of a tetrasiloxane ring is 4.68 A, which is within 0.03 A of the interpeptide distance of the anti-parallel p-sheet). The structural stability of a protein p-sheet results from a combination of regular interpeptide hydrogen bonds and large barrier to peptide bond rotation [3]. These conformational effects are not as prominent in the polysaccharides, due in part to pure hydrocarbon backbones. RESULT S AND DISCUSSIO N The process of silonate bridge formation and biosilicification were modelled through, respectively, two and six 4-step dehydrolysis reactions. These consistent stages of each reaction include 1) hydrogen-bonding of reactants, 2) saddle point (activation state), 3) low-energy pentacoordinate metastable transition state, and 4) products. Similar mechanisms were found in other studies [10,11]. Complexation of silanols with hydroxyl groups form ether-like bonds (i.e. C-O-Si) while carboxyl groups formed ester-like bonds (i.e. C-OO-Si) (see Figure 1). Figure 2 shows a bar graph of the activation barriers and net degrees of stabilization fir each of the silonate bridge models. Only the systems that formed ester-like bonds resulted in products that were more thermodynamically stable than their reactants. The dramatic difiFerence between barriers of silanol complexation with glutamic acid (17.8 kcal/mol) and an acidic monosaccharide (0.2 kcal/mol) are believed to be due to different mechanisms of reaction rather than the different substrate structures. Figure 3 shows saddle points of the two reaction complexes. During ligation of silicic acid to glutamic acid, the resonance form shown above, Figure 3 a, was maintamed, and the -OH oxygen of the carboxyl group ligated with silicic acid. During reaction with the acidic saccharide, however, the carboxyl group adopted the resonance form shown in Figure 3b, and the non-protonated oxygen ligated with silicic acid. The more favorable energetics of the system in 3b may be due, in part, to 1) a more readily donated hydrogen from the cationic oxygen of 3b, and 2) the fact that simultaneous donation of H^ in 3 a to form water, and ligation with Si via the same -OH is not as sterically favorable as the latter case, where these two events occur via separate oxygens. Figure 2 also shows the activation barriers and net degrees of stabilization for five models of template-mediated biosilicification. Formation of a tetrasiloxane ring in the absence of an
Silicon in ConnectiveTissue: Semi-Empirical Molecular Orbital Models: K. Lobel et al. 559
25x Biosilicificatio n
NHCOCH3
C H , OH
«^ ,-# 1
C H j OH
^
0
’
Si(OH) 1
Models
1
COO"
1 ^
Sllonate-Bridg e ^
0f
2
Si(OH) 1
.^^
2
1
COO"
Barrier
NHCOCH3 Figure 1. Two hyaluronic acid monomers with two potential silonate cross-links
Stabilization
Figure 2. Thermodynamics of silonate bridge formation and biosilicificatio n models.
O-H
R-c 80 %
t
"^Wate r Formatio n
Resonanc e O-H
R-C %.
20 % -Wate r Formatio n Figure 3. Saddle structures for silonate bridge formation betwee n acidic monomers, (a) complexatio n to glutamic acid through the more prevalent (80%) resonance form of COOH . (b) complexatio n to an acidic monosaccharid e through the less prevalent (20%) resonance form of COOH .
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organic template has the greatest thermodynamic stabilization (-41.8 kcal/mol). Regulation by acidic organic templates was found to be more favorable than hydroxyl-rich templates in terms of both net stabilization and activation barriers. This data supports the theory of organic matrix mediated biosilicification: the initial nucleation of a silica-rich phase may occur via acidic sites on an organic template, as suggested by the low activation barrier for this system. Subsequent autopolymerization of silicic acid may proceed without further matrix influence to build the siliceous skeleton by the matrix-free model. Such a model meets the necessary role of lowering the energy of nucleation for biomineralization [5,12]. A similar functional role of collagen and sulfonated macromolecules of ground substance during the mineralization of bone is believed to exist [13]. Unfortunately the AMI model is currently not parameterized for calcium, making theoretical analysis difficult. Advances in molecular orbital chemistry will likely make fiulher analysis possible in the near future. A recently developed theory to account for the enhanced osteogenic properties of Class A bioactive glasses and glass-ceramics relative to Class B materials (e.g. hydroxyapatite) focuses on the differences in soluble silicon production [14]. The biomolecular and/or cellular mechanisms by which silicon may promote tissue production remams to be identified - indeed, a potential structural/metabolic dichotomy should be recognized. Whether the element induces rapid proliferation by merely supplying one of the necessary biulding blocks of connective tissue [1,2], or by actively triggering an increased metabolic activity of formative cells (e.g. chondroblasts, osteoblasts) [15], one can assume that highly specific interactions with biomolecules of these tissues will ultimately dictate the biological response. A more complete understanding of these interactions may help reveal the key to the unique behavior of silicate-based bioactive glasses. ACKNOWLEDGMEN T The authors would like to thank the Air Force OflBce of Scientific Research for financial support of this work, under Grant No. F49620-92-J-0351. REFERENCE S 1. Carlisle, E.M., Science 1972, 178, 619-21. 2. Schwarz, K. and Mihie, D.B., Nature1972, 239, 333-4. 3. Mathews, C.K. and van Holde, K.E., In: Biochemistry,Benjamin/Cummings Pub. Co., Redwood City, CA. 1990, 287. 4. Hecky, R., Mopper, K., Kilham, P. & Degens, E., Mar. Biol 1973, 19, 323-331. 5. Perry, C.C, In: Biomineralization:Chemicaland BiochemicalPerspectives,VCH, London 1989, 223-56. 6. Volcani, B.E., In: Silicon and SiliceousStructuresin BiologicalSystems,SpringerVerlag,New York 1981, 157-200. 7. Hench, L.L. and West, J.K., Life ChemistryReports1996, 13, 187-241. 8. Hench, L.L. and West, J.K., In: Bioceramics6,Buttenvorth-Heinemann, Oxford, 1993, 35-40. 9. Lobel, K.D., West, J.K., and Hench, L.L., Mar. Biol. 1996, 126, 353-60. 10. Burggraf, L.W., Davis, L.P. and Gordon, M.S. In: UltrastructureProcessing of AdvancedMaterials,Wiley, New York, 1992, 47-55 11. West, J.K. and Wallace, S., J. Non-Cryst.Solids1993, 152, 109-117. 12. Mann, S. et al.. Science1993, 261, 1286-92. 13. Lowenstam, H.A. and Werner, S., In: On Biomineralization, Oxford Univ. Press, New York 1989. 14. Hench, L.L., In: Bioceramics7, Buttenvorth-Heinemann, Oxford, 1994, pp. 3-14. 15. Keeting, P.E. et al., J. ofBone and Min. Res.1992, 7, 1281.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
TH E EFFEC T OF HEA T TREATMEN T ON BON E BONDIN G ABILIT Y OF ALKALI-TREATE D TITANIU M Shi gem Nishiguchi \ Takashi Nakamura\ Masahiko Kobayashi \ Wei-Qi Yan\ Hyun-Min Kim^, Fumiaki Miyaji ^, and Tadashi Kokubo^ ^Department of Orthopedic Surgery, Faculty of Medicine, Kyoto University, Shogoin-kawaharacho 54, Sakyo-ku, Kyoto 606-01, Japan ^Department of Material Chemistry, Faculty of Engineering, Kyoto University, Yoshida-honmachi, Sakyo-ku, Kyoto 606-01, Japan ABSTRAC T The purpose of this study is to evaluate the bone-bonding ability of alkali-treated titanium without heating. We made three types of titanium plate. 1) control group; pure titanium 2) alkali-treated group 3) alkali-and heat-treated group. The reaangular plates were inserted transcortically into the proximal met^^hyses of both rabbit tibiae. The tensile failure loads between implant and bone were measured by a detaching test. The tensile failure loads of alkali-and heat-treated group was 2.67 kgf and 4.13kgf, at 8 and 16 weeks, respectively. In contrast, those of control group and alkali-treated group were nearly 0 kgf even at 16 weeks. Histological examination revealed that alkali-and heattreated group was in direct contact with bone, but the other groups have thin intervening fibrous tissue between the implants and bone. A previous in vitrostudy demonstrated that alkali treatment without heating provide titanium an ability to form apatite on its surface. However our data showed alkali-treated titanium had no bone bonding ability. This discrepancy was due to the unstable surface reactive layer of alkali-treated titanium, which might be lost during preservation or implantation. In conclusion, both alkali and heat treatment are essential for preparation of bioactive titanium in practical use. KEYWORD S titanium, alkali treatment, heat treatment, bioactivity , bone-bonding IIVTRODUCTIO N In vitrostudy Kokubo et al showed that via a simple chemical treatment of alkali treatment and heat treatment, titanium and its alloys formed bonelike ^atite on their surface in simulated body fluid (SBF), which has ion concentration neariy equal to human blood plasma(1-4) In those articles, alkali-treatment alone also can provide titanium an ability to form apatite. Apatite formation on the material surface is believed to be a prerequisite for its bioactivity, that is, direct bone bonding.(5) In vivo study, Yanet al reported that alkali-and heat-treated titanium can bond to bone directly , and also showed that titanium treated alkali plus heat and SBF soaking has bone-bonding ability.(6,7) If alkali-treated titanium bond to bone without heat treatment, preparation of bioactive titanium will become much easier. And the effect of heat treatment to alkali-treated titanium on bone bonding ability is unknown. The purpose of this study is to investigate whether alkali-treatment alone can provide titanium bioactivity, and to investigate the influence of heat treatment on bone bonding strength. MATERIAL S AN D METHOD S Implant preparation We made three kinds of rectangular titanium plates (size; 15X10X2 mm). 1) control group; Commercially pure titanium plates (Kobe Steel Co., Kobe, Japan) were abraded 561
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with #400 diamond paste and washed with distilled water. And they were dried at room temperature. 2) alkali-treated group ; Pure titanium plates were soaked in 5 M NaOH aqueous solution for 24 hours. After the alkali treatment, the plates were gently washed with acetone and distilled water, and dried at 40*C for 24 h in an air atmosphere. 3) alkali-and heat-treated group ; After alkali-treatment, the plates were heated to 600*t at a rate of 5 C/min. in an elearic furnace, kept at a given temperature for Ih, and cooled to room temperature in the furnace. Implantation The implants were conventionally sterilized with ethylene oxide gas. They were implanted into the metaphyses of the tibiae of mature male Japanese white rabbits. The surgical methods was the same as that reported previously.(8) Using a dental burr, a 16X2 mm hole was made from the medial cortex to the lateral one parallel to the longitudinal axis of the tibial metaphysis. After irrigation of the hole with saline, the titanium plates were implanted in the frontal direction, perforating the tibia and protruding from the medial to lateral cortex. Four rabbits of each group were sacrificed at 8 and 16 weeks after the operation. In this study the guideline for animal experiments of Kyoto University was observed. Measurement of the detaching failure load After sacrifice, segments of proximal tibial metaphyses containing the implanted plates were cut out and prepared for the detaching test.(Figure 1) Traction through hooks holding the bone segments was applied vertically to the implant surfaces at a crosshead speed of 35 mm/min using an Instron-type autograph (Model 1011, Aikoh Engineering Co., Ltd., Nagoya, Japan). (8) The detaching failure load was measured when the plate was detached from the bone. If the plate was detached before the test, the failure load was defined as 0 Kgf. Data were expressed as mean – standard deviation (SD) and assessed using a one-way ANOVA. Differences at p < 0.05 were considered to be statistically significant. Histological Examination After the detaching test, specimens were fixed in 10% phosphatebuffered formalin and dehydrated in serial concentratbns of ethanol. Then, they were embedded in polyester resin. Sections 500 u m thick were cut with a band saw (BS-3000, EXACT cutting system, Norderstedt, Germany) perpendicular to the axis of the tibia, and were ground to a thickness of 150-180 u m for CMR and Giemsa surface staining using a grinding-sliding machine (Microgrinding MG-4000, EXAKT). Several 500 u m sections were polished with diamond paper and coated with a thin layer of carbon for observation by a scanning electron microscope (S-800, Hitachi Co. Ltd., Tokyo, Japan) attached to an energy-dispersive X-ray microanalyzer (EMAX-3000, Horiba Co. Ltd., Kyoto, Japan) (SEM-EPMA). RESUL T Detaching test In detaching test breakage always occurred at the bone-plate interface. The detaching failure load of each material at 8 and 16 weeks after implantation are summarized in Table 1. At 8 weeks after the operation the failure load of alkali-and heat-treated group was significantly higher than those of the other groups. At 16 weeks after the implantation, alkali-and heat-treated group showed a failure load of 4.13 kgf This value was significantly higher than that of the same group at 8 weeks and those of alkali- treated group and control group at the same weeks. The failure loads of alkali-treated and control group at 16 weeks did not differ significantly from those at 8 weeks. Table 1
Detaching test
failure loads (kgf; mean – SD) 8w 2.71+1.47 alkali-and heat-treated group (n=8) 0.52+0.52 alkali-treated group(n=8) control group(n=8) 0.02+0.03
16w 4.13 –1.25 0.49+0.38 0.33+0.36
Effect of Heat Treatmenton Bone Bonding Ability of Alkali-Treated Titanium:S. Nishiguchi et al.
