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ANNÉE 2014
THÈSE / UNIVERSITÉ DE RENNES 1
sous le sceau de l’Université Européenne de Bretagne
pour le grade de
DOCTEUR DE L’UNIVERSITÉ DE RENNES 1
Mention : Chimie
Ecole doctorale Sciences de la Matière, Rennes
présentée par
Mostafa Mabrouk Mohamed
préparée à l’unité de recherche : UMR CNRS 6226 Sciences Chimiques de Rennes
Composante universitaire : S.P.M
Preparation of
PVA/Bioactive Glass
nanocomposite
scaffolds. In vitro
studies for
applications as
biomaterials.
Association with
active molecules.
Thèse soutenue à Rennes le 11 juin 2014
devant le jury composé de :
Hicham BENHAYOUNE Professeur, Université de Reims, France / rapporteur El-Sayed Mahmoud EL SAYED Professor, University Ain Shams, Egypt / rapporteur Mohamed EL GOHARY Professor, University Al Azhar, Egypt / Examinateur Sylvie JEANNE Professeur, Université de Rennes 1 / Examinateur Amany MOSTAFA Professeur, National research Centre (NRC) , Egypt, Co-Directeur Hassane OUDADESSE Professeur, Université de Rennes 1 / Directeur de Thèse
Statement of original authorship
i
Statement of original authorship
The work contained within this thesis has not been previously
submitted for a degree or diploma at any other higher
institution. To the best of my knowledge, and belief, the thesis
contains no materials previously published or written by
another person except where due reference is made.
Mostafa Mabrouk
ACKNOWLEDGMENT
ACKNOWL EDGMENT
I would like to express the deepest gratitude to
Prof. Hassane Oudadesse Professor and Head of
Biomaterials group - the University of Rennes 1,
SCR, UMR CNRS 6226, France for suggesting the point
of this search and for his continuous advice and
enhancement throughout this work.
I would like to express the deepest gratitude to
Prof. Amany Mostafa Professor of materials science -
Department of biomaterials - National Research Centre
for her supervision, guidance and suggesting the point
of this search and for her continuous advice and
enhancement throughout this work.
I wish to express my sincerest appreciation to Prof.
Dr. Mohamed I. El-Gohary Prof of Biophysics Faculty
of Science-AL-Azhar University for his supervision ,
continuous advice and support, encouragement, and
reviewing throughout the manuscript which have
rendered the realization of this work to be possible.
I would like to express the deepest gratitude to Dr
Azza Mahmoud researcher of Pharmaceutical
Technology Dept., National Research Centre, for
assisting me in chosen the appropriate drug , teaching
ACKNOWLEDGMENT
me the drug incorporation methodology and how to
assess the drug release throughout this work.
Many thanks to all my colleagues in the
Department of Biomaterials and the head of
Biomaterials department for the facilities offered and
continuous encouragement in various ways.
I'm very grateful to the research team at
university of Rennes 1.
Last, but not least, I would like to thank all the
members of my family for their continuous love,
encouragement, and support that kept me motivated
during my studies.
This work was financially supported by National
Research Centre, Cairo, Egypt and Campus of France.
Listofpublications
IV
List of publications
1. M. Mabrouk , A.A. Mostafa, H. Oudadesse, A.A.
Mahmoud and M.I. El-Gohary , Effect of
ciprofloxacin incorporation in PVA and PVA
bioactive glass composites scaffolds, Ceramics
International 40 (2014) 4833–4845.
2. Mabrouk M, Mostafa AA, Oudadesse H, Mahmoud
AA, Gaafar AM, and El-Gohary MI. (2013)
Fabrication, Characterization and Drug Release of
Ciprofloxacin Loaded Porous Polyvinyl
Alcohol/Bioactive Glass Scaffold for Controlled
Drug Delivery. Bioceram. Dev. Appl. S1: 009. doi:
10.4172/2090-5025.S1-009.
3. Mabrouk M., Oudadesse H., Mostafa A.A., and El-
Gohary M. I. In vitro assays: comparative study of
nanobioactive glass system by sol-gel. J.
Bioceramics development and applications (In
press).
Poster presentation :
4. Preparation of Polyvinyl Alcohol/Bioactive Glass
(PVA/BG) Nanocomposite Scaffolds and in-vitro
Assays for Applications as Biomaterials in
Orthopedic and Maxillofacial Surgery. Poster
Presented during journey due doctoral at
university of Rennes 1, France.
Listofabbreviations
V
List of abbreviations
Serial symbol 1 3D Three dimensional2 46S6 Bioactive glass with system of (46 % SiO2,
24% CaO, 24 % Na2O, 6 % P2O5 wt %) 3 BG-COL-PS Bioglass –Collagen- Phosphatidyl Serine 4 Ciprofloxacin 1-cyclopropyl-6-fluro-1, 4-dihydro-4-
oxo-7- (1-pipera Zinyl)-3- quinoline carboxylic acid
5 CG Chitosan Gelatin 6 Cs Chitosan 7 DNA Deoxyribo Nucleic Acid 8 DSC/TG Differential Scanning
Calorimetric/Thermo Gravimetric 9 DTA/TG Differential Thermal Analysis/Thermo
Gravimetric 10 ECM Extra Cellular Matrix 11 EDX Electron Dispersive X-rays 12 ESB European Society for Biomaterials 13 FTIR Fourier Transform Infrared 14 GPa Giga Pascal 15 HA HydroxyApatite 16 HE Hematoxylin and Eosin 17 HREM High Resolution Electron Microscope
18 ICP-OES Inductively Coupled Plasma -Optical Emission Spectrometer
19 KBr Potassium Bromide
20 kN Kilo Newton
21 kV Kilo Volt
22 mA Mille Amper
23 MBG Mesoporous Bioactive Glass 24 MCT Micro Computed Tomography
VI
25 MB Melting Bioactive glass 26 MIP Mercury Intrusion Porosimetry
27 μm Micron28 MPa Mega Pascal 29 MTT 3-(4,5-dimethylthiazol-2-yl)-2,5-
diphenyltetrazolium bromide, a yellow tetrazole
30 nBGC nanoparticles Bioactive Glass Ceramic 31 nm nanometer 32 OOKP Osteo Odonto- Kerato Prosthesis33 P % Porosity Percentage
34 PBS Phosphate Buffered Saline
35 PCL Poly Capro Lactone 36 PEG Poly Ethylene Glycol
37 PEO Polyethylene Oxide
38 PHB Poly Hydroxy Butyrate
39 PHBV Poly Hydroxy Butyrat Hydroxy Valerate
40 PLCA Poly Lactic Coglycolic Acid
41 PLA Poly Lactic Acid
42 PGA Poly Glycolic Acid
43 PLLA Poly L- Lactic Acid
44 PMMA Poly Methyl Meth Acrylate 45 PP Polypropylene Poly(ethylene terephtalate) 46 PPM Particles Per Million
47 PTFE Poly Tetra Fluoro Ethylene
48 PVA Poly Vinyl Alcohol 49 PVA/BG Polyvinyl alcohol/ Bioactive glass50 PVP Poly Vinyl Pyrrolidone
Listofabbreviations
VII
51 RP Rapid Prototyping
52 SBF Simulated Body Fluid
53 SG-B Sol-Gel Bioactive Glass 54 SCID mice Severe Combined Immuno Deficiency
mice 55 SC/PL Solvent Casting/ Particulate Leaching
56 SEM Scanning Electron Microscope57 SFF Solid Freeform Fabrication
58 SLA Stereo Lithography Analysis
59 TE Tissue Engineering 60 TEM Transmission Electron Microscope 61 TEOS Tetra Eth Oxy Silane
62 Tc Temperature of crystallization
63 Tf Temperature of fusion
64 Tg Temperature of glass transition
65 UV Ultra Violet 66 XRD X-rays diffraction 67 XRF X-rays fluorescence
ListofFigures
VIII
List of Figures
Serial Fig. No. Figure caption Page 1 Fig. 2.1 Schematic diagram of the
different phases in tissue engineering, from scaffold
fabrication and cell isolation to in
vivo implantation.