563
Histological examination Histlogical examination by Giemsa surface staining and CMR revealed that alkali-and heat-treated group showed direct contact between bone and plate without any intervening fibrous tissue in all samples at 8 weeks(Figure 2). And at 16 weeks amount of bone directly contacted to the plate increased. In contrast, there wasfibroustissue layer between bone and the implant in both control and alkali-treated groups. At 16 weeks new bone formation increased around the plate, but intervening layer still remained in these groups SEM examination showed almost same findings as Giemsa and CMR Control and alkali-treated groups had a gap between bone and titanium, while alkali-and heat-treated plate contacted bone directly. Alkali-and heattreated group showed thin new bone layer formed on intramedullary portion of titanium far from cortex even at 8 weeks. Thisfindingwas not seen in alkali-treated and control groups. DISCUSSIO N Alkali- and heat-treated titanium was reported to bond to bone and supposed to be clinically applicable as orthopaedic implant material. (6.7) If alkali-treatment without heat treatment could induce bioactivity on titanium ,it would become possible to prepare bioactive titanium implant with soaking titanium implant in NaOH solution even at the operation theater. In this study, the bonding strength of alkali-and heat-treated titani um increased to 4.17kgf at 16 weeks by a detaching test. In contrast, alkali-treatedfitaniumshowed almost no bonding. And pure titanium did not bond to bone even at 16 weeks in accordance with the previous report. (9) Kim et al reported that alkalitreated titanium without heating had an ability to form apatite on its surface in SBF just like alkali-and heat-treated one.(3) And they also showed that once apatite layer formed, the tensile strength of the apatite layer and alkali-treated titanium is comparable to that of alkali-and heat treated one. (10) Thus if apatite layer forms, there will be no difference between bonding strength of alkali-and heat treated titanium and that of alkali-treated one. In this study, however, the bonding strength of alkali-treated titanium was significantly lower than that of alkali-and heat-treated one. This discrepancy can be explained as follows. Although alkali titanate hydrogel layer forms on titanium when soaked in alkali solution, this layer is so unstable as to lose its apatite forming ability during preservation or implantation. Throughout this study the implants were treated manually like as Figure 1. Detaching test
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orthpaedic implants and not treated so roughly. Although alkali-treated titanium has an ability to form apatite in SBF, it does not serve as bone bonding material in practical use. If alkali-treated titanium bond to bone, preparation of bioactive titanium would become easier. However this is not recommended from these results. In practical use both alkali and heat treatment is essential for inducing bioactivity on titanium. REFERENCE S 1. Kim, H. M., Miyaji, R, Kokubo, T., Nakamura ,T. J.Biomed.Mater.Res., 1996, 32,409-417. 2. Kokubo, T., Miyaji, F., Kim, H. M., Nakamura, T. JAm.Ceram.Soc.1996,79, 1127-1129. 3. Kim, H. M. , Miyaji, P., Kokubo, T., Nakamura, T. 7. Ceram.Soc. Japan, 1997, 105, 111-116. 4. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T., Yamamuro, T. J.Biomed.Mater.Res.1990,24, 721-34. 5. Kokubo, T., Ito, S., Huang, Z. T., et al. J. Biomed.Mater.Res. 1990,24, 331-43. 6. Yan, W. Q., Nakamura, T., Kobayashi, M., Kim, H. M., Miyaji, P., and Kokubo, T. J.Biomed. Mater.Res., to appear 7. Yan, W. Q., Nakamura, T., Kobayashi, M., Kokubo, T., Kim, H. M., Miyaji, P. In: Bioceramics volume 9, Elsevier, Oxford, 1996, 305-308. 8. Nakamura, T., Yamarumo, T., Higashi, S., Kokubo, T., Itoo, S. J.Biomed.Mater.Res.1985, 23, 631-648. 9. Takatsuka, K., Yamamuro, T., Nakamura, T., Kokubo, T. J Biomed.Mater.Res. 1995,29, 15763. 10. Kim, H. M., Miyaji, P., Kokubo, T., Nakamura, T. In: Bioceramicsvolume 9, Elsevier, Oxford, 1996, 301-304. Pigure 2. Giemsa surface staining image at 8 weeks; a.Alkali-and heat-treated group. Titanium implant has direct contaa with bone. b. Alkali-treated group. Thin intervening layer exists at the interface between bone and titanium. (X200)
B=bone T=titanium I=intervening layer
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
THERMA L PROCESSIN G OF COMPAC T BOVIN E BON E G. Vargas, M. Mendez, J. Mendez, and J. Lopez. CINVESTAV-IPN Unidad Saltillo, Apartado postal 663, 25000 Saltillo, Coahuila, Mexico.
ABSTRAC T In the present work, chemical and structural transformations undergone by compact bovine bone thermally processed up to 1500 C were studied using x-ray diffraction, scanning electron microscopy, and optical microscopy. The results showed that most of the organic matter was burned at temperatures below 500 C. The residual carbon content of bone was ~1 wt% at this temperature. The carbonate hydroxyapatite phase remained stable up to 600 C. At 1000 C, hydroxyapatite plus 0.6 wt% Na, 0.7 wt% Mg, and -0.1 wt% C, were found. Up to this temperature, 37 % of the initial weight of the boiled and skinned compact bovine bone had been lost. Between 1200-1500 C, mixtures of a-tricalcium phosphate, tetracalcium phosphate and hydroxyapatite, were observed. At these temperatures, the relative amount of each phase was affected by heating time and cooling rate. KEYWORDS : Bone calcination, calcium phosphates, hydroxyapatite, carbonate hydroxyapatite. INTRODUCTIO N When bones are thermally processed, water, living cells and other organic components, are removed, and inorganic matter suffers chemical and structural changes. Bovine bone calcining is a promissory process to obtain different calcium phosphates for biomedical applications such as drug delivery systems [1], bone repair and reconstructive surgery [2], as well as for manufacturmg of cements used in temporary or permanent dental and orthopaedic implants fixation [3]. Several research works have been carried out on bone calcmmg [4,5,6]. However, as different types of calcium phosphates may show different bioactivities inside a living being, it is necessary to fully understand their formation process during bone calcining. Thus, further studies about the chemical and structural transformations taking place during bone thermal processing, are required. MATERIAL S AND METHOD S Thin round sections of compact bovine leg bones were used. The first stage of bone processing was degreasing in boiling water. Subsequently, the bones were skinned before the remaining organic matter was burned at 300 C. After burning, each bone section was thermally processed for up to 24 hours at temperatures between 500 and 1500 C. Once cooled down to room temperature, the calcined bones were separated into two parts, one of which was subjected to a milling process, until the particles were smaller than 5|im, to carry out X-Ray Diffraction studies. The other bone piece was prepared using standard ceramographic techniques in order to carry out 565
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Studies of optical and scanning electron microscopy. More detailed information about the experimental equipment and techniques employed can be found in a previous work [7]. RESULTS AND DISCUSSION The following results were obtained from a previous work involving the chemical and microstructural characterisation of compact bovine bone processed up to 1000 C [7]: 1. Bone is dehydrated, fat is burned, and volatile matter is eliminated below 350 C, involving a total loss of-12.5% of the initial weight of the compact bovine bone. 2. Thermal cycles at 500 C yield products with up to ~lwt% of residual carbon, involving a total weight loss of ~34wt%. 3. The thermal treatment should be carried out between 500 and 600 C if carbonate hydroxyapatite has to be preserved, because the occurrence of carbonate decomposition was observed to take place between 600 and 700T. 4. The residual carbon content at SOOT is -0.14 wt%. 5. By subjecting powdered samples to x-ray diffraction, as well as through the examination of monolithic samples employing optical and scanning electron microscopy, only a single homogeneous phase was detected in bone processed at temperatures between 700 and 1000 C, which was identified as hydroxyapatite. 6. After treatment at lOOOT, 0.6 wt% Na, 0.7 wt% Mg, and 0.12 wt% C, were found. The weight of bone calcined at this temperature was approximately 63% of the initial weight of the boiled and skinned compact bovine bone.
(a)
(b)
100 l^m
Figure 1. XRD pattern (a) and Optical photomicrograh (b) of bone calcined at 1200 C for 24 hrs.
Thermal Processing of Compact Bovine Bone: G. Vargas et al.
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Figure 2. XRD pattern (a) and Optical photomicrograph (b) of bone calcined at 1500 C for 2 hrs. In the present work, the results obtained from further studies carried out at temperatures between 1200 and 1500 C, are reported. According to the x-ray diffraction pattern shown in Figure 1(a), hydroxyapatite was the main phase remaining after bone processing for up to 24 hours at 1200 C. However, as it is shown in Figure 1(b), at this temperature the incipient formation of a new phase was noticed on the periphery of the Haversian channels under the optical and the scanning electron microscopes. Later on, EDS microanalysis on the SEM allowed to identify the new phase as tricalcium phosphate. The rate of formation of this phase was very slow at 1200 C. Thus, the presence of tricalcium phosphate could not be detected neither by x-ray diffraction, even after 24 hours of treatment, nor by differential thermal analysis [7], although it was followed, employing monolithic bone samples, by sequential observations made under the optical microscope. After 2 hours at 1500 C, and allowing the samples to cool down rapidly outside the furnace, the presence of three different phases (hydroyapatite, tricalcium phosphate and tetracalcium phosphate) was confirmed by x-ray diffraction, as shown in Figure 2(a). The EDS spectrum obtained on the SEM for the light grey phase observed in Figure 2(b) is given in Figure 3(a). The semiquantitative determination of the Ca/P molar ratio obtained from this EDS spectrum gave a value of 1.6, which is close to the stoichiometric value of 1.667 corresponding to hydroxyapatite. The dark grey phase observed in Figure 2(b) showed a Ca/P molar ratio of 1.3 from the EDS pattern displayed in Figure 3(b). Since a suitable standard to carry out a more accurate determination could not be employed, this value was assumed to be reasonably close to the stoichiometric value of 1.5 corresponding to tricalcium phosphate. A small amount of Na was also detected in this phase. From the XRD diffraction pattern shown in Figure 2(a), and from the EDS studies, the white phase observed in Figure 2(b) was inferred to be tettacalcium phosphate. For treatment times longer than 24 hours at 1500 C, only mixtures of tricalcium phosphate and tetracalcium phosphate were found. On the other hand, by using slow cooling rates, the reversion into oxyapatite of some of the tticalcium phosphate and tetracalcium phosphate formed was observed, which was in agreement with previous observations made by Gottschling et al [8].
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REFERENCES 1. Itokazu M., Matsunaga T., Kumazawa S. and Wenyi Yang, Journal of Applied Biomaterials, 1995,6, 167-169. 2. Ricci J.L., Blumenthal N.C., Spivak J.M. and Alexander H., J. Oral Maxillofac. Surg., 1992, 50, 969-978. 3. OonishiH., Biomaterials, 1991, 12, 171-178. 4. Webster A. V., Cooper J. J., Hampson C. J. and Cubbon R. C. P., British Ceramic Transactions, 1987,86,91-98. 5. Hill R. G., Webster K., May C. and Mandal A., British Ceramic Transactions, 1994, 93, 16-20. 6. Cooper J. J., British Ceramic Transactions, 1995, 94, 165-168. 7. Vargas G., Mendez M., Mendez J. and Lopez J., sent to Journal of Applied Biomaterials. 8. Gottschling S., Kohl R., Engel A., Oel H. J., Bioceramics: Materials and Applications, Ceramic Transactions, 1995, 48, 201-213.