26
2 Fig. 2.2 Three-dimensional reconstructionof a cross-section of a long bone
showing the cortical and cancellous regions.
29
3 Fig. 2.3 Schematic representation of solvent casting/particulate
leaching (SCPL) method. The SEM image illustrates the morphology of a porous
hydroxyapatite/PLGA scaffold obtained using this method.
37
4 Fig. 2.4 Schematic representation of emulsion/freeze drying technique.
The SEM image illustrates the morphology of PCL scaffolds obtained using this method.
41
5 Fig. 2.5 Tissue engineering of patient-specific implant (e.g. bone graft)
via SFF technique.
43
6 Fig. 2.6 The chemical structure of PVA. 48 7 Fig. 2.7 Sol–gel processing and potential
processing methods. 53
8 Fig. 3.1 Preparation method of melted bioactive glass.
60
9 Fig. 3.2 Sol-gel method for bioactive glass preparation.
62
10 Fig 3.3 Schematic diagram for PVA/BG-Cip preparation method.
68
ListofFigures
IX
11 Fig. 3.4 DSC/TG instrument and some of its results.
69
12 Fig. 3.5 XRF instrument. 70 13 Fig . 3.6 TEM device and some of its
results. 70
14 Fig . 3.7 Zetasizer device and some of its results.
71
15 Fig.3.8 TEM device and some of its results.
72
16 Fig. 3.9 MIP instrument. 72 17 Fig. 3.10 Universal testing machine. 73 18 Fig. 3.11 XRD device and example of its
results. 74
19 Fig. 3.12 FTIR instrument and example of its results.
75
20 Fig. 3.13 ICP-OES device and example of its results.
76
21 Fig. 4.1 The thermal behaviour of SG-B with reference to MB.
81
22 Fig. 4.2-a
XRD of SG-B at different treatment temperatures with
reference to MB.
83
23 Fig. 4.2-b
The Influence of the sintering temperature on SG-B with
reference to MB.
84
24 Fig. 4.3 FTIR of SG-B and MB before immersion in SBF.
86
25 Fig. 4.4.a
TEM of MB. 89
26 Fig. 4.4-b
TEM of SG-B. 89
27 Fig.4.5.a
XRD for MB before and after immersion in SBF for 2, 5and 7
days.
90
28 Fig. XRD for SG-B before and after 91
ListofFigures
X
4.5.b immersion in SBF for 2, 5and 7 days.
29 Fig.4.6.a
FTIR for MB before and after immersion in SBF for 2, 5and 7
days.
92
30 Fig. 4.6.b
FTIR for SG-B before and after immersion in SBF for 2, 5and 7
days.
93
31 Fig. 4.7 SEM micrographs; a, c, e and g for sample MB before and after
immersion in SBF for 2, 5, 7 days and b, d, f and h for sample SG-B
before and after immersion in SBF for 2, 5, 7 days.
94
32 Fig. 4.8.a
Ca ions concentration after 2, 5 and 7days of immersion in SBF.
97
33 Fig.4.8.b
P ions concentration after 2, 5 and 7days of immersion in SBF.
97
34 Fig. 4.8.c
SBF P ions concentrations after soaking of the prepared samples
for different periods.
98
35 Fig. 4.9 The MTT assay of MB and SG-B. 99 36 Fig.
4.10 DTA of samples (MB, PVA
biocomposite and PVP biocomposite).
102
37 Fig. 4.11-a
XRD patterns of PVA biocomposite with reference to
PVA and MB.
104
38 Fig. 4.11-b
XRD patterns of PVP biocomposite with reference to
PVP and MB.
105
39 Fig. 4.12.a
FTIR of PVA biocomposite with PVP and MB.
107
40 Fig. 4.12.b
FTIR of PVP biocomposite with PVP and MB
108
ListofFigures
XI
41 Fig. 4.13-a
DLS of PVA biocomposite and PVP biocomposite with reference
to MB
109
42 Fig. 4.13.b
Zeta potential of PVA biocomposite and PVP
biocomposite with reference to MB.
110
43 Fig.4.14-a
FTIR of MB before and after soaking in SBF.
111
44 Fig. 4.14-b
FTIR of PVA biocomposite before and after soaking in SBF.
111
45 Fig. 4.14.c
FTIR of PVP biocomposite before and after soaking in SBF.
112
46 Fig. 4.15
Fig. (4.15), SEM images for a) MB before soaking in SBF, b) PVA biocomposite before soaking in
SBF, c) PVP biocomposite before soaking in SBF, d) MB after 5
days of soaking in SBF, e) PVA biocomposite after 5 days of
soaking in SBF, f) PVP biocomposite after 5 days of
soaking in SBF, g) MB after 7 days of soaking in SBF, h) PVA
biocomposite after 7 days of soaking in SBF and i) PVP
biocomposite after 7 days of soaking in SBF.
114
47 Fig. 4.16.a
SBF Ca ions concentrations after soaking of the prepared samples
for different periods.
116
48 Fig. 4.16.b
SBF P ions concentrations after soaking of the prepared samples
for different periods.
117
49 Fig. SBF Si ions concentrations after 117
ListofFigures
XII
4.16.c soaking of the prepared samples for different periods.
50 Fig. 4.17
SEM images for a) PVA scaffold, b) 1PVA:2MB scaffolds, c) PVA
loaded with 20% of drug , d) 1PVA:2MB loaded with 20% of
drug e) 1PVA:2SG-B and f) 1PVA:2SG-B loaded with 20% of
drug scaffolds with magnifications of Χ15and Χ100.
122
51 Fig. 4.18
The compressive strength of the prepared scaffolds before and
after drug incorporation.
123
52 Fig. 4.19.a
XRD of PVA and PVA/MB scaffolds before immersion in
SBF.
124
53 Fig. 4.19.b
XRD of PVA and PVA/SG-B scaffolds before immersion in
SBF.
125
54 Fig. 4.20.a
FTIR of PVA and PVA/MB scaffolds before immersion in
SBF.
127
55 Fig. 4.20.b
FTIR of PVA and PVA/SG-B scaffolds before immersion in
SBF.
128
56 Fig. 4.21
XRD of the prepared scaffolds before and after soaking in SBF.
130
57 Fig. 4.22
FTIR of the prepared scaffolds before and after soaking in SBF
134
58 Fig. 4.23
SEM image of the prepared scaffolds after immersion in SBF
for 21 days.
135
59 Fig. 4.24
ICP-OES analysis of the bioactivity solution.
138
60 Fig. XRD of the prepared scaffolds 140
ListofFigures
XIII
4.25 before and after drug loading 61 Fig.
4.26 FTIR of the prepared scaffolds before and after drug loading.
143
62 Fig. 4.27 Reaction mechanism PVA and ciprofloxacin.
144
63 Fig. 4.28
SEM image and EDX of a) ciprofloxacin , b) PVA 20% Cip ,
c) 1PVA:2MB 20% Cip and d) 1PVA:2SG-B 20% Cip.
145
64 Fig. 4.29
Biodegradation rate of the prepared scaffolds before and
after drug loading.
147
65 Fig. 4.30
The cumulative ciprofloxacin release for a) PVA scaffolds
loaded with 5,10 and 20% Cip, b) 1PVA:2MB scaffolds loaded with 5,10 and 20% and c) 1PVA:2SG-B scaffolds loaded with 5,10 and
20%.
150
66 Fig. 4.31
SEM of the prepared scaffolds after soaking in PBS.
151
List of tables
XV
List of tables
Serial Table No. Title of table Page 1 Table 2. 1 Mechanical properties of human
cortical bone.
29
2 Table 3.1 The used materials. 53 3 Table 3.2 The different compositions of the
prepared composite scaffolds. 64
4 Table 4.1 The chemical analysis of MB and SG-B determined by XRF analysis.
87
5 Table 4.2 Porosity percentage and pore diameter of the samples measured by Hg porosimeter and liquid displacement techniques.
121
Contents
Contents
Page
Statement of original authorship ------------------- I
Acknowledgement ------------------ II
Publications and conferences -------------------- IV
List of abbreviations ------------------------------------- V
List of figures ---------------------------------------- VIII
List of tables ------------------------------------ XIV
Contents ---------------------------------------------------- XV
Summary --------------------------------------------------- XX
Introduction and Aim of the Work ----------------- 1
Chapter (1)
Literature Review
Literature Review -------------------------- 8
Chapter (2)
Theoretical Aspects
2.1 Biomaterials Background-------------------------- 19
2.2 Tissue engineering ----------------------------------- 23
2.3 Bone and Bone Tissue Engineering------------ 27
2.3.1 Bone Structure --------- 27
2.3.2 Bone Tissue Engineering------------------------- 31
2.4Scaffolds and its role in tissue engineering ------- 32
2.4.1Biocompatibility of scaffolds ---------- 33
Contents
2.4.2. Biodegradability of scaffolds --------- 34
2.4.3. Preparation methods ------------- 34
2.4.3.1. Conventional scaffold fabricating techniques -----
35
2.4.3.2. Advanced scaffold fabricating techniques- - 39
2.4.4 Scaffolds as drug delivery system------- 41
2.5 Biomaterials for tissue engineering
applications---------- 43
2.5.1Polymers in orthopedic and maxillofacial
surgeries ---------- 46
2.5.2 Inorganic materials in orthopedic and
maxillofacial surgery ---------
49
Chapter (3)
Experimental Techniques
3.1. Materials--------------------- 56
3.2. Methods ---------- 61
3.2.1Preparation of Bioactive glass------ 61
3.2.1. a Melting molding technique ------------- 61
3.2.2.b Sol-gel method-------------- 62
3.3. Polymer route technique ------- 65
3.4. Scaffolds preparation ------- 66
3. 5. Preparation of Simulated Body Fluid---- 69
3.6. Characterizations techniques------ 70
3.6.1. Differential thermal analysis by (DSC) ----- 70
3.6.2 Elemental composition analysis (XRF) ------ 71
3.6.3. Transmission electron microscope (TEM) --- 72
3.6.4. Particle size distribution and charge using
Zetasizer-------- 73
Contents
3.6.4. Morphological and microstructural
properties---------- 74
3.6.5 Mechanical properties of the prepared
scaffolds------------ 74
3.6.6. Bioactivity Assessment ------- 74
3.6.7. Drug loaded scaffolds In-vitro degradation
studies-------- 77
3.6.8. Ciprofloxacin release behavior-------- 77
3.6.9. Mechanism of ciprofloxacin release----- 78
Chapter (4)
Results & Discussion
4.1. Characterization of 46S6 bioactive glass
prepared by melting and sol-gel methods ----------- 81
a) DSC/TG analysis---------------- 81
b.1) XRD analysis before immersion in SBF:----- 83
b.2) Influence of the sintering temperature on
the prepared powder by sol gel method----- 84
c) FTIR before immersion of MB and SG-B in
SBF solution----- 86
d) X-rays Fluorescence (XRF) analysis-------- 87
e) Morphology and particle size of bioactive
glass using TEM - 88
f) Bioactivity Assesment ------- 90
f.1) XRD after immersion of MB and SG-B
in SBF at different periods ------- 90
f.2) FTIR of MB and SG-B before and after
immersion in SBF for different time intervals-- 92
f.3) SEM evaluation before and after immersion in
SBF at different periods------------ 94
f.4) Chemical reactivity investigation using ICP-
OES------------ 96
g) Cytotoxicity and cellular viability ------------ 99
Contents
4.2 Characterization of polymer technique for
Composites Preparation- 102
a) DSC/TG analysis----- 102
b) XRD before immersion in SBF------- 104
c) FTIR before immersion in SBF --------- 106
d) Dynamic light scattering (DLS) and zeta
potential ---------- 109
e) Bioactivity Assessment ----------- 111
e.1) FTIR before and after immersion in SBF ----- 111
e.2) SEM before and after immersion in SBF------- 113
e.3) Ions concentrations in SBF by ICP-OES ----- 116
4.3 Scaffolds Results---------- 120
4.3. BG/PVA scaffold with and without drug----- 121
4.3.1. Morphological and microstructural
properties--------- 122
4.3.2.Mechanical properties -------- 123
4.3.3 XRD before immersion in SBF--------- 124
4.3.4. FTIR before immersion in SBF------------ 126
4.3.5. Bioactivity Assessment------------------------ 129
a) XRD after immersion in SBF------------ 129
b) FTIR after immersion in SBF--------- 132
c) SEM with EDX after immersion in SBF------ 135
d) Evaluation of elemental concentrations in SBF-- 137
4.3.6. Ciprofloxacin incorporation---------- 138
a) XRD analysis before and after drug loading-- 139
b) FTIR spectra before and after drug loading - 142
Contents
C) SEM coupled with EDX-------- 145
4.3.7. Scaffolds Degradation ------------- 147
4.3.8. Release behavior of ciprofloxacin-------- 149
Conclusion----------------------------------------- 153
References --------------------------------------------- 156
Résumé:
Les verres bioactifs élaborés par fusion et par sol-gel présentent un grand
intérêt lorsqu�ils sont utilisés en tant qu�implants osseux. Les travaux effectués
dans notre groupe de recherche « Biomatériaux» ont montré leur bonne
biocompatiblité. Le dopage des verres 46S6 (46% SiO2-24% CaO-24% Na2O-
6% P2O5) avec des éléments tels que le magnésium, le strontium ou le zinc ont
permis de faire varier leur cinétique de réactivité chimique et de bioactivité.
Ainsi, celles-ci peuvent être adaptées aux patients et à leur métabolisme osseux
qui varie avec l�âge entre autres. De même, l�association des verres bioactifs
avec un bio polymère tel que le chitosan a montré que ces composites (verres-
chitosan) peuvent aussi servir à délivrer ces biomolécules dans le squelette et
traiter certaines pathologies osseuses.
Ce travail de thèse est basé sur la préparation de verres bioactifs (BG) par
différents procédés tel que la fusion, la voie sol-gel et le scaffolds. La synthèse
de verres bioactifs par le procédé scaffolds est une nouvelle méthode de
synthèse dans notre groupe de recherche. Par ce nouveau procédé, les
biomolécules introduites pourront être véhiculées vers les cellules et dans le
squelette de manière relativement contrôlée.
L�avantages des verres scaffolds réside dans leur micro architecture et dans la
maîtrise de la porosité induite dans ces biomatériaux. La matrice de base
constituant le verre bioactif utilisé dans ce travail est le 46S6 formé de 46 %
SiO2- 24% CaO- 24% Na2O � 6% P2O5. Le choix de cette composition chimique
est basé sur les compositions déjà étudiées dans le groupe Biomatériaux en site
osseux, pour pouvoir faire des comparaisons rigoureuses et interpréter les
phénomènes induits suite aux modifications des paramètres de synthèse de
dopants ou de polymères de manière objective.