EVALUATIONS METHODS AND NEW APPLICATIONS
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Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CHARACTERIZATIO N OF SYNTHETI C AND BIOLOGICA L CALCIU M PHOSPHAT E MATEIUAL S BY MICRO-RAMA N SPECTROMETR Y G. Penel^ ^ G. Leroy\ G. Couraot^ andE.Bres^ ^U 279 INSERM 1 rue du Prof. Calmette, 59019 Lille Cedex, France; ^UFR de Chirurgie-Dentaire, place de Verdun 59000 Lille, France; ^ CNRS (URA583)-INRA-LNSA, 78350 Jouy en Josas; \SPES URA CNRS 234 USTL Bat.C6 59655 Villeneuve d’Ascq. France ABSTRAC T In this study the composition of biological calcium phosphates obtained in vivo (enamel, dentine and bone, implants of bone substitutes) or in vitro (resorption cavities) was investigated by micro-Raman spectrometry and compared to synthetic compounds. This technique provides high resolution molecular informationns in a non-destructive manner and without possible artifacts due to specimen preparation. KEYWORD S MicroRaman spectrometry - calcium phosphate - biomaterials - cell culture - prothesis. INTRODUTIO N Ciystallographic and chemicals investigations of calcium phosphate of biological interest are extensively performed with different methods (chemical analysis. X-ray diffraction, infrared spectrometry, MAS-NMR, TEM, etc.; [1-5]) that allowed to know their composition and structure, but numerous unresolved problems persist about the fine characterisation of the apatites. It has been shown that Raman spectroscopy, mainly focused to the vibration modes of P04^", OH’, HP04^’ and COa^’, can provide complementary informations in a non-destructive manner. In the biological calcium phosphates however this analysis is difficult because of the fluorescence caused by the organic components. This fluorescence can be reduced using high spatial resolution micro-Raman spectrometry. The purpose of the present study was to investigate by micro-Raman spectrometry the composition of biological calcium phosphates obtained in vivo or in vitroas compared to synthetic apatites. METHOD S Micro-Raman analysis were all performed with a OMARS 89 and a LABRAM microspectrometers from DELOR (Lille-France). RESULT S Synthetic bioapatites. Apatites of biological interest are mainly represented by tricalciumphosphate p (p-TCP), hydroxyapatite (GHAp) and carbonated apatite (CarAp). Micro-Raman spectra of these compounds are shown in figure 1. The spectra are dominated 571
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SUMMAR Y This study shows that micro-Raman spectrometry allows to strictly identified the different apatites contained in biological samples, by comparaison with the spectra obtained in synthetic compomids. Fmther studies will be required to characterize organic components. REFERENCE S 1. Arends J., Christoffersen J., ChristoflFersen M.R., Eckert H., Fowler B.O., Heughebaert .C, NancoUas G.H., Yesinowski J.P. and Zawacki SJ. J. of CrystalGrowth 1987,84,515-532. 2. Kleebe H.J., Bres E., Bemache-Assolant D., and Ziegler G. J.Am. Ceram. Soc,1997, 80, 37-44. 3. Elliott J.C. In Structureand chemistryoftheapatitesand othercalcium orthophosphates. Studiesin inorganicchemistry,Elsevier, London 1994, 191-301. 4. Penel G., Leroy G., Key C, Sombret B., Huvenne J.P. and Bres E. J. Mater.Sci. : Mater, in Med.,1997 {in press). 5. Bertoluzza A., Cacciari S., Tinti S., Vasina M. and Morelli M.A. J. Mater.Sci. 1995, 6, 76-79. 6. Fateley W.G., McDevitt N.T. and Bentley F. Appliedspectroscopy1971, 25, 155-173. 7. Tsuda H. and Arends J.J. Dent. Res.1994,11, 1703-1710. 8. Nelson D.G.A. and Williamson B.E. Aust.J. Chem. 1982, 35, 715-727. 9. Rehman I., Smith R., Hench L.L. and Bonfield W. J.ofBiomed.Mater.Res. 1995, 29, 1287-1294. 10. Sauer G.R., Zunic W.B., Durig J.R. and Wuthier R.E. Calcif Tissue.Int.1994, 54, 414420.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
BIOLOGICA L EVALUATIO N OF GLAS S REINFORCE D HYDROXYAPATIT BY FLO W CYTOMETR Y
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M.A. Lopes^’2, J.C. Knowles 3, K.A. Hing^ , J.D. Santos^, FJ. Monteiro^ and I. Olsen’* ^IRC in Biomedical Materials, Queen Mary and Westfield College, Mile End Road, London, El 4NS, UK; ^instituto de Engenharia Biom6dica, Faculdade de Engenharia, Universidade do Porto, Pra^a do Coronel Pacheco, N 1, 4050 PORTO, Portugal; Departments of Biomaterials^ and Periodontology*, Eastman Dental Institute, 256 Gray’s Inn Road, London, WCIX 8LD, UK
ABSTRAC T How cytometry (FCM) was used for evaluating the biological response of human cells to direct contact with materials with potential medical applications. The osteoblast-like cell line MG63 was grown on hydroxyapatite (HA) and P205-based glass reinforced hydroxy^atite composites (GRHA). Results suggested that the composites caused a delay in the progression of cells from GO/Gl into the S phase of the cell cycle. Moreover, although the cellular expression of collagen type I (COL I) was found to be similar in both materials, the level of osteocalcin (OC) was downregulated in the GR-HA compared with HA alone. The results of this study suggest that the FCM is a useful technique for assessing cell-biomaterial interactions. KEYWORDS : flow cytometry; cell culture; biocompatibility; hydroxyapatite INTRODUCTIO N A number of GR-HA materials with improved mechanical properties have been developed. However, their possible use for the repair and reconstruction of natural bone requires a critical biological evaluation. In vitrostudies of cytocompatibility have been carried out using several different assay methods, but the technique of flow cytometry (FCM) has previously not been widely used in thisfield.The present study has therefwe examined the use of the FCM technique to investigate cell-biomaterial interactions. More specifically, the effect of HA and GR-HA on the cell cycle and on the exiM-ession of COL I and OC were examined in a human osteosarcoma cell line. MATERIAL S AND METHOD S Preparation of discs A glass with a composition of 50.0%P2O5, 16.5%CaO and 33.5%MgO (mol %) was prepared. The composites were obtained by mixing 2.5 and 4.0 wt% of this glass with commercial hydroxyapatite. A detailed procedure for the preparation of this glass and composites has been reported previously [1]. The HA and the composites powders were uniaxially p-essed at 288MPa to form disc shaped samples. The discs were sintered at 1300 C for Ih, followed by natural cooling inside the furnace. All disc specimens were mechanically polished to 1 ^m finish, ultrasonically degreased and cleaned in ethanol followed by deionized water. Prior to cell culture the discs were sterilized in a dry atmosphere in a furnace at 180 ^C far 60 min. 575
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Cell culture The MG63 cells were routinely cultured at 37 C in a humidified atmosphere of 5% CO2 in air, in 75cm^ flasks containing 10ml of alpha Minimum Essential Medium (a-MEM), 10% foetal calf serum (PCS), 2mM L-glutamine, 50IU/ml penicillin and 50 |ig/ml streptomycin. Media were change every third day and, for sub-culture, the cell monolayer was rinsed twice with phosphatebuffered saline (PBS) and incubated with trypsin-EDTA solution for 10 min at 37 ^C. The trypsin was inactivated by adding the complete medium at room temperature, the cells washed twice by centrifugation and resuspended in complete medium for re-seeding and growing in new culture flasks. The 30nmi discs were placed into 6-well culture plates. Aliquots of 500|il of cells were carefully placed onto the discs and allowed to settle for 2h in the incubator at 37 ^C, after which 2.5ml of complete medium was added. In control cultures the cells were placed directly into the plastic culture dishes in 2.5ml of medium, at the same density of cells per unit area as placed onto the discs. Cell cycle progression The cells were plated at 5000 cells/cm^ on the 30nun discs and the plastic dishes as described above. Approximately 36h after plating they were stained for DNA content using propidium iodide (PI), as follows. The cells were detached and suspended in 2ml of 70% ethanol for 30 min, then centrifuged and resuspended in 400|Lil of PBS. They were incubated with 100|Lil of RNase A (Img/ml) and 100|il of PI (400|il/ml) at 37 ^C for 45min and analyzed by PCM as described bellow. Immunofluorescent staining of antigens The cells were plated at 10000 cells/cm^ on 30nmi discs and directly on plastic dishes. The expression of collagen type I (COL I) and osteocalcin (OC) were measured after 6 days of culture. The cells were washed twice with PBS and detached using 20mM EDTA in PBS for approximately 5 min at 37 C. They were centrifuged at 400g for 7min and the pellet was resuspended and again centrifuged. They were fixed in l%(w/v) paraformaldehyde in PBS for 30 min, then washed by centrifugation and resuspended in washing buffer. This contained PBS, 2% PCS and 0.05% sodium azide. Aliquots of 10^ fixed cells were used to measure the level of each antigen by PCM. Cells were first permeabilized for 10 min using washing buffer with 0.1%(w/v) saponin, then washed, centrifuged and resuspended in washing buffer. Cells were incubated for 60 min at room temperature in either mouse monoclonal antibody (mAb) against human COL I, diluted 1:10, or in rabbit polyclonal antibody against human OC, diluted 1:100. Mouse IgGl and normal pre-inmiune rabbit serum were used as negative controls. Cells were then washed, centrifuged and resuspended again in washing buffer with 0.1% saponin. Secondary antibodies, fluorescein isothiocynate (PITC)conjugated rabbit anti-mouse IgG and FITC-conjugated swine anti-rabbit IgG, diluted 1:20, were added for 30 min at room temperature. Cells were washed again and resuspended in 400|xl of washing buffer and analyzed. FCM analysis The light scattering properties andfluorescenceof cells stained with PI and PTTC were measured on a PACScan flow cytometer (Becton Dickinson). Analysis were performed on 10000 cells. Data were collected, stored and analyzed with CELLQuest Software (Becton Dickinson).
Biological Evaluation of Glass ReinforcedHydroxy apatiteby Flow Cytometry:M.A. Lopes et al.
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RESULT S AND DISCUSSIO N The light scattering properties of FCM can be used to examine certain fundamental moq)hological characteristics of cells, since the intensity of light scattered at small angles (forward scatter; FSC) is considered proportional to cell size and the orthogonal light scatter (side scatter; SSC) is proportional to the granularity [2,3]. In the present experiments there were no significant differences in either FSC or SSC between cells grown on the GR-HA, HA or plastic dishes (data not shown). The cell cycle is the entire process by which a cell divides into two identical daughter cells. It is divided into five main phases: GO and Gl (pre-DNA synthesis; resting or interphase), S (DNA synthesis), G2 (post-DNA synthesis) and M (mitosis). The red fluorescence signals of cells stained with PI were used to obtain histograms of the DNA content of the individual cells in the MG63 cultures grown in the presence of the different materials, as described in the Materials and Methods. A representative histogram can be seen in Figure 1 in which the first peak corresponds to cells in GO/Gl, having a 2n DNA content, and the second peak corresponding to cells with G2+M DNA content of 4n. Cell scored in the trough have a DNA content intermediate between Gl and G2+M, and they usually considered to represent cells in S phase. As can be seen from the results shown in Figure 2, a higher proportion of quiescent cells with a lower proportion of cells in S phase was detected in the cultures grown on GR-HA compared with HA and control cells. This findings suggest that the GR-HA causes a delay in the progression of the MG63 cells through the cell cycle, although further studies are required to determine whether these materials affect the subsequent proliferation capacity of the cells. We also used the FCM to evaluate the effect of the GR-HA on cellular function, as measured in the present study by the levels of expression of COL I and OC, both of which have previously been shown to be key antigens in bone development, repair and regeneration. The results in Figure 3 show thefluorescencevalues of the GR-RA and HA cultures normalized to that of the cells grown directly on plastic culture tissue dishes. Although there was little difference in COL I expression between the different cultures, it is notable that OC levels were significantly lower in the GR-HA composites compared with HA. This may be explained at least partly by a delay in differentiation in the GR-HA cultures resulting from the apparently inhibitory effects of this material on the progression of the MG63 cells into the S phase of the cell cycle. Further studies were carried out to clarify the biological effects of the GR-HA materials. X-Ray diffraction (XRD) revealed a beta-tricalcium phosphate phase ((J-TCP) in addition to HA in all the composites. Table 1 shows the Rietveld analysis of these phases. Since it is known that P-TCP has a higher dissolution rate than HA, different cellular response would be expected. Moreover, we also found a slightly higher water contact angles of the GR-HA composites compared with HA (unpublished results), suggesting their higher hidrophobicity may also have an important influence on the progression of the cell cycle progression and hence on the functional activity of the cells.
Table 1 - Rietveld results for HA and GR-HA. HA
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CONCLUSION The results of this study show that while the GR-HA materials appeared to have no major deleterious effects on MG63 cells, they nevertheless caused some perturbation in the progression of the cells through the cell cycle and also appeared to alter the regulation of a pivotal bone antigen. Moreover, our results show that the FCM technique is likely to be powerful new tool for assessing biocompatibility. ACKNOWLEDGMENTS The support of the JNICT (BD 1355 and project PBICT-C-CTM-1890-95) is gratefully acknowledged. REFERENCES [1] J. D. Santos, J. C. Knowles, F. J. Monteko, G. W. Hastings; Biomaterials, 15 (1994) 5. [2] M. G. Ormerod; How Cytometry, 1994, IRL Press. [3] H. M. Shapiro; Pratical Row Cytometry, 1995, Wiley-Liss Inc.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
EVALUATIO N OF MACROPHAG E RESPONS E TO CERAMI C PARTICLE S BY FLOW CYTOMETRY : ANALYSI S OF PHAGOCYTOSI S AND CYTOTOXICIT Y I. Catelas\ R. Ma^chand^ L’H. Yahia\ O. L. Huk^ ^Ecole Polytechnique de Montreal, Institute of Biomedical Engineering, P.O. Box 6079, Succ. Centre-Ville, Montreal, Quebec, Canada, H3C-3A7, ^Microbiology and Immunology Department, Ste-Justine Hospital, Montreal, Canada, ^Lady Davis Institute for Medical Research, Jewish General Hospital, Montreal, Canada, ABSTRAC T Using the J774 cell line, we designed an in vitro model to analyse the effects of size, concentration and composition of ceramic particles (AI2O3 and Zr02) on phagocytosis and cell mortality by flow cytometry. The kinetic of phagocytosis was also analysed to determine at what time cell stimulation is maximal. Results reveal that phagocytosis increases with size and concentration for particles up to 2 |Lim. For larger particle range (up to 4.5 |im), phagocytosis reaches a plateau, which suggests a saturation of phagocytosis, most likely dependant on overall particle volume ingested. Cytotoxicity studies revealed that macrophage mortality increases with particle size and concentration for size greater than 2 |im and for concentrations up to 500 particles per macrophage. Smaller particles (0.6 |Lim) cause significant cell mortality only at higher concentrations (up to 1250 particles per cell), and the mortality is still very low (< 10%). No significant difference appears between AI2O3 and Zr02. Kinetic studies revealed that phagocytosis of the particles begins very early after cell exposure, increasing with time and particle concentration, and reaching a plateau at 15 hours, which implies that the optimum period to evaluate cellular response to particulate debris should be between 15 and 24 hours of incubation. KEYWORD S Ceramics, particles, macrophages, flow cytometry INTRODUCTIO N Aseptic loosening, initiated by wear particles, is without a doubt, the main cause of Total Hip Arthroplasty (THA) failure. Particle composition has been thought to play a role in the inflammatory reaction to wear debris generated from THAs. In this regard, ceramics have been considered bioinert. However, recent studies have shown that even ceramic particles can induce a macrophage foreign body response [1], and stimulate the production of bone resorbing mediators in vitro[2]. While analysis of pseudomembranes retrieved at revision surgery has provided much information on the particle sizes generated in vivo and has identified a myriad of inflammatory mediators that are released by macrophages, very little is understood about the exact parameters to which the macrophage responds and about the kinetics of this response. Indeed, many confounding factors are inherent in clinical studies, and it is difficult to determine the relative importance of each parameter, such as size, concentration and composition of the particles on ultimate macrophage response. Using flow cytometry, we developed a new in vitro model for 579
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testing cellular response [3] and the standardisation of flow cytometry for this type of in vitro model should be applicable to the study of cellular response to any kind of biomaterial. The purpose of this study was to analyse the effects of size, concentration and composition of ceramic particles (alumina AI2O3 and zirconia Zr02) on phagocytosis and cell mortality, using flow cytometry. The kinetics of phagocytosis was also analysed to determine at what time cell stimulation is maximal. MATERIAL S AND METHOD S Particles AI2O3 (Alcoa (Bauxite,Arkansas)), Alcan (Jonquieres,Quebec),Durmax (Chatellerault, France)) and Zr02 particles partially stabilised with yttrium (Unitec Ceramics (Stafford, England)) were commercially obtained in the phagocytosable range: 0.6 (Alcan), 1.3 (Alcoa), 2.4 (Alcan) and 4.5 jim (Durmax) for AI2O3 and 0.6 jim for Zr02 (Unitec Ceramics). Particles were sterilised by ethylene oxide. Cell culture The J774 cell line (ATCC, USA) was used in this in vitromodel. Macrophages were cultured and maintained in RPMI 1640 tissue culture medium (Gibco) with 10% foetal bovine serum and 2% antibiotics (gentamicin, penicillin, streptomicin). Cells were exposed to ceramic particles in tubes, containing 4.10^ cells in 1 ml of culture media with a particle suspension at concentrations varying from 30 to 1250 particles per macrophage. Culture tubes containing macrophages but no particles served as negative controls. Incubations were conducted at 37^C, in a 5% CO2 environment, for 24 hours. The tubes were then washed with PBS at pH 7.2, and cells were stained with propidium iodide (PI) for 4 minutes, at room temperature. Tubes were again washed and data were collected using flow cytometry. Methods Phagocytosis and cytotoxicity tests were conducted using flow cytometry (FACScan, Becton Dickinson). Flow cytometry analysis of phagocytosis determines the percentage of macrophages that have phagocytosed the particles (stimulated macrophages), by detecting changes in cell granularity and/or size following particle ingestion. This percentage is called the phagocytosi s index. The cytotoxic effect of particles is evaluated as cell mortality, and is obtained by measuring changes in cell fluorescence, after staining stimulated macrophages with propidium iodide. Kinetic studies allowed us to evaluate, when controlling for size, the progression of phagocytosis with time and concentration of particles. Kinetic tests (5 min to 24 hours) were conducted on AI2O3 at 1.3 jxm. RESULT S AND DISCUSSIO N Phagocytosis Figure 1 presents the effect of particle sizeand concentrationon phagocytosi s index. In the small particle range (< 2 |Lim), the phagocytosis index increases with particle size and concentration. In the larger particle range, the index reaches a maximum, independent of concentration. This suggests a saturation of phagocytosis, most likely dependant on particle volume ingested. As particle concentration increases, the percentage of stimulated macrophages increases. For particles larger than 2 |im, more than 80 % of macrophages are stimulated, independent of concentration. This later effect of maximal phagocytosis for large particles that was independent of concentration suggests that the larger the particle, the higher its chance of being seen by the cell and phagocytosed.