Le Poly Vinyl Alcohol (PVA) a été associé aux verres élaborés dans un
système quaternaire (BG) par les procédés cités (fusion, sol-gel et sacffolds).
Différents paramètres intervenant dans les synthèses des verres bioactifs ont été
étudiés, nous citons à titre d�exemple : la température, le pH, la taille des
particules, le rapport Polymère / verres, la microstructure, la porosité et la
biodégradation. Les caractéristiques thermiques des verres élaborés ont été
également déterminées après chaque synthèse par analyse thermique
différentielle (DSC). Ainsi, la température de fusion, la température de transition
vitreuse ainsi que la température de cristallisation ont été élucidées. Ces
caractéristiques thermiques changent lorsque la composition chimique du verre
est modifiée. A ce titre, les compositions chimiques ont été étudiées par
Fluorescnece (XRF) et Inductively Coupled Plasma-Opticale Emission
Spectroscopy (ICP-OES) après chaque synthèse pour s�assurer de la pureté des
verres bioactifs élaborés et destinés à des applications médicales. Plusieurs
techniques physico chimiques d�analyses (DRX, MEB, MET, FT-IR, XRF, ICP-
OES) ont été mises en �uvre pour déterminer les propriétés physico chimiques
de nos verres bioactifs avant et après expérimentations « in vitro ». Le nano
composite Polymère - Verres scaffolds que nous avons obtenu présente des
particules de tailles comprises entre 40 et 61 nm et une porosité d�environ 85%.
La biodégradation des verres scaffolds décroît lorsque la teneur en verre
scaffolds dans le nano composite croît. Les expérimentations « in vitro »
montrent qu�après immersion de ces nano composites dans un liquide
physiologique synthétique (SBF), une couche d�apatite (phosphate de calcium)
se forme à leur surface. L�épaisseur de la couche formée dépend clairement de la
taille des particules et du rapport polymère / verre scaffolds.
Les résultats obtenus après synthèse par les différents procédés montrent
bien des matériaux amorphes élucidés par DRX avec présence des liaisons Si-O-
Si, P-O. L�analyse thermique et les diagrammes des rayons X ont montré que
pour le procédé sol-gel, la température appropriée pour l�obtention du matériau
amorphe 46S6 est de 600°C.
Pour les verres élaborés par scaffolds, la porosité de 85% décroit
légèrement lorsque le pourcentage de verre dans le nano composite augmente.
Un réseau micro structuré de pores a été mis en évidence par microscopie
électronique à balayage et à transmission (MEB et MET), il varie entre 145 µm
et 6,3 nm.
Les différents verres élaborés on été mis en contact avec un liquide
physiologique synthétique, le Simulated Body Fluid (SBF), de composition
chimique similaire à celle du plasma sanguin. Les délais d�immersion sont
compris entre quelques heures et environ 30 jours. Après retrait des verres
bioactifs du liquide SBF, des évaluations physico chimiques et biologiques ont
été réalisées pour les différents délais d�immersion. Le relargage du Si vers la
solution SBF et l�utilisation du Ca et du P nécessaires à la formation d�une
couche d�apatite biologique ont été évalués pour chaque type de verre et pour
chaque délai d�immersion.
La formations d�une couche de phosphate de calcium sous forme
d�hydroxyapatite (Ca10(PO4)6(OH)2 et de phosphate tricalcique-β a été élucidée
par les techniques citées ci-dessus. La qualité de cristallisation et l�épaisseur des
couches formées dépendent largement du mode de synthèse. Il en de même pour
la cinétique de bioactivité. Tous les verres bioactifs ont montré un
comportement cellulaire avec une prolifération et une adhésion comparables à
celles du témoin utilisé lors des tests biologiques.
Une autre molécule a été associée au verre bioactif et a été étudiée dans ce
travail, il s�agit de la ciprofloxacine. Le biocomposite scaffolds composé du
Poly Vinyl Alcohol et du verre bioactif chargé avec de la ciprofloxacine
présente une porosité inter connectée et bien structurée. Son relargage du verre
élaboré par scaffolds vers le SBF a été relié à la nature de la porosité du verre
support et au caractère hydrophilique de cette molécule.
Summary
i
Summary
Scaffolds are implants used to deliver cells, drugs, and genes
into the body in a local controlled release pattern which offers many
advantages over systematic drug delivery. The ideal scaffolds should
have appropriate microstructures to facilitate cellular attachment,
proliferation and differentiation. In addition, the scaffolds should
possess adequate mechanical strength and biodegradation rate
without any undesirable by-products.
The aim of the present work is the preparation of Bioactive
Glass (BG) 46S6 by different techniques. Fabrication of composite
scaffolds by using of Poly Vinyl Alcohol (PVA) and quaternary BG (two
methods melting and sol-gel) with different ratios to the prepared
scaffolds was carried out. This drug has antibacterial and osteogentic
effects. Different factor affecting the final properties of the prepared
composite scaffolds were investigated in this study such as;
temperature of treatment, BG particle size, polymer/glass ratio,
microstructure, porosity, biodegradation, bioactivity, and drug release.
The thermal behavior of the prepared bioactive glass by sol-gel
and melting techniques were identified using Differential Scanning
Calorimetric/Thermo Gravimetric (DSC/TG) or Differential Thermal
Analysis/Thermo Gravimetric (DTA /TG). Moreover, the glass transition
temperature Tg, glass crystallization temperature Tc and the glass
fusion temperature Tf were also determined.
The elemental composition of the prepared bioactive glasses
was determined by X-rays Fluorescence (XRF) to confirm that the
prepared bioactive glasses have the same elemental compositions.
The prepared bioactive glass by sol-gel method has higher purity than
those of bioactive glass prepared by melting technique. The particle
Summary
i
size of the prepared bioactive glass was determined by Transmission
Electron Microscopic (TEM). Nano-bioactive glass could be obtained by
modified sol-gel and the obtained particle size ranged between 40 to
61 nm. However, the prepared BG samples by sol-gel in nanoscale
shows bioactivity and biocompatibility more than those prepared by
melting technique as confirmed by bioactivity test in SBF and MTT
tests.
The investigation of the transformed phases was conducted by
X-rays Diffractometer (XRD) and Fourier Transmission Infra Red
spectroscopic (FTIR) techniques. The prepared bioactive glass by both
applied methods has the same amorphous phase and all identified
groups as well as. Besides it has been confirmed by XRD that the
appropriate temperature for preparing of 46S6 system of bioactive
glass by sol-gel technique was 600ºC.
The porous scaffold has 85% porosity with a slight decrease by
increasing the glass contents. The microstructured pore network was
formed between 145 µm to 6.3 nm relatively in uniform size pores and
with thin pore wall. The thickness of the pore walls increased by
increasing the glass contents in the prepared scaffolds. The BG
particles were embedded in the polymer matrix as it was confirmed by
Scanning Electron Microscopic (SEM) and Energy Dispersive X-rays
(EDX). The degradation rate decreased by increasing of glass content
in the prepared scaffolds.
The bioactivity of the prepared composite scaffolds was
evaluated by XRD, FTIR, SEM coupled with EDX and Inductively
Coupled Plasma -Optical Emission Spectroscopic (ICP-OES). It has
been observed that after soaking in Simulated Body Fluid (SBF), there
was an apatite layer formed on the surface of the prepared samples
Summary
i
with different thickness depending on the glass particle size and
polymer/glass ratio. Also, the concentration of Ca and P ions of SBF
solution decreased due to their consuming in formation of apatite
layer on the surface of the samples. Bioactive glass dissolution was
confirmed by increasing of Si ions concentration in the SBF.