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Cytotoxicity Figure 4 presents the effects of AI2O3 particle size and concentration on cell mortality. Results reveal that macrophage mortality increases with particle size and concentration,and the later is only significant for particle size greater than 2 |im for concentrations up to 500 particles per macrophage. Smaller particles (0.6 |im) cause significant cell mortality only at higher concentrations (up to 1250 particles per cell), and the mortality is still very low (< 10%). However, macrophage stimulation was elevated at that size, as evidenced by the phagocytosis index as high as 80%. Despite evident stimulation at this combination of parameters, the cell is not sufficiently triggered to die. This demonstrates a distinction between the toxic and stimulatory effects of particulate debris on macrophages. With regards to composition, there was no difference in mortality between AI2O3 and Zr02 at the size (0.6 \xm)and concentrations we analysed. CONCLUSION S The evaluation of the macrophage inflammatory response in relation to various particulate debris parameters has not been thoroughly studied at the cellular level. Whereas microscopic evaluation of phagocytosis provides qualitative information, flow cytometry allows us to quantify the morphologic changes observed. The results of this study demonstrate that flow cytometry is a quantitative and objective technique to evaluate the inflammatory response to particulate debris. More specifically, we have shown that macrophage response to ceramic debris depends on size and concentration, which is in agreement with results in the literature [5] and is independent of ceramic composition. Our kinetic studies reveal that phagocytosis reaches a plateau at 15 hours of incubation, which implies that the optimum period to evaluate cellular response to particulate debris should be between 15 and 24 hours of incubation. Thus far, the clinical implications of this study indicate that if particulate debris in the size range up to 4.5 |Lim is generated in sufficient concentration, a macrophage response will be triggered, and that this response is most likely dependant on the volume of phagocytosed particles. Our results strongly support the concept that macrophages do not perceive absolute particle numbers, but rather, respond to overall particle volume. Future prosthetic design efforts should also focus on wear resistant interfaces to reduce debris. While not truly bioinert at the particulate level, the superiority of the ceramic-ceramic articulating interface in terms of wear resistance, merits continued study of this joint couple for THAs. REFERENCE S 1. Lerouge, S., Huk, O.L., Yahia, L’H., and Sedel, L., J. Biomed.Mater. Res. 1996, 32, 627633. 2. Nakashima, Y., Shuto, T Hayashi, K., Hotokebuchi, T., Yasuda, K., and Sugioka, Y., ORS, 1995, 780. 3. Catelas, I., Huk, O.L., Marchand, R., and Yahia L’H., In: BioceramicsVolume 9, Otsu, Japon 1996, 93-96. 4. Devrets, D.A., and Campell, P.A., J. of Immunol Methods1991,142, 31-38. 5. Shanbhag, A.R., Jacobs, J.J., Black, J., Galante, J.O., and Giant T.T., J. Biomed.Mater.Res. 1994,28,81-90.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
STUD Y OF POROU S INTERCONNECTION S OF BIOCERAMI C ON CELLULA R REHABITATIO N IN VITRO AND IN VIVO J.X. Lu\ B. Flautre\ K. Anselme\ A. Gallur^, M. Descamps^ B. Thierry^ and P. Hardouin^ 1. IRMS, Institut Calot, Rue du Docteur Calot. 62608 Berck-Sur-MerFRANCE. 2. CRITT Ceramiques Fines, Z.I. du Champ de 1’Abbesse. 59600 MaubeugeFRANCE.
ABSTRAC T Our study would find the role of porous interconnections (PIC) on bone ingrowth and material degradation in tricalcium phosphate beta (P-TCP) with about 50% porosity, a size of 100-300 ^m macropores and of 30-100 ^im PIC. In vitro,human osteoblasts were cultivated on discs with two delays: 14 and 28 days. In vivo,implants were implanted in middle diaphysis of both femurs with two delays: 12 and 24 weeks. The in vitro and in vivo samples were observed with histomorphometry (HMM) and scanning electronic microscopy (SEM). In vitroresults show that human osteoblasts penetrate sizes of PIC over 20 ^m diameter (dia.), set up and grow inside the bioceramic macropores. In vivo HMM results show that PIC directly influence bone ingrowth inside pores and material degradation. We notice that a PIC size over 50 \imdia. allows formation of new bone ingrowth inside the pores. The PIC density expresses the link between pores inside porous materials, assures the cells proliferation and tissular diflFerentiation by extra cellular and vascular exchanges. In resorbable materials, pores and PIC initial densities play a more important role than their sizes. KEYWORD : Bioceramic, porosity, porous interconnections, osteoblasts. INTRODUCTIO N The physico-chemical composition of tricalcium phosphate ceramics is identical to the mineral constituents of our skeleton. As far back as 1920, they were used as bone substitutes and showed the ability to be recolonized by bone tissue.^^^ Then, their use has been given up. But since 1970, a lot of studies investigated their behaviour in biological fluids. They were considered biocompatible either in vitroor in vivoP’^^ Meanwhile, the biological behaviour of these ceramics depends on factors involved in the processs such as the raw powders used, mineral phases, micro (pores size < 10 ^m) and macro (pores size > 100 ^m) porosity. Except for the chemical composition, degradation rate is directly influenced by the specific surface and the volume of microporosity; Macropores size and macroporosity volume influence bone ingrowth in these porous ceramics.^^’^^ Previous studies have mainly investigated bone ingrowth inside pores, but only a few ones have investigated the relationship between porous interconnections (PIC) and bone ingrowth. PIC is literally a pathway between pores to favour the cellular and vascular penetration which assures the bone ingrowth inside pores. The aim of our study was to observe the PIC influence on cellular and tissular bone ingrowth in porous materials. MATERIA L AND METHOD S Biomaterial: Beta tricalciimi phosphate (p-TCP) with a macroporosity of about 50 % and a microporosity of about 4.2% (measured by Hg porosimeter), a pores size of 100-300 ^im and a PIC 583
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size of 30-100 ^m (measured by morphometry), and a Ca/P ratio of 1.55 – 0.03 (measured by Inductive Coupling Plasma) was used in the study. In vitrostudy: Osteoblasts were obtained from cancellous bone tissue of iliac crest of a 6-year-old boy. Cells were used after a second passage. p-TCP discs of 13.2 mm dia. and 1.4 – 0.2 mm height were sterilised by dry heat at 180 C during one hour. Discs were soaked in culture medium under vacuum. An osteoblasts suspension was prepared with culture medium in order to obtain 4 x 10^ cells/ml. One ml of cells suspension was then inoculated on the discs in each well. Complete medium was changed twice a week. After 14 and 28 days of incubation (37.0 – 0.2 C temperature, 5.0 – 0.1% CO2 and > 98% humidity), the discs were recuperated for HMM and SEM analysis. In vivostudy: p-TCP cylinders were implanted in rabbit. Implants size was 3 mm dia. and 6.1 – 0.2 mm length, and they were sterilised l^ 25 KGY gamma radiation. 8 white female New Zealand adult rabbits were 4 – 0.3 Kg body weight and 6-month-old before implantation. In rigorous asepsis conditions and under general anaesthesia, at the lateral external side of each femur a cavity of 3 mm dia. and 6 mm depth was made with an electric drill ((j) 2.8 mm, 200/minute speed) and then (() 3.0 mm manually in each middle femoral diaphysis. The created defect was washed from bone debris with saline solution. One implant was inserted in each cavity perpendicular to diaphysis. 17 and 2 days before sacrifice a double label of oxycycline and alizarin complexon was performed. After 12 and 24 weeks of implantation, the animals were sacrificed by an overdose of thiopental sodium (Nesdonalfi). The middle femoral shafts were removed and fixed in 10% neutral buffered formol for HMM analysis. Histomorphometry : Undecalcified bone preparation and polymethylmetacrylate embedding were used for each sample. Microradiographs were performed from 100 ^m thick sagittal sections. Then 50 ^m sections were stained with May Griinwald Giemsa staining for in vitrostudy and with Van Gieson’s Picro-Fuchsine staining for in vivo study. The following HMM parameters were measured: mineral apposition rate (MAR, mcm/d), new bone volume (BV/IV, %), residual material volume (MV/IV, %), material degradation rate (MDR, % = 100% - measured MV/IV / initial MV/IV), relative osteoid tissue volimie (OV/BV, %), macropores (MP) size (^m) and density (number/mm^), PIC size (^mi) and density (nimiber/mm^). Statistical analysis: Results were expressed by means and standard deviations. Impaired bilateral Mest (SYSTATfi) was used to compare the delays. Correlations (PIC mean size, MP mean size, PIC mean density, MP mean density and the BV/IV) were studied with Pearson correlation coefficient (r). RESULT S In vitrostudy: SEM analysis (A7 = 2): a cellular layer spread on the materials area, and covers the opened MP. On a broken surface of cylinders, osteoblasts are observed in the MP. After 28 days of culture the cells penetration is deeper than after 14 days. Some osteoblasts inside pores go through the PIC (Fig. 1). Nevertheless, no modification of the microstructure of the ceramic was observed after culture with SEM. On stained sections observed with transmitted light microscope (n = 4), cells are seen inside of PIC. The size of PIC containing is between 19.15 and 215.43 ^m dia. with a mean size of 78.67 – 37.41 ^im. The maximalfrequencyPIC size with cells is 60 ^m dia. after 28 days of culture. In vivo study: MAR is higher in receiving bone near the implant than inside the implant. Inside pTCP, MAR and BV/IV increases (p < 0.05), and OV/BV and MV/IV decreases {p < 0.05) with time (Table 1). On stained sections, the pathways size with bone tissue is 19.15 - 201.06 ^m dia..
Study of PIC of Bioceramics on Cellular RehabitationIn Vitro and In Vivo: J.X. Lu et al.
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but no significant difference (the mean value of size mean of each sample and their standard deviation: 74.40 – 5.29 for T12 weeks and 77.32 – 5.47 for T24 weeks) is observed between the delays. Table 1.The HMM parameters before and after Implantation Delay n Bone P-TCP Implant Weeks MAR(mcm/d)MAR(mcm/d) BV/IV(%) OV/BV(%) MV/IV(%) MDR (%) TO 8 59.79 –3.23 12 8 1.78 + 0.32 0.91+0.17 35.00 + 6.64 7.48 + 5.16 52.89 + 4.58* 11.49 –7.66 24 8 1.7410.2 1 1.25 –0.10* 52.07+7.53* 3.42 + 2.93* 42.57 –7.71* 28.77 –12.90* * Compare between the delays with/? < 0.05. In microradiography, MP size and density increase with time. The curve of PIC distribution shows a translation in the biggest sizes with aflattenedcurve (Fig. 2). Mean size and density of PIC have increased (p < 0.05) with time, but no significant difference is observed for the PIC density between 12 and 24 weeks (Table 2). Figure 1. SEM in vitro p-TCP, 28 days: osteoblasts (short arrows) go througji the PIC (lc»ig arrow).