The effects of the glass particle size, percentage and drug
concentrations on the mechanical strength of the prepared scaffolds
were determined by universal testing machine. It could be noted that
the mechanical properties were more enhanced by incorporation of
nano BG than micro BG. Also, the increase of the drug concentrations
enhances the compressive strength.
Assessment of the drug loaded scaffolds was evaluated in PBS by
UV-Spectrophotometer. Release rate of ciprofloxacin was enhanced as
the glass polymer ratio was increased.
The PVA/BG biocomposite scaffolds loaded with ciprofloxacin
with well interconnected pore structure and appropriate porosity were
fabricated via freeze drying technique as confirmed by the results of
SEM, Mercury Intrusion Porosimeter (MIP) and liquid displacement
method this was assured for orthopedic and maxillofacial surgeries.
Introduction & Aim of the work
1
Introduction and Aim of the work
1. Introduction
There is a growing demand for replacing bone substance
that has been lost due to traumatic or non traumatic events
Suchanek and Yochimara (1998). In order to achieve a
satisfactory result and have an appropriate host response at the site
of implantation, suitable implanted biomaterials should have
certain desired properties. The microstructred features and the
mechanical properties of the bone must be thoroughly understood
for the preparation of successful candidate in order to mimic the
natural bone structure.
Reconstructive treatment of bone defect in orthopedic and
maxillofacial surgery is a wide spread practice. Osteoclasts and
osteoblasts cells ensure a balanced control of bone resorption and
formation, resulting in bone repair, renewal and growth. The self
healing capacity of bone is widely used for the repair of small
fractures. However, bone grafts are needed to provide support, fill
lacunae and enhance biological repair when the skeletal defect
reaches a critical size. Orthopedic and maxillofacial surgeons
employ bone grafts or substitutes for non-union defects and the
replacement of diseased tissue after trauma, infection and tumor
resection or prosthetic revision. The worldwide market for bone
replacement and repair is estimated at ~ €300 million, including
Introduction & Aim of the work
2
autologous, allogenic, xenogenic and synthetic bone materials
Laurencin, et al (1999).
Several kinds of material can be used in the following
procedures: (a) Autografts (from patient itself), (b) allograft
(homografts from human or xenograft from animals) (c) implants
from synthetic bone bonding biomaterials.
Biomaterial is defined as any substance or combination of
substances that can be used for any period, as a whole or as a part
of a system for use in the human body to measure, restore and
improve physiological function and quality of life. Metals,
ceramics, composites and polymer (natural or synthetic) have been
used as an artificial heart valves, synthetic blood vessels, artificial
limps, dental composites and polymers for controlled drug release.
Biomaterials should be compatible with body in order to exhibit
their function probably. Incompatible materials may induce
unfavorable immune reactions, undesirable interactions with the
blood and the body fluids.
Bone implants have their advantages and disadvantages.
Polymer has low mechanical strength compared to bone, however,
metals have superior mechanical properties but they are corrosive.
Ceramics are brittle with low fracture toughness despite their other
desired properties such as wear resistance. Biocompatible
composite materials are considered the reasonable approach to
achieve reasonable combined properties Oudadesse, et al (2011).
Introduction & Aim of the work
3
Bone is a composite material composed of an organic
matrix; made essentially of collagen type (I) mineralized with
hydroxyapatite. The composite nature of bone has a complex
microstructure difficult to imitate which gives most of the superior
mechanical properties of bone. Extensive research has been
conducted on bone substitute composite materials composed of
bioactive materials and polymer Chu (2007).
One of the key issues in designing clinically transplantable
regenerative tissue is the generation of a functional microvascular
network within the engineered constructs to provide oxygen and
nutrients to facilitate growth, differentiation, and tissue
functionality Kneser et al (2006) and Brey et al(2002). An
inadequate microvascular network will result in the hypoxic cell
death of engineered tissues Stahl, et al (2004) leading to total
implant failure Cassell, et al (2002).
The specific criteria for ideal scaffolds used in bone tissue
engineering are the following; ability to deliver cells, excellent
steoconductivity, good biodegradability, appropriate mechanical
properties, highly porous structure, irregular shape fabrication
ability, in addition to potential commercialisation.
Development of composite scaffold materials has an
advantageous property of two or more types of materials can be
used to suit better the mechanical and physiological demands of
the host tissue. By taking advantage of the formability of polymers
Introduction & Aim of the work
4
and including controlled-volume fractions of a bioactive ceramic
phase, mechanical reinforcement of the fabricated scaffold can be
achieved Boccaccini and Maquet (2003) and Ramakrishna, et
al (2001). At the same time, the poor bioactivity of most polymers
can be counteracted.
The most important driving force behind the development
of polymer/bioactive glass composite scaffolds for orthopedic and
maxillofacial surgeries and for most of bone tissue engineering is
the need for conferring bioactive behavior to the polymer matrix,
which is achieved by the bioactive inclusions or coatings. The
degree of bioactivity is adjustable by the volume fraction, size,
shape and arrangement of inclusions Maquet, et al (2004) and
Wang, et al (2003). It has been shown that increased volume
fraction and higher surface area to volume ratio of inclusions favor
higher bioactivity; hence in some applications the incorporation of
fibers instead of particles is favored Jiang, et al (2005) and
Jaakkola, et al (2004). Addition of bioactive phases to
bioresorbable polymers can also alter the polymer degradation
behavior, by allowing rapid exchange of protons in water for
alkali in the glass or ceramic. This mechanism is suggested to
provide a pH buffering effect at the polymer surface, modifying
the acidic polymer degradation Li, et al (2005).
BG has various applications in repair and reconstruction of
bone tissue; however, it has week mechanical properties especially
Introduction & Aim of the work
5
in porous form. One approach to enhance the mechanical
properties of materials is the elaboration of BG with polymer to
form composites Yazdanpanah, et al (2012). This way leads to
an excellent combination between strength and toughness, as well
as improved characteristics, when compared to their individual
components Sokolsky-Papkov, et al (2007) .The composites of
BG/polymer are able to provide construct with excellent
osteogensis and angiogenesis Yazdanpanah, et al (2012).
Local application of antibiotic release systems is important
for hard tissue engineering because of both poor vascularity in
bone tissue for oral or intravascular therapy and easiness of
microbial attack in dental sites where it is open area to
environment Garima and Bikramjit (2012) and Swapnika, et al
(2012). The composite scaffolds could be used as drug-delivery
systems for antibiotic treatment of osteomylitis, a common bone
disease caused by bacterial infection of bone modullare cavity,
cortex, and /or periosteum upon implantation. These systems have
the advantage that no second surgical procedure is required for
implant removal. Ciprofloxacin is a fluroquinolone derivative,
widely used in osteomyelitis because of its favorable penetration
and bactericidal effect on all the probable osteomyelitis pathogens.
Ciprofloxacin act by inhibiting the bacterial enzymes DNA gyrase
Robert, et al (2012).
Introduction & Aim of the work
6
Aim of the present work
The objectives of this study is the preparation and
characterization of polymeric /bioactive porous composite
biomaterials by using PVA and BG 46S6
First, BG 46S6 will be prepared by different methods;
melting and sol-gel. The characterization of the prepared powders
after heating at 1300 ºC for melting technique, and at 600 ºC for
sol-gel will verified. Thus, the chemical composition, grain size
and bioactivity in vitro will be investigated. The prepared
bioactive glass by both methods will be used in the fabrication of
biocomposite scaffold using Freeze drying technique. The
microstructure of the scaffolds will be examined by SEM and Hg
prosometer. The bioactivity of the prepared scaffolds will be
tested in vitro by immersion in SBF solution for different intervals
up to 30 days. Loading the prepared scaffolds with water-soluble
drug (ciprofloxacin) will be achieved during the fabrication of the
scaffolds without the use of possibly toxic surfactants. The effect
of BG content and the drug percentage on the development of
tailored medicated scaffolds during freeze drying will be
investigated. The chemical, morphological, mechanical,
biodegradation rate and release properties of ciprofloxacin loaded
scaffolds will also declared.