Figure 2. Porous Interconnections Distribution of p-TCP
10
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Interconnection s size Oim)
Table2. Characterisatio n of (3-TCP microstructur e before and after implantation n Delay Sample Macropores Porous Interconnection s n Density {nlvam^) (Weeks) Size (^im) Density (w/mm^) Size (fim) TO 8 14.46 –1.67 165.62 –6.60 26.45 –1.90 63.93 –3.81 12 8 17.58 –1.11* 178.15 –11.56 * 21.72–2.15 * 78.59–5.21 * 24 8 18.74 –2.49 202.63 – 12.20 * 18.41 –1.77* 93.83 –7.63*
D The mean value of size inean(or density meam) of every sample and their standard deviation. * Comparison between delays with/? < 0.05. Correlations: A correlation between initial MP size and density is observed (r = -0.789, p < 0.05). No correlation is foxind between the initial PIC size and density, neither between the PIC and the MP. At T12 weeks, a conelation between PIC density and MP density is seen (r = 0.765, p < 0.05), also a correlation between BV/IV and the PIC size (r = 0.733, p < 0.05), or between BV/IV and the pores density (r = -0.736, p < 0.05). At T24 weeks, there is also a correlation between the PIC size and MP size (r = 0.880, p < 0.01) or between the PIC size and MP density (r = -0.786, p < 0.05). DISCUSSION GALOIS et al.^^^ did not found rats dermfibroblastsinside HA (hydroxyapatite) or p-TCP implants with different pores size after 5 and 40 days of culture. However, our in vitrostudy shown that human osteoblasts have spread in the pores of porous ceramics and are able to through inside the
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PIC. However the osteoblasts are seen only in the first 3 rings of ceramic (1.2 – 0.2 mm depth) but never in the center. To avoid the hydrophobic phenomenon of porous ceramics, the air bubbles were extracted with vacuum and stained coloration was justified. According to our results, the osteoblasts can cross through 20 fim PIC size. KLAWITTER and HULBERT^^^ have noted that the PIC size of calcium aluminate should be at least 100 ^m size to favour the mineralised bone tissue growth, between 40 and 100 |iim for osteoid tissue and at last 5 - 15 ^m for fibrous tissue. SHIMAYAKI and MOONEY^^^ have studied comparatively 230 and 600 ^m dia. of pores with respectively 190 and 260 ^m dia. of PIC in porous HA implanted in rabbit femurs. Results shown a higher bone ingrowth for 600 ^m than 230 ^m pores size. In our study, MAR and BV/IV increase significantly with time. PIC size and density increase in time and facilitate the pathway of vessels. Mean size of PIC with mineralised bone tissue inside was about 70 ^m diameter. We have measured the Volkman canals size inside the ceramic on microradiography. Their size was 10-20 \\mdia. with a maximal frequency of 20 ^m size. It could be the reason why a correlation is observed between new bone formation and PIC size or and pores density after 12 weeks. Because of an increase of p-TCP degradation at 24 weeks no correlation with new bone formation is observed at this time, but close correlation yet in their microstructure. WHITE and SHORS^^^ considered that PIC size must be superior than 100 ^m. But, our results indicate that a PIC size greater than 50 ^m allows the mineralised bone formation inside pores of material. Because our two-dimensions measurement we could precize the maximal PIC size, but we failed to precize the minimal PIC size. CONCLUSIO N These results show that in vitro,human osteoblasts are able to cross through PIC and to spread inside the pores. A minimal PIC size of 20 fim dia. is required but the ideal size is about 60 ^m diameter. In vivo,a PIC size greater than 50 fim allows formation of new mineralised bone tissue. Density and size of PIC and MP can be modified by the biodegradation of the (3-TCP. New PIC can improve the cells, vessel and biological liquids penetration and facilitate the material degradation and bone ingrowth inside the material. Therefore for resorbable materials the MP and PIC initial density play a more important role than their size. REFERENCE S 1. F.H.ALBEE,^«w.5’wrg. 71(1920)32. 2. L. Galois, D. Mainard, K. Bordji, H. Membre, L. Marchal, B. Foliquet, D. Clement and J.P. Delagoutte, in "Actualitesen materiau"(Romillat, Paris,), 3 (1990) 361. 3. J.J. Klawitter and S.F. Hulbert, J. BiomedMater.Res.Symp.2 (1971) 161. 4. K. Shimazaki and V. Mooney, J. Orthop.Res.3 (1985) 301. 5. E. White andE.C. Shors, Dent Clin. Nor. America.30 (1986) 49. 6. D. Groot, in the Materialcharacteristicsversusin vivo behavior,P. Ducheyne and J. Lemons (eds.), N.Y. Acad. Sci., 523 (1988) 227. 7. R.Z.L. Geros, JR. Parsons, G. Daculsi, F. Drissens, D. Lee, ST. Liu, D. Peterson and M. in vivo behavior,P. Ducheyne and J. Lemons Walker, in the Materialcharacteristics versus (eds.), N.Y. Acad. Sci., 523 (1988) 268.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
REPAI R OF OSTEOCHONDRA L DEFEC T USIN G ARTIFICIA L CARTILAG E
ARTICULA R
M. Hasegawa, A. Sudo, Y. Shikinami* and A. Uchida Department of Orthopedic Surgery, Mie University, Faculty of Medicine, 2 -174 Edobashi, Tsu City, Mie 514, and *Takiron CO., LTD., 405, Nagano, Yasutomicho, Shisou-gun, Hyogo 671-24, Japan
ABSTRAC T A new artificial articular cartilage was used for to repair an osteochondral defect. To test its biocompatibility, a large full-thickness defect was created in the patellar groove of a rabbit, and was filled with artificial articular cartilage tightly. Macroscopically, the repair tissues appeared as glistening smooth surface partially covering the implants. On histological and immunohistochemical evaluation, the ingrowth of bone and hyaline cartilage-hke tissue was found surrounding and within the implants. All the repairs with or without implants were associated with slight synovitis of the knee joint. The artificial articular cartilage [three-dimensional fabric (3-DF)] has bulk and surface biocompatibility, and could serve as both a scaffold for cartilage formation and a prosthesis. KE Y WORD S artificial articular cartilage, osteochondral defect, histological evaluation INTRODUCTIO N It has been reported that large articular cartilage defects of mature animals have little capacity for repair. Repair of full-thickness osteochondral defects has been attempted by various methods. Some experiments have achieved successful repair [1-7], but there is no highly acceptable method for complete repair of hyaline cartilage. The purpose of this study was to test the biocompatibility and repair potential of a threedimensional fabric (3-DF) developed as a new implant for artificial articular cartilage [8] to enhance the healing of large osteochondral defects in the rabbit knee. MATERIAL S AN D METHOD S The 3-DF used in this experiment is a prototype of FABRICUBE^^ (Takiron Co., Ltd). This triaxial-three dimensional (3A-3D) fabric was fabricated by the method described below. Low density linear polyethylene (LLDPE) was melted and used to the surface of ultra-high molecular weight polyethylene (UHMWPE) fibers that was a yam of 500 denier (50 filaments) to prepare a yarn with an average diameter of 400|im (Yam A). Yam A was treated with a hand operation practice machine which was designed to produce a 3 A-3D structure. The 3-DF is a block with an orthogonal fiber alignment and a size 10 x 10 x 3mm (Figure 1). Its compression behavior (mechanical strength and profile) was quite 587
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similar to that of natural cartilage. The surface of Yam A was oxidized by corona-discharge. Thereafter, the LLDPE was softened by heating, and a micro powder of unsintered hydroxy apatite was sprayed onto the surface. After cleaning the surface with water, partial exposure of the powder from the surface was confirmed. Fourteen skeletally mature female J^anese white rabbits, weighing 3.2 - 3.9 kg, were used After shaving, anteromedial arthrotomy was performed under intravenous pentobarbital sodium anesthesia (20mg/kg body weight). The patella was dislocated laterally, and the articular surface of the distal femur was exposed A full-thickness osteochondral defect, measuring 10mm from proximal to distal transversing the entire width of the patellar groove and 3mm in depth penetrating the subchondral bone plate generated in the patellar groove of the femoral condyle using osteotome, rongeur and curette. The defect on one knee was filled with sterile artificial cartilage with the surface of the implant was parallel to the original cartilage surface (group A; n= 14). The defect on the other knee was left empty as a control (group B; n = 14). Joint capsule and skin were sutured as separate layers. After the operation, all rabbits were allowed free cage activity without immobilization. Two rabbits with artificial articular cartilage on one side and control on the other were excluded because one died 11 days after the operation, and one sustained a dislocation of the patella on the knee with the implanted artificial cartilage. The remaining the 12 animals were sacrificed with a lethal dose of pentobarbital sodium administered intravenously from 2 to 12 weeks postoperatively. The distal part of the femur and synovium were examined macroscopically. For microscopical evaluation, the distal femur was removed, fixed in 10% buffered formalin, decalcified in Plank-Richlo solution at A C and embedded in paraffin. Sagittal sections 5-10|im thick were cut and stained with hematoxylin andeosin (HE) and alcian-blue. Immunostaining with an antibody to type n collagen was visualized by the avidin-biotin immunoperoxidase method Synovial tissue samples were collected from the supr^atellar pouch, fixed and embedded in the same way, then stained with HE to evaluate inflammation and immune responses.
Figure 1. artificial (FABRICUBE ).
articular
cartilage
Figure 2. Photomicrograph showing the ingrowth of cartilage into the implant at 8 weeks postoperatively (alcian-blue, original magnification x 100).
Repair of OsteochondralDefect Using Artificial Cartilage: M. Hasegawa et al.
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RESULTS Macroscopic Observations Neither joint contracture nor infection was found in any animals. Slight intra-articular effusion was observed only at 2 weeks postoperatively in both groups, but decreased with time. In group A, the implant was covered with white glistening tissue on the medial and lateral side 2 weeks postoperatively. Tissues around the implant became thicker with time, but no thick tissues were formed on the center of the artificial cartilage by 12 weeks postoperatively. In group B, the defect was filled with soft brown tissue, and the surface was very rough and depressed below the surrounding original articular surface 2 weeks postoperatively. The defect was filled with white and opaque tissue, which did not reach the level of the original articular surface during the experimental period The synovial membrane was slightly hyperemic and hyperplasic in both groups 2 weeks postoperatively, but these observarions became more normal with time. Microscopic Observations In group A, cartilaginous rissue was found at the base and side of the implanted material at 2 weeks postoperarively. Ingrowth of fibrous tissue into the implant was seen, and the newly formed tissue on top of the implant was fibrous. The ingrowth of bone and cartilage into the implant was seen 4 weeks postoperatively (Figure 2). Hyahne cartilage-like tissue and bone were observed covering part of the implant in 8-week specimens (Figure 3). Hyaline cartilage-like tissue was stained by alcian-blue and immunostained by the antibody to type II collagen. Most cartilage in contact with the implant was replaced by bone. Abundant bone ingrowth into the implant was seen in the deep region 8 weeks postoperatively (Figure 4). No bone was found at the articular surface by 12 weeks postoperatively. No inflammatory reaction was observed In group B, the surface of the defect was depressed below the normal level by 12 weeks postoperatively. Fibrous tissue including scant hyaline cartilage-like tissue was seen in the superficial region and bone was observed in the deep region. Synovial rissue showed mild inflammation including lymphocyte infiltration and subsynovial fibrosis at 2 weeks in group A. However, these changes tended to diminish with time, and the same findings were seen in group B. No foreign body giant cells or debris particles were found.
Figure 3. Photomicrograph showing neocartilage formation covering part of the implant at 8 weeks postoperarively (HE, original magnification x 40).
Figure 4. Photomicrograph showing the ingrowth of bone into the implant at 8 weeks postoperatively (HE, original magnificarion x 40).
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DISCUSSIO N Various methods for to repair of osteochondral defects have been reported using organic [1-3] and inorganic materials such as collagen matrix [4], carbon fiber [5], polytetrafluoroethylene [6], polyester [6] and polylactic acid matrix [7]. Artificial materials can have the mechanical function at an early stage and be easily shaped to fit the defect, but have drawbacks such as fixation and the risk of debris particle generation. When carbon fiber was used as a scaffold for repair of full-thickness osteochondral defects, cartilage repair was not obtained and minimal synovitis with pigmentation were observed [5]. Implantation of polytetrafluoroethylene (Tefron) and polyester (Dacron) showed early neocartilage formation without normal cartilage morphology and synovitis [6]. When polylactic acid matrix was implanted in cartilage defect with and without periosteal grafting, polylactic acid matrix with periosteal grafting showed formation of cartilage resembling articular cartilage [7]. The experiments described above involved small defects (3.0-3.7mm in diameter) produced in the knees of rabbits. Small osteochondral defects less than about 3mm in diameter have been suggested to be repaired naturally, although repaired cartilage never fully resembles normal hyaline cartilage biochemically [9]. In this study, we found early formation of cartilage in contact with the artificial cartilage in the osteochondral defects. The implant coated by hydroxy apatite was fixed firmly to the subchondral bone which is necessary for artificial cartilage. Inflammatory reactions against the implant were slight, shown as synovial hyperplasia in 2-week histological specimens. Complete repair of articular cartilage defects was not achieved during the experimental period because of large the size of the defect. However, this implant could be used as a prosthesis. We conclude that this implant has good bulk and surface biocompatibility, and could serve as both a scaffold for bone and cartilage formation and a prosthesis.
REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9.