Chapter 1 Literature Review
7
Chapter 1
Literature Review
Bioactive glass (BG) has various applications in repair and
reconstruction of bone tissue; however, it has week mechanical
properties especially in porous form. One approach to enhance the
mechanical properties of materials is the elaboration of BG with
polymer to form composites. This way leads to an excellent
combination between strength and toughness, as well as improved
characteristics, when compared to their individual components
.The composites of BG/polymer are able to provide construct with
excellent osteogensis and angiogenesis. Application of a drug to a
specific region using drug loaded scaffold produce high
concentration of the drug in the required site of action which
eliminate the side effects that prohibit the administration of large
oral dose. Ciprofloxacin (Cip) is a fluroquinolone derivative,
widely used as an antibiotic and in osteomyelitis because of its
favorable penetration and bactericidal effect on all the probable
osteomyelitis pathogens. The main purpose of the current study
was to develop and fabricate a construct of bioactive scaffold
combining an antibiotic (ciprofloxacin) as a target drug delivery
system.
Dietrich et al., (2008) fabricated bioactive glasses in the
system SiO2–CaO–Na2O–P2O5 pure and doped with magnesium or
zinc by melt-derived method. The bioactivity was studied during
Chapter 1 Literature Review
8
in vitro assays: the ability of hydroxyl Carbonate Apatite (HCA)
layer to form on the glass surface was examined after contact with
simulated body fluid (SBF). The X-ray diffraction (XRD), Fourier
Transform Infrared (FTIR) and scanning electron microscopy
(SEM) studies were performed before and after immersion in vitro
assays. The SBF solutions were also analyzed using inductively
coupled plasma-optical emission spectroscopy (ICP-OES).
Introduction of magnesium and zinc as trace element induces
several modifications on the observed phenomena at the glass
surface and in SBF solution after immersion of the samples. The
chemical durability of the glasses, the formation of the silica-rich
layer and the crystallization of the HCA layer were affected, but
not present the same modifications as the introduced doping
element.
Oudadesse et al., (2010) synthesized a pure bioactive glass
(46S6) and zinc-doped bioactive glass (46S6Zn10) with 0.1 wt%
zinc by melting and rapid quenching. Cylinders of both types of
glasses were soaked in a simulated body fluid (SBF) solution to
determine the effect of zinc addition as a trace element on the
chemical reactivity and bioactivity of glass. Several physico-
chemical characterization methods such as x-ray diffraction,
Fourier transform infrared spectroscopy and nuclear magnetic
resonance methods, with particular focus on the latter, were
chosen to investigate the fine structural behaviour of pure and Zn-
doped bioactive glasses as a function of the soaking time of
Chapter 1 Literature Review
9
immersion in SBF. Inductively coupled plasma-optical emission
spectroscopy (ICP-OES) was used to measure the concentrations
of Ca and P ions in the SBF solution after different durations of
immersion. The effect of the investigated samples on the
proliferation rate of human osteoblast cells was assessed by the 3-
(4,5-dimethyl-2-thiazolyl)-2,5-diphenyl-2H-tetrazolium bromide
(MTT) assay, and tested on two different sizes of pure and zinc-
doped glasses in powder form, with particle sizes that ranged
between 40 to 63 μm and 500 to 600 μm. The obtained results
showed the delay release of ions by Zn-doped glass (46S6Zn10)
and the slower CaP deposition. Cytotoxicity and cell viability
were affected by the particle size of the glass. The release rate of
ions was found to influence the cell viability.
Peter, et al., (2010) prepared a novel nanocomposite
scaffold of chitosan (CS) and bioactive glass ceramic
nanoparticles (nBGC) by blending nBGC with chitosan solution
followed by lyophilization technique. The particle size of the
prepared nBGC was found to be 100 nm. The composite scaffolds
showed adequate swelling and degradation properties. The in-vitro
biomineralization studies confirmed the bioactivity nature of the
composite scaffolds. Cytocompatability of the composite scaffolds
were assessed by MTT assay, direct contact test and cell
attachment studies. Results indicated no toxicity, and cells
attached and spread on the pore walls offered by the scaffolds.
These results indicate that composite scaffolds developed using
Chapter 1 Literature Review
10
nBGC disseminated chitosan matrix as potential scaffolds for
tissue engineering applications.
Peter, et al., (2010) prepared a novel nanocomposite
scaffold of chitosan (CS)–gelatin (CG) with nBGC was prepared
by blending of chitosan and gelatin with nBGC. The prepared
CG/nBGC nano-composite scaffolds showed macroporous
internal morphology in the scaffold with pore size ranging from
150 to 300µm. Degradation and swelling behavior of the
nanocomposite scaffolds were decreased, while protein adsorption
was increased with the addition of nBGC. Biomineralization
studies showed higher amount of mineral deposits on the nano-
composite scaffold, which increases with increasing time of
incubation. MTT assay, direct contact test, and cell attachment
studies indicated that, the nano-composite scaffolds are better in
scaffold properties and it provides a healthier environment for cell
attachment and spreading. So, the developed nano-composite
scaffolds are a potential candidate for alveolar bone regeneration
applications.
Boccaccini, et al., (2010) presents the state of art of the
preparation of nanoscale bioactive glasses and corresponding
composites with biocompatible polymers. The recent
developments in the preparation methods of nano-sized bioactive
glasses are reviewed, covering sol–gel routes, microemulsion
techniques, gas phase synthesis method (flame spray synthesis),
laser spinning, and electro-spinning. Then, examples of the
Chapter 1 Literature Review
11
preparation and properties of nanocomposites based on such
inorganic bionanomaterials are presented, obtained using various
polymer matrices, including polyesters such as
poly(hydroxybutyrate), poly(lactic acid) and poly(caprolactone),
and natural-based polymers such as polysaccharides (starch,
chitin, chitosan) or proteins (silk fibroin, collagen). The physico-
chemical, mechanical, and biological advantages of incorporating
nanoscale bioactive glasses in such biodegradable nanocomposites
are discussed and the possibilities to expand the use of these
materials in other nanotechnology concepts aimed to be used in
different biomedical applications are also highlighted.
Chengtie, et al., (2011) found that Mesoporous bioactive
glass (MBG) /silk scaffolds have better physiochemical properties
(mechanical strength, in vitro apatite mineralization, Si ion release
and pH stability) compared to non-mesoporous bioactive glass
(BG) /silk scaffolds. MBG and BG both improved the in vivo
osteogenesis of silk scaffolds. microcomputed tomography (lCT),
hematoxylin and eosin (HE) analyses showed that MBG/silk
scaffolds induced a slightly higher rate of new bone formation in
the defects than did BG/silk scaffolds and immunohistochemical
analysis showed greater synthesis of type I collagen in MBG/silk
scaffolds compared to BG/silk scaffolds.
Caixia, et al., (2011) revealed that the in vivo results
showed that Bioglass-Collagen-Phosphatidylserine (BG-COL-PS)
composite scaffolds exhibited good biocompatibility and extensive
Chapter 1 Literature Review
12
osteoconductivity with host bone. Moreover, the BG-COL-PS/MS
cells constructs dramatically enhanced the efficiency of new bone
formation than pure BG-COL-PS scaffolds or BG-COL/MSC
constructs. All these results demonstrate the usefulness of PS
composited BG-COL-PS scaffolds for inducing enhanced bone
formation. The BG-COL-PS scaffolds fulfill the basic
requirements of bone tissue engineering scaffold and have the
potential to be applied in orthopedic and reconstructive surgery.
Poursamar , et al., (2011) prepared porous scaffolds with
three dimensional microstructures, and in vitro experiments with
osteoblast cells indicated an appropriate penetration of the cells
into the scaffold’s pores, and also the continuous increase in cell
aggregation on the scaffolds with increase in the incubation time
demonstrated the ability of the scaffolds to support cell growth.
According to the obtained results, the nanocomposite scaffolds
could be considered as highly bioactive and potential bone tissue
engineering implants.
Kaisa , et al., (2011) prepare a synthetic keratoprosthesis
skirt for use in osteoodonto- keratoprosthesis (OOKP) surgery,
bioactive glass and polymethyl methacrylate (PMMA)-based
composites, by using of different bioactive glasses (45S5, S53P4
and 1-98) with two different forms ( particles and porous glass
structures). The results indicated that the bioactive composites
could be stable synthetic candidates for a keratoprosthesis skirt in
the treatment of severely damaged or diseased cornea.
Chapter 1 Literature Review
13
Xin, et al., (2011) fabricate a bioactive glass (13-93)
scaffolds with promising microstructure and mechanical response
for potential use in the repair of load-bearing bones using a
method based on unidirectional freezing of camphene-based
suspensions. Annealing the frozen constructs for 0–72 h at --34 °C
(slightly below the solidification temperature of the suspension)
resulted in coarsening of the camphene crystals, which provided a
method by which to control the pore diameter of the constructs (in
the range 15–160 μm after sublimation of the camphene).
Coarsening of the camphene crystals during the annealing step can
be described by a diffusion-controlled coalescence model.
Sintering resulted in a decrease in the porosity and the pore
diameter, giving scaffolds with porosities of 20–60% and pore
diameters of 6–120 μm for annealing times of 0–72 h. The
sintered scaffolds had compressive strength and elastic modulus
values in the freezing (orientation) direction which varied from
180 MPa and 25 GPa (porosity = 20%), respectively, to 16 MPa
and 4 GPa (porosity = 60%) which were 2–3 times larger than
those measured in the direction perpendicular to the orientation
direction.
El-Kady, et al., (2012) synthesized glass nanoparticles
containing 1, 3, 5, and 10 wt% of Ag2O (coded; GAg1%, GAg3%,
GAg5%, and GAg10%, respectively) through a quick alkali
mediated sol–gel method. All samples had an antibacterial effect
against different types of bacteria and the extraction of silver ions
Chapter 1 Literature Review
14
from them followed a diffusion-controlled mechanism, which
could demonstrate their ability to treat bone infection.
El-Kady and Ali Ashraf (2012) synthesized bioactive glass
nanoparticles in the system (SiO2–CaO–P2O5–ZnO) following the
sol–gel technique. The prepared glass nanoparticles of 1, 3 and 5
wt% of ZnO (coded: GZ1, GZ3 and GZ5, respectively). All glass
powders were highly porous (75, 76 and 75%) with surface areas
of 233, 94 and 118 m2/g for GZ1, GZ3 and GZ5, respectively. All
glass powders induced an appetite layer on their surfaces upon
immersion in simulated body fluid (SBF).
Zhengmao, et al., (2012) prepared a three-dimensional
lamellar structured bioactive glass powders using nonionic block
copolymer surfactants as structure-directing agents through a sol–
gel method. The biomineralized products on the surfaces of the
bioactive glass powders were apatite microcrystals with a low
crystallinity, the composition and morphologies of the apatite
microcrystals changed with the immersion time increased.
Garima , et al., (2012) used the polymer sponge replication
method to prepare the macroporous hydroxyapatite scaffolds with
interconnected oval shaped pores of 100–300 µm with pore wall
thickness of about 50 µm. The compression strength of 60 wt. %
HA loaded scaffold was 1.3 MPa. The biological response of the
scaffold was investigated using human osteoblast like SaOS2
cells. The results showed that SaOS2 cells were able to adhere,
proliferate and migrate into pores of scaffold. Furthermore, the
Chapter 1 Literature Review
15
cell viability was found to increase on porous scaffold compared
to dense HA.
Swapnika, et al., (2012) evaluated and modeled the
viscoelastic characteristics of chitosan and chitosan–gelatin
scaffolds prepared using a freeze-drying technique. Chitosan and
chitosan–gelatin solutions (0.5 and 2 wt. %) were frozen at -80 ºC
and freeze-dried. Using the scaffolds, uniaxial tensile properties
were evaluated under physiological conditions. The models were
used to fit the experimental stress-relaxation data and the
parameters obtained from modeling were used to predict their
respective cyclic behaviors, which were compared with cyclical
experimental results. These results showed that the model could be
used to predict the cyclical behavior under the tested strain rates.
The model predictions were also tested using cyclic properties at a
lower strain rate of 0.0867% s-1 (5% min-1) for 0.5 wt. % scaffolds
but the model could not predict cyclical behavior at a very slow
rate. They summarizing that the pseudo-component modeling
approach can be used to model the sequential strain-and-hold stage
and predict cyclical properties for the same strain rate.
Robert, et al., (2012) studied the effect of hydroxyapatite
reinforcement on the architecture and mechanical properties of
freeze-dried collagen scaffolds. They had found that HA whisker
reinforced scaffolds exhibited a nearly four-fold greater modulus
compared to the equiaxed HA powder, while there were no
differences with the HA reinforcement morphology at high and
Chapter 1 Literature Review
16
low reinforcement levels. Therefore, the elongated morphology of
HA whiskers enabled a reinforcing effect at a lower level of
reinforcement compared to a conventional, equiaxed HA powder.
Puga, et al., (2012) suggested that prevention and treatment
of osteomyelitis could be achieved through local drug delivery
using implantable devices, which provide therapeutic levels at the
infection site with minimum side-effects. Physical blends of
polycaprolactone (PCL) and poloxamine (Tetronic_) were
prepared by applying a solvent-free hot melting approach to obtain
cytocompatible implants with a tunable bioerosion rate,
ciprofloxacin release profile and osteoconductive features.
Incorporation of the block copolymer at weight ratios ranging
from 25 to 75 wt. % led to matrices with viscoelastic parameters
in the range of those of fresh cortical bone. Once immersed in
buffer the matrices underwent a similar weight loss in the first
week to the content of poloxamine, followed by a slower erosion
rate due to PCL. The initial rapid erosion and the increase in
porosity partially explain the observed burst of ciprofloxacin
release. The matrices sustained ciprofloxacin release for several
months (<50% released after 3 months) and showed in vitro
efficacy against Staphylococcus aureus, eradicating the bacteria in
less than 48 h.
Chang, et al., (2013) fabricate a novel hybrid hydrogels by
introducing nano-hydroxyapatite into chitin solution. Their
structure and morphology were characterized by FTIR spectra,
Chapter 1 Literature Review
17
wide-angle X-rays diffraction (WAXD), TGA, SEM, and TEM.
Their results revealed that hydroxyapatite nano-particles were
uniformly dispersed in chitin hydrogel networks. The hybrid
hydrogel exhibited about 10 times higher mechanical properties
(compressive strength: 274 kPa) than that of chitin hydrogel.
Moreover, COS-7 cell culture experiment proved that cells could
adhere and proliferate well on the hybtid hydrogels, suggesting
good biocompatibility. All these results signified that these
biomaterials could be potential candidates as scaffolds for tissue
engineering.