O’Driscoll, S. W., Keeley,F. W. and Salter, R. B. J. Bone JointSurg. 1986, 68-A, 1017-1034. Billings, Jr., E., von Schroeder,H. P., Mai, M. T., Aratow, M., Amiel, D., Woo, S. L.-Y. and Coutts, R. D. Acta OrthopScand.1990, 6 1, 201-206. Grande, D. A., Pitman, M. I., Peterson, L., Menche,D. and Klein, M. J. Orthop.Res. 1989, 7, 208-218. Speer, D. P., Chvapil, M., Volz, R. G. and Holmes, M. D. Clin. Orthop.1978, 25, 326-335. Muckle, D. S. and Minns, R. J. J. Bone Joint Surg. 1990, 72-B, 60-62. Messner, K. and Gillquist, J. Biomaterials1993, 14, 513-521 Von Schroeder, H. P., Kwan, M., Amiel, D. and Coutts, R. D. J. Biomed.Mater.Res. 1991, 2 5, 329-339. Shikinami, Y. and Kawarada, H. Biomaterials(submitted). Furukawa, T., Eyre, D. R., Koide, S. and Glimcher, M. J. J Bone Joint Surg. 1980, 62.A, 79-89.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
CALCIUM PHOSPHATE CERAMICS AS CONTROLLED RELEASE SYSTEMS FOR FGF-2 Midy v.* +, Hollande E.+, Rey C*. and Dard M§. * Laboratoire des Mat^riaux et Interfaces. Phosphates et Biomat^riaux, UPRESA CNRS 5071, 38 rue des 36 Ponts, 31400 Toulouse, France ^Laboratoire de Biologie Cellulaire, 38 rue des 36 Ponts, 31400 Toulouse, France § Merck Biomaterial Research, Frankfurter Str. 250,64271 Darmstadt, Germany KEY-WORDS : Calcium phosphate ceramics, Fibroblastic Growth Factor-2, Adsorption. ABSTRAC T The aim of our study was to determine the capacity of calcium phosphate ceramics to serve as carriers for Fibroblastic Growth Factor-2 (FGF-2), which is able to locally stimulate bone formation in vivo. Two different apatitic substrates were tested: stoichiometric hydroxyapatite (HA), widely used as an osteoconductive biomaterial, and nanocrystalline carbonated apatite (CA), very close to bone mineral crystals. Adsorption of the growth factor was quantified by ^^^I and reached 85% of the initial concentration for CA, and 10%, for HA. Release of FGF-2 was about 20%, of the adsorbed amounts for both compounds. In fetal bovine aortic endothelial cell culture, the FGF-2 released showed a strong loss of bioactivity possibly due to the presence of ceramic particles and/or to conformational changes. INTRODUCTIO N Bioactivity is an essential characteristic of several calcium phosphate materials. It involves four steps: surface modification of the ceramic, nucleation of carbonate apatite on the surface of the material from supersaturated body fluid, adsorption of protein, and adhesion and differentiation of bone cells. Althought calcium phosphate materials generally used as biomaterials differ substantially from the mineral part of bone, the precipitation, on their surface, of carbonated apatite crystals very analogous to bone mineral crystals is believed to play a dominant part in the biological response. The relationship between this neoformed layer and bone proteins or growth factors and their action on bone cells however remains largely unknown [1]. It is widely accepted that a multitude of factors, from systemic hormones to local regulatory factors, cytokines, and prostaglandins, act together to regulate the coupling between bone formation and bone resorption. Fibroblastic Growth Factor-2 (FGF-2) is one of the growth factors which can locally stimulate bone formation in vivo [2]. It is produced by osteoblasts and stored in skeletal tissues, and it has been shown to stimulate proliferation of osteoblasts, chondrocytes and periosteal cells [3,4,5]. In vivo studies with FGF-2 also suggest a potential therapeutic role in the treatment of bone loss [6]. Based on these findings, it is speculated that the association of FGF-2 with calcium phosphate ceramics may be useful to facilitate fracture repair. The aim of this report is to quantify the adsorption and release of FGF-2 from two different apatitic substrates, stoichiometric hydroxyapatite (HA) and nanocrystalline carbonated apatite (CA). In additioin the activity of the 591
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FGF-2 released was evaluated. This last point appears particularly important to assess the ability of the calcium phosphate ceramics to act as controlled release systems. MATERIA L AND METHOD S Hydroxyapatite (HA) powder was obtained from Merck KGaA (Darmstadt, Germany), carbonated apatite (CA) powder was prepared as previously described [7]. Recombinant human FGF-2 obtained from Merck Biomaterials (Darmstadt. Germany), was added to the ceramics at various concentrations (0.5-1-2.5-5 Mg/ml). The growth factor in the bound and in the unbound fractions was evaluated by ^^^I counting. The amount of FGF-2 released in a Tris Buffer Salt (TBS) solution was also determined by using ^^^i labeled FGF-2. The in vitroactivity of FGF-2 released after adsorption, from the calcium phosphate ceramics into TBS buffer was determined by using proliferation assay on FBAE (foetal bovine aortic endothelial) cells and on osteoblast-like cells. At the end of the period, the cells were counted using a Coulter Counter and the amount of active FGF-2 was determined by reference to a previously obtained standard curve. RESULTS AND DISCUSSION CA is a calcium-deficient, poorly crystalline apatite. Infra-red spectroscopy showed labile, non-apatitic carbonate and phosphate enviromnents probably located at the surface of the crystals [8], These characteristics are very similar to those of mature human bone mineral. Figure 1 shows the adsorption isotherm of FGF-2 on HA and CA. The adsorption of FGF2 reached 85% of the amount contained in solution for CA, and about 10% for HA. The experimental conditions did not allow saturation of the adsorption sites.
1 . ^ ^3000-
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20
30
Time of release (hour) Figure 1. Adsorption of FGF-2. Grow* factor Figure 2. Release at different time of was incubated at various concentrations (0-0.5FGF-2 after wash from CA and HA. 1-2.5-5 ng/ml) with 5 mg of CA and HA. Adsorption was of 1 ^g/ml of FGF-2. Several physical-chemical considerations may explain the higher efficiency of CA. These apatites exhibit a large proportion of reactive non-apatitic enviromnents of COs^’and lYPO/^’ions that have been shown to be easily exchanged and it has been suggested that they share the same adsorption sites as proteins at the surface of the crystals. Moreover, the surface irregularities on CA crystals, may facilitate protein attachment by exposing a large number of adsorption sites.The
Calcium Phosphate Ceramics as ControlledRelease Systemsfor FGF-2: V. Midy et al.
593
specific surface areas of the two apatitic substrates was also seen to be different (15 m^.g"^ for HA, andl56m^.g-l forCA). The release of FGF-2 is shown in figure 2. FGF-2 release from CA ranged from 17 to 27% of the adsorbed amount depending on the rinsing time. HA gave a similar relative amount released, nevertheless a much larger absolute amount of growth factor was released from CA than from HA. It is noticeable that most of the adsorbed growth factor was not released into the buffer solution we used and appeared to be strongly bound to the 2q)atitic substrate. 4>
relFGF-2/CA relFGF-2/HA
> a
So
2.5
’Z S
^
*s "^
1.5
9i U
a o O
0
0,5
4-
0
10
’ ^ ^ U ^ ’^
-relFGF-2/HA -relFGF-2/CA 20
10
30
20
t
30
Time of release (hour) Time of release (hour) Figure 3. Bioactivity of FGF-2 on FB AE Figure 4. Bioactivity of FGF-2 on osteoblast-like cell proliferation from 0 to 24 hours of cell proliferation from 0 to 24 hours of release. release. Adsorption was of 1 |xg/ml. Adsorption was of 1 pig/ml. The biological activity of FGF-2 released from the substrates, was tested by FBAE cell proliferation to detect any variation of the growth factor properties (Figure 3). The concentration of active FGF-2 released was found to decrease with time, from 2.6 to 1.7 ng/ml for CA, and from 1.8 to 0.83 ng/ml for HA. The same experiment was performed on osteoblast-like cells (figure 4) -’-’ ’ ""» »
FGF-2/HA -FGF-2/CA 4-
-h 0
1
2
3
4
5
Initial concentratio n of bFGF (ng/ml) Figure 5. Bioactivity of FGF-2 proliferation at different concentrations (0-0.5-1-2.5-5 ^ig/ml), after 24 hours of release.
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Bioceramics Volume10
and led to a similar curve indicating that the loss of biological activity was not dependent on the cell type, or on the cell type receptor. The amount of active FGF-2 released was also found lo depend on the amount initially adsorbed from solutions at different concentrations (Figure 5). The amount of active factor released increased with increasing initial concentrations of growth factor. A drastic decrease of the activity of FGF-2 appeared after the adorption-release process. Several mechanisms might be proposed to explain this phenomenon. Degradation products of Uie ceramics, resulting from the contact with culture medium and the presence of cells, led to debris of various sizes which could strongly inhibit cell proliferation [9]. The decrease of biological activity of the growth factor after adsorption-release appeared much stronger for CA than HA and this observation could be related to the higher degradability of CA . This factor cannot however explain the decrease of activity observed on HA as this sample only released very little debris. Physical-chemical changes of the proteins may also occur due to ionic interactions during adsorption of the growth factor. Such modiHcations have been found, for example in TGF-P2 [10]. This behaviour seems hovewer inapplicable for FGF-2 as this growth factor can be separated and purified by chromatography on HA. The loss of activity could nevertheless be assigned to alterations following adsorption as it showed a relative increase with the release time. CONCLUSIO N These studies show that apatite characteristics, in particular poorly crystalline apatites, greatly influence FGF-2 binding and release. Although the released growth factor was found to be essentially inactive, it should be pointed out that most of the adsorbed fraction was not released and that the activity of this part has yet to be determined. The mode of alteration of FGF-2 on apatites has to be analysed before envisaging any future applications of growth factor-apatite associations for bone repair. ACKNOWLEDGEMEN T The present work was supported by the Centre National de la Recherche Scientifique (Grant n^l41053500-CNRS9543) REFERENCE S 1. Rey, C , Biomaterials,1990,11, 13-15. 2. Frenkel, S., Herskovits, M.S., and Singh, I.J., Acta Anat,1992,145, 265-268. 3. Canalis, E., Centrella, M., and Mc Carthy, T., J, Clin, Inves.,1988, 81,1572-1577. 4. Globus, R.K., Patterson-Buckendahl P., and Gospodarowicz, D., Endocrinology,1988, 23, 98-105 5. Hauschka, P.V.,. Mavrakos, A.E., lafrati, M.D., Doleman, S.E., and Klagsbrun, M., /. Biol. Chem..1986, 261, 12665-12674 6. Kawaguchi, H., Kurokawa, T., Hanada, K.,Himaya, Y., Tamura, M., Ogata, E., and Matsumoto, T., Endocrinology,1994,135, 774-781. 7. Rey, C , Renugopalakrishnan, V., Collins, B., and Glimcher, MJ., Calcif. Tissue Int,1991, 49, 251-258. 8. Midy, V., Rey, C , Bres, E., and Dard, M., submitted 9. Alliot-Licht, B., Gregoire, M., Orly, I., and Menanteau, J., Biomaterials,1991,12, 752-756. 10. Griffith, D.L., Keck, P.C, Sampath, T.K., Rueger, D.C., and Carlson, W.D., Proc. Natl Acad. Sci. USA, 1996, 93, 878-883.
Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd
C-SR C ONCOGEN E mRNA EXPRESSIO N IN POROU S HYDROXYAPATIT
E
CERAMIC S K. Mishima^^ H. Ohgushi^\ T. Yoshikawa^\ H. Nakajima^^ E. Yamada^^ S. Tabata’^^Y. Dohi^^ and K. Ichijima’^ Departments of^^Pathology,^^Orthopaedic Surgery,^^Chemistry, and^^Public Health, Nara Medical University Kashihara City, Nara 634, Japan.
ABSTRAC T The C-src oncogene is a tyrosine kinase and has an important role in bone resorption. To investigate the role of C-src in osteogenic response of porous hydroxyapatite(HA), we utilized the experimental model which can show de novo bone in rat marrow/HA composite. HA without marrow and marrow/HA composites were implanted at rat subcutaneous sites and total RNAs were extracted from the implants for Northern blot analysis. Both HA and composites showed C-src mRNA expression at 2 weeks post implantation, and the signal of the mRNA in the composite was much higher than that in HA. These results indicate the important role of tyrosine kinase in the osteogenic marrow/HA composites. KE Y WORDS : C-src, hydroxyapatite, marrow, implantation.
INTRODUCTIO N The C-src is a non-receptor tyrosine kinase and plays an important role in signal transduction of cells and it is expressed ubiquitously but with elevated levels in neurons and platelets[l,2]. Recent reports indicate its important role in bone metabolism. The mice having impaired C-src gene show osteopetrosis because osteoclasts, which express C-src at high level, lose the bone resorption ability[3]. It is also reported that C-src is expressed in bone forming cells(osteoblasts). Therefore, it is very important to investigate the role of c-src in bone formation process. We have reported that neither the marrow cells nor the HA by themselves show bone formation at subcutaneous sites , whereas composites of HA and marrow cells result in consistent bone formation which starts directly on the HA surface[4]. In the present report, expression of the C-src gene in the osteogenic marrow/HA composites was determined by Northern blot analysis. 595
596
Bioceramics Volume10
MATERIAL S and METHOD S Implantation of marrow/HA composite s Detailed procedures for marrow cell preparation have already been reported[57]. Briefly, marrow from the femoral and tibial diaphysis of Fisher rats were hydrostatically forced into centrifuge tubes containing heparinized phosphate buffer saline(PBS). The marrow was disagregated and concentrated by centrifiigation resulting in 5x10^ nucleated cells/mL. Porous hydroxyapatite ceramic(HA) discs (the average pore size is 230 nm in diameter and average void volume is 50-60%, the disc shape is 5 mm in diameter and 2mm in thickness; Interpore International, Irvine CA,) were soaked in the marrow cell suspension. HA(without marrow) and HA soaked in the cell suspension were subcutaneously implanted and harvested at 2 and 4 weeks post implantation. The harvested implants were used for decalcified histological section and RNA extraction. Subcloning of C-src cDNA in rat Total RNA fraction was extracted by the acidic guanidinium thiocyanatephenol chlorofonn (AGPC) method from rat cancellous bone and mRNAs from the RNA fraction were isolated with Oligotex^^-dT30 (Roche, Japan). First strand cDNA was synthesized from the mRNAs by reverse transcriptase (first strand cDNA synthesis kit, Amersham, International pic, England). cDNA fragment of c-src was amphfied using the first strand cDNA by polymerase chain reaction(PCR). Oligonucleotide primers for the PCR were synthesized according to the sequence of c-src cDNA in mice(EMBL M17131). Amplified fragment of c-src cDNA was inserted into a pCR vector (Invitrogen Corporation, USA) and the DNA sequence was performed with 373A DNA sequencing systems(Applied Biosystems,USA). Northern blot analysis After homogenization of the harvested implants, total RNA was extracted by AGPC method and separated on a 1.2% agarose gel containing 20mM MOPS, 5mM sodium acetate, ImM EDTA, and 6% fonnaldehyde. After gel electrophoresis, RNA was transferred to a nylon membrane (Immobilon N, Millipore Ltd., Japan) in 10 x SSC and cross-linked with ultraviolet light. Prehybridization and hybridization of the blots were carried out at 68 C in QuickHyb^^ solution (Stratagene, CA) for 20 and 120 min, respectively. Hybridization was done in the presence of a ^^P-labeled cDNA probe for the subcloned rat c-src. A cDNA probe was labeled with ^^P using a multirandom primer labeling kit( Amersham, Japan) and [^^P] dCTP. RESUL T AND DISCUSSIO N Subcloned cDNA fragment of the rat c-src was 453 bp and had 96% homology to mouse C-src cDNA of peripheral form. The cDNA fragment lacked 18 bp of neurospecific insertion reported in mouse C-src cDNA(Fig. 1). The C-src mRNA expression was detected in both HA(without marrow) and the marrow/HA composites at 2 weeks post implantation. The intensity of the C-src mRNA signal in the composite was higher than that in HA. At 2 weeks post implantation, we reported that composite can show bone specific mRNA(osteocalcin) signal whereas the HA (without marrow) do not show the signal[4] nor show bone formation[4-7].
C-SRC OncogenemRNA Expression in Porous Hydroxy apatiteCeramics: K. Mishima et al.
597
i TCAAGAAAGGGGAGCGGCTGCAGATTGTCAATAACAC AGAGGGAGA CTGGTGGCTGGCACACTCGCTGAGCACCGGACAGACCGGTTACATCCC CAGTAACTATGTGGCGCCCTCCGACTCCATCCAGGCTGAGGAGTGGTACTT TGGCAAGATCACTAGACGGGAATCAGAGCGGCTACTTCTCAACGCCGAGA ACCCCAGAGGGACCTTCCTCGTGAGGGAGAGTGAGACCACAAAAGGTGCC TACTGCCTCTCTGTATCCGACTTTGACAATGCCAAGGGCCTAAATGTGAAA CACTACAAGATCCGGAAACTGGACAGTGGCGGACTCTACATC ACCTCCCGCACGCAATTCAACAGCCTGCAGCAGCTTGTGGCTTACTACTCC AAACATGCTGATGGCCTGTGTCACCGTCTCACTACCGTGTGTCCCACATCC AAGCCTCAGACCC Fig. 1. The sequence of fragment of c-src cDNA. It lacks specific insertion in mice.(arrow shows the position of insertion)
18 bp of
neuron-
Therefore we suppose the high intensity of the C-src mRNA in the composite is attribute to the osteogenic function resides in the composite. As shown in Fig. 1, some multinucleated giant cells could be observed in both HA and marrow/HA implantation. In this regard, osteoclasts is multinucleated cells and can resorb both bone and biological HA[8], and the osteoclasts show the high signals of C-src expression. Therefore, there is a possibility that the C-src mRNA expression of HA(without marrow) is partly derived from the multinucleated giant cells on the HA. Present data of Northern blot analysis were obtained from extracted RNA fractions, and thus identification of the gene expression of each cell was not clear. Therefore, to confimi the role of C-src at cellular level, an another analytical approach such as in situ hybridization technique might be needed.
28S 18S
Fig. 2. Northern blot analysis of mRNA for c-src. + : marrow/HA composites at 2 weeks post implantation : HA without marrow at 2weeks post implantation
598
Bioceramics Volume10
3L
Fig. 3. Photomicrograph (x200 originally) of the composite of porus hydroxyapatite(HA) and marrow, a: 2 weeks after implantation. Obvious bone formation does not appear at this time and fibrovascular tissue fomiation emerges in the pore areas. A few multinucleated cells are found(arrow), b: 4 weeks after implantation. Abundant bone fonnation together with many osteoblasts and osteoclasts.
REFERENCES 1. Levy, J.B., and Brugge, J.S. Mol. Cell. Biol. 1989, 9 , 3332-3341. 2. Golden, A., Nemeth, S. P., and Brugge, J. S. Proc. Natl. Acad Sci.USA 1986, 83,852-856. 3. Lowe, C , Yoneda, T., Boyce, B.F., Chen, H., and Mundy, G.R. Proc. Natl. Acad. Sci. USA 1993, 90, 4485-4489. 4. Ohgushi, H., Dohi, Y., Tamai, S., and Tabata, S., J. Biomed Mat. Res. 1993, 27,1401-1407. 5. Ohgushi, H., Okumura, M., Tamai, S., and Shors E. C. J. Biomed. Mat. Res. 1990,24, 1563-1570. 6. Okumura, M., Ohgushi, H., and Tamai, M. Biomaterials, 1991, 12, 411-416. 7. Okumura, M., Ohgushi, H., and Tamai, M., and Shors, E.C. Cells and Materials 1991,1, 29-34. 8. Yamada, S., Nakamura, T., Kokubo, T., Oka, M., and Yamamuro. J. Biomed. Mater. Res. 1994, 28, 1357-1367.
AUTHOR INDEX
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Index of Authors Afonso, A. Aguado, E. Aitasalo, K. Akahane, M. Akao, M. Anselme, K. Antolotti, N. Apfelbaum, F. Araki, N. Arbore, L. Asaoka, N. Aurelle, J. L. Avila, G. Azumi, R.
489 71 429 97, 109, 117 229, 407 583 455 397 91 379 447 275 553 545
Caccavale, F. 287 Cales, B. 135, 271 Cannas, M. 139 Capello, W. N. 157 Carl, G. 195 Catelas, I. 579 Cavaiheiro, J. 489 Chai, C. S. 175 Chamberland, D. L. 349 Champion, E. 75, 549 Chang, Y. S. 147 Charissoux, J.-L. 113 Chen, X. F. 391 Chevalier, J. 135, 271 Cho, S. B. 407, 545
Bakki,B. Baquey,Ch.
53 383
Su’^’^e’e ? " " « ’ f ’ ^-
if’o^ ?^l261
Basse-Cathalinat, B. 383
ClemenU-
553
Ben-Nissan, B. 175 r. u r> i-tc Benaben, P. 275 Beneytout,J.-L.113 Berger,G. 53,367
/r, XT Correia, R. N. <- u . r. Couchourel, D. Cournot, G.
/ii? 415 oi 23 571
^""I"
415, 515
D
XT-
D
T^
Bernache-Assollant, D. 75, CAQ
Conte, P.
113, 523,
Bertini, S. 455 447 Best, S. Best, S. M. 19 Betchem, F. 179 Bevis, M. J. 515 Bizot, P. 131 Boltong, M. G . 279, 481 Bonel, G. 143 Bonfield, W. 15, 19, 447, 519 Bosetti, M. 139 Bouler, J.-M. 23, 27, 71 Branco, R. 489 Bres, E. 571 61 Brink, M. Buttery, L. D. K. 105
I’
Cunha, A . M .
493
. ,.
,
D’Antonio, J. A. 157 Daculsi, G. 23, 27, 71, 79, 191 Dard, M. 591 De Santis, E. 467 Delecrin, J. 27, 79, 191 Deloge, G. 161 Demonet, N. 275 Descamps, M. 387, 583 Di Silvio, L. 19, 519 Dohi, Y. 97, 595 Dolcini, L. 503 Doria, C. 467 Dorozhkin, S. V. 187 Driessens, F. C. M. 279, 481
Drouin, J. M. Ducheyne, P. Dupraz, A. Dupuy, B.
135, 271 45, 101 191 383
Eckert, K.-L. 3,451 Edidin, A. 157 Edidin, A. A. 317 Epinette, J. A. 157 Ezhova, Z. A. 237 Pages, J. 493 Falaize, S. 45 Fernandes, M. H. 415 Fernandez, E. 481 195, 199 Ferraris, M. Filser, F. 433 Flautre, B. 387, 583 Forest, B. 275 Forest, N. 219 Frayssinet, P. 143,371,379,493 Freche, M. 375 Freeman, M. A.. R. 305 Fricain, J. C. 383 Fujishiro, Y. , 283 Fujita, H. 497 Gallur, A. Galy-Fourcade,, Gatti, A. M. Gauckler, L. Gauthier, O. Gautier, S. Geesink, R. G. Gibson, I. R. Gil, F. J. Gildenhaar, R. Ginebra, M. P. Gineste, L. Gisep, A. Gonella, F.
Greenspan, D. Gross, U. Gross, K. A. Gruner, H. Ha, S.-W. Hack, G. D.
387, 583 D. 507 287 433 71 549 157 19 553 367 481 143, 371, 493 203 287
C. 265,391,411 53 175 203, 451 3, 203, 451 411
Hamadouche, M. 131 Hamagami, J.-i .463 Hamagami, J. 207 Hamanishi, C. 419 Hammond, A. 45 7 Handa, S. 337 Hang, Y. S. 291, 295, 471 Hara, T. 387, 583 Hardouin, P. Harmand, M. F. 455 Hartmann, P. 57 Hasegawa, M. 587 Hastings, G. W . 511 Hatim, Z. 179, 375 Hayakawa, S. 33, 41, 527 Hayashi, K. 291, 295, 471 245, 283, 287, 541, 557 Hench, L. L. 19, 575 Hing, K. A. Hoellrich, R. G\. 349 Hohmann, D. 169 HoUande, E. 591 Hong, C. Y. 337 83, 261 Hong, K. S. Hosoda, K. 97 Huang, J. 519 Huk, 0 . L. 579 Hukkanen, M. 105 Ichijima, K. Ido, K. lida, H. Ikeda, N. Imamura, T. Imayoshi, N. Imoto, K. Inoue, M. Ishida, H. Ishii, Y. Ishikawa, K. Itoh, F. Iwaki, H. Iwaki, Y. Iwamoto, Y.
109, 117,595 497 87, 357, 497 313, 357 291 41 333 531 233 345 301 215 403 333 291, 295, 471
Jana, C. 195 Jha, L. J. 19 Johnson, G. S. 321 Joseph, R. 15
Kadoya, Y. Kageyama, Y. Kahana, F. Kakutani, Y. Kameyama, T. Kato, K. Kawamoto, Y. Kawanabe, K. Kawashita, M. Khairoun, I. Khor, E. Kijima, K. Kikuchi, M. Kikutani, T. Kim, C.-E. Kim, C. Y. Kim, D.-J. Kim, H.-M. Kim, Y. H. Kin, N. Kita, Y. Kitamura, Y. Knowles, J. C. Kobayashi, A. Kobayashi, M. Kobayashi, T. Kocher, P. Kokubo, T.
305, 325 253 397 357 105, 329 91, 329 105, 329 87, 459 531 279 511 241 407 313 211 37 211 11,215, 561 153, 309 333, 403 253 497 575 305, 325 313, 497, 561 407 433 7, 11,215,219,253, 257, 313, 459, 531, 561 Koval, E. M. 237 407 Koyama, Y. 165 Krahl, H. Krajewski, A. 139, 195, 199, 503 Kushitani, S. 333, 403 Kwon, S. 37 Lacout, J. L. 375 Laine, P. 61 LaTorre, G. P. 391 Latour Jr, R. A. 541 Leboutet, M.-J . 113 83, 261 Lee, C. K. 211 Lee, M.-H. 387 Legrand, O. 477 Lemaitre, J. Lerch, A. 371, 493 571 Leroy, G. 57 Leuner, B.
Liagre, B. Lickfield, G. C. Lin, F. H. Litkowski, L. J. Liu, H. C. Lobel, K. D. Lopes, M. A. Lopez, J. Lorier, M. A. Loty, C. Loty, S. Lowery, G. Lu, J. X. Luthardt, R. Luthy, H.
113 541 337 411 337 557 575 565 321 219, 477 219 65 387, 583 437 433
Mandai, Y. 361, 403, 485 Mangano, C. 503 Manley, M. T. 157,317 Manunta, A. 467 Marchand, R. 579 Marie, P. J. 523 Martinetti, R. 503 553 Martinez, S. 295 Mashima, T. 419 Matsuda, N. 497 Matsuda, Y. Matsumoto, M .545 Matsushita, K. 253 203 Mayer, J. Mazzocchi, M. 139 McLoughlin, S . W. 349 565 Mendez, J. 565 Mendez, M. 179 Michaud, P. Midy, V. 591 Minamigawa, K. 361, 403 Mishima, K. 109, 595 Miyaji, F. 7, 11,215,257, 531 561 Miyamoto, Y. 301 Miyazaki, T. 11 Moisescu, C. 199 Monari, E. 287 Monteiro, F. J., 49, 575 Morita, S. 535 Moroni, A. 455 Mucalo, M. R. 321
Mulier, M. Miiller-Mai, C. Murakami, H. Murata, N. Musil, R.
161 53 357 333 437
Nagata, F. Nagatomi, K. Nagayama, M. Nakahira, A. Nakajima, H. Nakamura, S. Nakamura, T.
105, 329 403 301 241 109, 117, 595 229 7, 11, 87, 147, 207, 215, 257, 313, 357, 459, 497, 561 Nakashima, Y. 471 Nardin, M. 523 Narva, K. 61 Nguyen, J. M. 79 Nishiguchi, S. 215, 561 Nishizawa, K. 105, 329 Nizard, R. S. 131 Nordstrom, E. 223 Ohashi, H. Ohgaki, M. Ohgushi, H.
305, 325 229 97, 109, 1117,223,233, 595 Ohta, T. 97 Ohtsuki, C. 19, 33, 4]I, 527 Ohura, K. 419 Ohzeki, K. 345 Oka, M. 147, 357, 497 Okada, K. 329 Okada, Y. 87, 313 Okazaki, M. 241 Okumura, M. 233 Okuyama, M. 329 Olsen, I. 575 Oonishi, H. 283, 333, 361, 403, 485 Orienti, L. 455 Orlovskii, V. P . 237, 401 Osaka, A. 41, 53, 527 Padrines, M. Paracchini, L. Park, J. S. Park, M. R.
23 195 309 309
Park, T. S. Passuti, N. Pedra, A. Peltola, M. Penel, G. Pieper, H.-G. Pilet, P. Pitto, R. P. Planell, J. A. Platzbecker, U. Plitz, W. Ploska, U. Polak, J. M.
309 191 445 429 571 165 71, 191 169 279, 481, 553 57 127 367 105
Quack, G.
165
45, 101 Radin, S. Railhac, J. J. 507 Ranz, X. 455 Rastellino, M. 139 Ravaglioli, A. 139, 195, 199, 503 Raynaud, S. 75 Redey, S. A. 523 Refior, H. J. 127 Reif, D. 57 Reis, R. L. 415, 515 Revell, P. A. 305 Rey, C. 455, 523, 591 Rieger, W. 437 Rieu, J. 275 Rinonapoli, G. 467 Ritter, S. 3 Rodriguez, F. 179 Rohanizadeh, 1R. 23, 27., 191 Rouquet, N. 143, 371, 379, 493 Royer, J. 79 Ryu, H. S. 261 Sakamoto, K. Salinas, A. J. Sano, T. Sans, N. Santin, M. Santos, E. M. Santos, J. D. Sarig, S. Sasaki, S. Sautier, J. M.
241 245 535 507 139 101 49, 575 397 345 219
Sbernardori, M . C. Scharer, P. 433 425 Schepers, E. Schulze, K.-J. 57 305 Scott, G. Scrivani, A. 455 Sedel, L. 131, Sekine, Y. 207 Serret, A. 245 Shapiro, I. 101 Sharrock, P. 447, Sheaffer, H. B. 411 Shenker, B. 101 Shibuya, T. 535 Shikinami, Y. 587 Shimoya, I. 249 Shinto, Y. 91 Shinzato, S. 313 Shon, J. H. 153 49 Silva, P. L. SHvka, 0 . I. 401 Sonoda, J. 535 Stefani, Y. 271 Stokes, K. E. 349 Sudo, A. 587 Suetsugu, Y. 249, Sugihar, F. 485 Sugihara, F. 403 Sun, J. S. 337 Suonpaa, J. 429 Sutter, B. 387 Suzuki, K. 253, Suzuki, T. 105, Suzuki, Y. 531
467
523
507
Thierry, B. 387, 583 Tian, Y. 541 Toda, T. 361, 485 Toguchida, J. 147 Toriyama, M. 105 Trecant-Viana, M. 27, 79 Trembley, S. D. 541 Trunec, M. 183 Tsuang, Y. H. 337 Tsuji, E. 403 Tsuru, K. 33 Uchida, A. 91, 535, 587 Uenoyama, K. 295 Umegaki, T. 207, 463 Ushio, K. 147 Vallet-Regi, M. 245 Van Landuyt, P. 477 Vargas, G. 565 Vasconcelos, M. 489 Verne, E. 195, 199 Viola, G. 455 Vogel, J. 57, 195 Voigt, C. 53
407, 545
301, 341 215
Tabata, S. 595 Tahara, Y. 345 Takadama, H. 257 Takakuda, K. 407 Takano, I. 345 Takaoka, G. H .531 Takashima, S. 527 Takegami, K. 535 97, 1109, 117,223,233 Tamai, S. Tanaka, J. 249, 407, 545 Tanaka, S. 419 Tanner, K. E. 519 Terradas, R. 553
Wakabayashi, H. 535 Wakitani, S. 333 Walter, A. 127 Wan, A. C. A. 511 Wang, M. 15, 519 West, J. K. 541, 557 Wheeler, A. P. 541 Wheeler, D. L. 349 Willfahrt, M. 367 Willmann, G. 123, 165, 353 Wilson, J. 65 Wintermantel, E. 3, 203, 451 Yahia, L’H. Yajima, H. Yamada, E. Yamada, I. Yamada, M. Yamaguchi, S. Yamakawa, T. Yamamuro, T. Yamano, Y.
579 233 595 531 341 241 91 357 305, 325
Yamashita, K. Yamashita, Y. Yamazaki, T. Yan, W.-Q. Yli-Urpo, A. Yokobori, T. Yokogawa, Y. Yonehara, E. Yoon, H. J. Yoshii, S. Yoshikawa, M.
207, 463 91 535 459, 561 61, 433 223 105, 329 463 83 357 361, 485
Yoshikawa, T. Yoshinari, H. Yoshino, A. Youn, H.-J. Yura, S. Yutani, Y.
97, 117, 109,233,595 223 253 261 147 325
Zahraoui, C. Zhong, J. Zhong, J. P.
507 265 391
KEYWORD INDEX
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Index of Keywords A
bioactive glass (BG)
a-tricalcium cement
229, 241, 361, 397, 403, 407 87, 497 A-W glass ceramic 357 acetabular dysplasia ACP (amorphous calcium phosphate) 253 coating adhesion 49, 275, 337, 523 adsorption 527, 541, 557, 591 Ag-doping 329 ageing 135, 175, 261 aliphatic alcohols 241 alkali treatment 561 alkoxide 175 all-ceramic crown 211 all-ceramic dental bridge 433 all-ceramic wear couple 165 all-ceramics-systems 437 allogeneic bone marrow1 97 123, 131, 237, alumina 271, 313, 325, 463, 549 alumina head 333 alveolar bone 361 AMI 245 animal experiments 57 antibacterial property 329 antibiotics 87,91 apatite 7, 11, 33,41, 187, 215, 257 apatite-collagen conjugate 403 arc sprayed titanium 291 arc spraying 291 artificial articular cartilage 147, 587
B
BG-bone interface
429
bioactive bone cement bioactive fillers bioactivity bioceramics biocompatibility biocomposites biodegradable biodegradable glasses biodegradation Bioglassfi biological property biomaterials biomechanical testing biomimetic process bioresorbable cement bioresorption biphasic ceramics Bis-GMA blood compatibility blood purification blood vessel bonding bone bone-analogue bone bonding bone calcination bone cement bone formation bone grafting bone graft substitute bone growth bone ingrowth bone remodelling bone replacement
37, 45, 199, 283, 287, 415, 425, 429 497 503 11,33,41, 101 215, 561 105, 223, 387, 583 19, 287, 575 195 415 553 493 265, 349, 391, 411,519 313 571 349 7, 257, 459 419 403 375 313, 497 527 527 253 261, 471 53, 321, 387 357, 515 459, 561 83, 565 117,481,535 233 109, 353 169 265 147, 295 147 415
bone stock bone substitute bone union bonelike crystal growth
c
169 71, 383 341 463
595 261 367 437 383 379 3, 23, 27, 71, 79, 329, 481, 527, 545, 565, 571 calcium phosphate cement 279, 301, 419, 493 calcium phosphate ceramics 367, 371, 591 calcium phosphate coatings 451 calcium phosphate debris 143 calcium phosphate glass target 207 cancellous bone defect 349 canine 497 CaO-Na20-P205 system 553 carbon fiber reinforced PEEK 203, 451 carbonation 523 catheterization 507 cell culture 101, 407, 519, 571, 575 cell differentiation 101 cells 223 cellulose 329 cement 179, 229, 241, 361, 397, 403, 407, 477, 489 cementless prostheses 301 ceramic 71, 127, 169, 275, 345, 579 ceramic composites 49 ceramic fracture 305 ceramic particles 305 ceramic sheet 211 ceramic surface 219 characterisation 175, 229 chemical mechanism 187 chemical treatment 215
C-src Ca/P ratio Ca2KNa(P04)2 CAD/CAM calcite calcium oxide calcium phosphate
chitin 511 chitosamine 507 chondrocyte 219 chondroitin-sulphate 48 5 Cis-diamminedichloroplatinum 345 clinical study 153, 157 co-polymerized polylactide 407 coagulation parameter 527 coating 175, 275, 291, 445, 447 collagen 401 complexing agent 447 composite 15, 329, 313, 401, 407, 415, 511, 515, 519, 549 composite coating 463 computational chemistry 541 computer mirror-image 161 coprecipitation 237, 401 coral 383 covercoat 37 crowns and bridges 437 cryopreservation 109 crystal nucleation 545 cytocompatibility 379 cytotoxicity 229, 519
D
defatting degradation dental ceramics dental implants dentin occlusion dentin dentistry deproteination dexamethasone
321 53, 265, 391 211,287 353, 445 411 411 437 321 109, 117
dicalciurn phpsphate dihydrate 371 differentiation 219 dissolution 45, 75, 187, 275, 375 drug delivery system 91, 345 dual energy absorptiometry (DXA) 387
E
EDX electrolysis
257 447
electron microscopy 191 electrophoretic coatings 463 enhancement of biocompatibility 397 extracellular matrix 523
F fat tissue 253 fatigue 75 femurs 471 ferromagnetic thermoseed 535 fibroblastic growth factor-2 591 fibrovascular tissue 429 flow cytometry 575, 579 fluorapatite 19, 113 fluorhydrohyapatite 455 follow-up 131 FTIR/NMR 321
G
glass and glass-ceramic coatings 199 glass 41, 65, 489 glass-ceramics 53, 357, 415 glass coating 139 glass fibre 61 glass reactions 61 glycerin 485 grain growth 183 green machining 433 grinding of zirconia-TZP 437 groundcoat 37 growth factor 101 guided tissue regeneration 407 guided bone regeneration 53, 407
H
hardening heat treatment hip arthroplasty
179 11, 561 305, 317, 467, 497 hip joint 123, 131, 271 hip prothesis 127, 153, 161, 169, 325, 333 hip prosthesis with spongiosatype metal surface 165 histological evaluation 587 histomorphometry 387 human frontal sinus obliteration 429 humeral pseudoarthrosis 341
hydroxy-carbonate apatite 249, 391, 565 15, 19, 37, 75, hydroxyapatite 83, 91, 97, 105, 109, 113, 117, 157, 175, 179, 203, 223, 229, 233, 237, 241, 261, 283, 291, 321, 353, 371, 375, 379, 401, 447, 463, 467, 481,511, 515, 519, 523, 549, 565, 575, 595 hydroxyapatite cement 507 hydroxyapatite ceramics 183 hydroxyapatite coating 143, 161, 207, 291, 295, 305, 309, 317, 353, 471 445 hydroxyapatite-silica hydroxyapatite-tricalcium phosphate 341 hyperthermia 535
I
341 Manufactured 161 157 287, 291, 309, 371, 425, 459, 489, 493, 595 321 implants 45, 219, 391 in vitro in-vitro biocompatibility 139 in vivo 79, 349 inflammation 301 insert surface 143 53 interface interleukin 113 531 ion implantation
iliac autograft IMP ( = Intra-operativ<; Prosthesis) system implant fixation implantation
K
kinetics
183
L lactic acid 507 Langmuir-Blodgett monolayer 545 lipoxygenase 113 load-bearing condition 147 loosening 325 low crystalline apatite 403 low temperature ceramic 397 low temperature crystallization 207 lysine 541
M macrophages 485, 579 marrow 109, 595 117,233 marrow cells 33,41 MAS-NMR material degradation in vivo 191 mechanical properties 79, 515, 553 mechanical resistance 179, 375 mechanical strength 75, 223, 477 mechanical testing 199, 459 metal implants 353 metal particles 325 metallosis 305, 325 Micro Raman spectrometry 571 mixing liquid composition 477 MO modelHng 245 modulus of elasticity 317 molecular modeHng 541, 557 monocalcium phosphate monohydrate 477 morphology 15 multi-nuclear giant cells 485
N N-carboxymethyl chitosan coating 503 NaOH treatment 3 near-equihbrium drying 265 non-coated devices 143, 161 nucleation 257
o obliteration material 429 octacalcium phosphate 403 organically modified silicates (ORMOSILs) 33 osseointegration 353, 467 osteoarthritis 113
osteoblasts osteocalcin osteochondral defect osteoconduction osteomyelitis osteoproduction ovariectomized rates
105, 337, 523, 583 97 587 283, 313, 455, 467, 471 87,91 283 471
p particles 579 particle size 15 particulate calcium phosphate 191 periosteum 233 phase transformation 135 phosphate glasses 57 phosphate ions 245 phosphorus 531 plasma-spraying 49, 229, 275, 379, 455 PM3 245 PMMA 489 polarized IR spectroscopy 249 polyethylene 15, 515 polyethylene cup thickness 333 polyethylene cups 143 polymers 415 porosity 57, 71, 147, 179, 367, 583 porous apatite 153, 463, 503 porous block 87 porous ceramic 371 porous coating 301, 317 porous interconnections 583 precipitation 23, 27 prostheses fixation 279 prosthesis 433, 467, 571 proteins 541, 557 proximal interface enhancements 317 pseudoarthrosis model 419 push-out test 291, 471 pyrost bone 337
R r.f. magnetron sputtering 207 rabbit 83 radioactive tracers 275
radiotherapy 531 range of motion 341 rapid prototyping 433 recombinant human bone morphogenetic protein-2 419 357 reconstruction resorbable compounds 57, 367 resorption 61, 383 revision 169 root canal sealer 361
S second phase SEM SEM-EPMA setting time sheep silica
261 519 143 477 467 557
silica-alumina composite 527 silicon 557 simulated body fluid (SBF) 7, 11, 33, 41, 215, 257 sintering slurries
183, 195, 261 549
sodium silicate sodium tantalate soft tissue sol-gel
7 11 61
solubility tests solution processing solvation effects
33, 175, 265, 349, "IQI 57, 455
511 245
space group 249 spinal fusion 65 spreading 337 stress shielding 317 subcutaneous tissue 485 subcutaneously implantation 301 surface macrotexture 301 surface modification 45, 455 surface structure 215
207, 295, 451, 471 titanium implants tooth hypersensitivity 411 total knee replacement 353 549 toughness 241 transformation 41 transition metal oxide 123, tribology 105, tricalcium phosphate 229, tricalcium phosphate, a 403, tricalcium phosphate, P 477
u U H M W P E implants ultrastructure
11 257 83 403
215, 291, 309, 447, 561
127 261, 291 241, 397, 407
143 23, 27, 53
V vacuum plasma spraying
ly^, 2U3, 4^1
volume of mixing liquid 477
w wear
123, 135, 271, 325 wear of polyethylene cu p 333
X X-ray diffraction
249
Y Y203-Al203-Si02 glass 531
Z
T tantalum TEM tensile strength tetracalcium phosphate
tetracalcium phosphate cement 361, 485 thermal decomposition 183 tissue culture 105, 253 61 tissue response 3 titania ceramics 37, 65, 203, titanium
zeta potential zirconia
545 123, 135, 139, 237, 271, 433, 445
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