Zhang et al., (2013) study the influence of porosity on
long-term degradation of PCL scaffold in phosphate buffer
solution (PBS). A 72-week degradation study of PCL scaffolds
with various porosities was conducted to elucidate the changes of
physico-chemical properties such as weight, molecular weight,
morphology and compressive modulus. Within 72 weeks, PCL
scaffolds experienced three stages: stable stage, mechanical loss
stage and structural collapse stage. The higher porosity induced
the severer loss of weight, molecular weight and compressive
modulus. It was found that a minimal acid autocatalysis also
happened in the scaffold samples with low porosities (less than
85%). Cellular response on the scaffolds with various porosities
was further evaluated. The cell ingrowth improved on the scaffold
with high porosity in contrast to those with low porosity. The
combined results demonstrated that an optimal porosity of PCL
Chapter 1 Literature Review
18
scaffolds should be designed greater than 90% due to the
appropriate degradation rate and good cell performance.
Wu et al.,
(2013) prepare a novel porous HA/b-TCP
bioceramics scaffold with micro-ribs structure extrusion
deposition technique and microwaves intering. Micro-ribs were
placed at the center and corners of scaffold along the direction of
load. Mechanical behaviors were studied to verify the
strengthening effect of micro- ribs. Compared to the scaffold
without micro-ribs, the average compressive strength of newly
developed scaffold was remarkably improved from 28.3 MP at
o45.6 MP under the porosity of 50%. Moreover, it also exhibited
more stable and longer lasting mechanical strength during
degradation in vitro. The effectiveness of micro-ribs on improving
the mechanical performance of scaffolds provided a structural
design reference for bone tissue engineering.
Chapter 2 Theoretical Aspects
19
Chapter 2 Theoretical Aspects
The success of orthopedic implants not only depends on its
mechanical properties but also on the biological osteointegration.
It is mandate to develop and evaluate new biomaterials in order to
improve and fasten the osteointegration process. Various strategies
are used to modify the implanted biomaterials for osteoconduction
mechanism and the formation of new bone tissue and bone
remodeling.
2.1 Biomaterials Background
Biomaterial is a material used in implants or medical
device, intended to interact with biological systems, [Williams,
(1987)]. Biomaterials did not become practical until the advent of
an aseptic surgical technique developed by in the 1860s. Earlier
surgical procedures, whether they involved biomaterials or not,
were generally unsuccessful as a result of infection. The earliest
successful implants were in the skeletal system. Bone plates were
introduced in the early 1900s to aid in the fixation of long bone
fractures. Many of these early plates broke as a result of
unsophisticated mechanical design; they were too thin and had
stress concentrating corners. Also, materials such as vanadium
steel, which was chosen for its good mechanical properties,
corroded rapidly in the body and caused adverse effects on the
healing processes. Stainless steels and cobalt chromium alloys in
Chapter 2 Theoretical Aspects
20
the 1930s were introduced with success in fracture fixation, and
the first joint replacement surgeries were performed [Park,
(1984)].
Polymethyl methacrylate (PMMA) became widely used
after that time for corneal replacement and for replacements of
sections of damaged skull bones. Following further advances in
materials and in surgical technique, blood vessel replacements
were tried in the 1950s and heart valve replacements and
cemented joint replacements in the 1960s. In the late of the same
year, ceramics, particularly alumina, were first introduced as
structural orthopedic biomaterials [Boutin, (1972)]. However,
limitations in processing technology, lack of quality control, high
levels of impurities and imperfections, caused a further reduction
in the strength of ceramics in tensile or shear resulted in failure in
a number of clinical cases [Holmer, et al (1993); Peiro, et al
(1991)].
Improvement in processing techniques for ceramics by
1977 resulted in smaller and less variable grain sizes and the true
chemical biocompatibility of these materials. Alumina and
zirconia have become the most popular ceramics for use in total-
joint replacement. Zirconia was introduced to reduce the risks of
component fracture and wear-particle production [Jazwari, et al
(1998)]. In total-joint arthroplasty either a polymer or another
ceramic both possible have been used.
Chapter 2 Theoretical Aspects
21
Implants constructed predominantly of ceramics,
particularly for total-knee replacement, are currently being
investigated. These designs are particularly useful in patients with
demonstrated metal sensitivities. At the opposite end of the
spectrum to the inert ceramics another category bioactive
materials that are designed to induce a reaction from the
surrounding tissue have been investigated. These bioactive
materials take advantage of the tissue‟s cellular physiology and
structural component materials to induce bone remodeling,
growth, and integration into the implant. An ideal bioactive
ceramic would actually promote integration of the bone with the
implant structure, and gradually biodegrade as healthy bone tissue
replaces the artificial structure [Brunski, (1996)].
Two general categories of bioactive ceramics have been
developed: calcium-based ceramics, such as calcium phosphate,
calcium sulfate, and hydroxyapatite; and bioglasses, mineral rich
structures that can be tailored to optimize the tissue response.
These bioactive materials can have either osteoinductive or
osteoconductive properties. The former refers to the ability of a
material to trigger bone cell differentiation and remodeling in
locations where bone cell proliferation and healing would not
normally occur (such as a large defect), whereas the latter
promotes bony ingrowth and vascularization, allowing for
integration and remodeling to take place. Calcium-based
composites have been used and the theory behind their use is that
Chapter 2 Theoretical Aspects
22
the body will see these materials as tissues that need to be
remodeled, allowing them to be integrated with and then replaced
by bone. Tricalcium phosphate [TCP, Ca3(PO4)2], calcium sulfate
[plaster of paris, CaSO4], and hydroxyapatite [Ca10(PO4)6(OH)2]
are all currently being used to fill bony deflects and stimulate or
direct bone formation [Tay, et al ( 1999)]. Hydroxyapatite has
also been combined with polymethyl methacrylate bone cement
with the goal of inducing bone growth into the cement. Bioglass
was introduced to the scientific world in the late 1960s by Dr.
Hench. These glass ceramics with varied proportions of SiO2,
Na2O, CaO, P2O5, CaF2, and B2O3, were designed to interact with
the normal physiology of bone to allow strong bone bonding
[Ducheyne, et al (1985)].
Initial work by [Greenspan and Hench (1976)] indicated
that an alumina implant coated with Bioglass showed substantially
improved attachment to bone and new bone formation when
implanted in rats compared with alumina-only controls. The
bonding mechanism was found to depend on the composition of
the glass, and this has sparked the development of other variations
of glass-ceramics. These include Ceravital (which contains K2O
and MgO in place of CaF2 and B2O3) Ducheyne, et al (1985) and
a form containing apatite and wollastonite [Nishio, et al (2001)].
Glass composites have also been investigated to reinforce the
glass-ceramic. The goal of these composites is to increase their
resistance to fracture by blunting crack growth and introducing a
Chapter 2 Theoretical Aspects
23
residual compressive stress within the material [Ducheyne, et al
(1985)].
2.2 Tissue engineering
The organs replacement has been subjected for researcher
interest, however, two decades ago the tissue engineering field in
vitro and in vivo to repair damaged tissue has been originated. In
fact, reconstructive surgery was the cause for development of
tissue engineering (TE) direct transplantation of (allogenic) donor
tissue is practiced to repair the function of damaged tissue. Many
difficulties arise with direct transplantation due to insufficient
donor organs, pathogen transmission and rejection of the donor
organ [Saltzman, (2004) and Badylak, (2007)]. Therefore,
patients can wait for an organ donor for years, and when they find
donor in time, they should take immunosuppressive medication
for the rest of their lives and risk the need of a replacement organ
from days to years after the surgery.
Autogenic tissue engineering transplant from patient‟s own
cells would overcome most limitations of direct transplantation
and avoid problems concerning rejection and infection. Moreover,
Autogenic tissue engineering transplant is not depends on the
donors. Therefore, an excellent alternative of direct transplantation
of donor organs is constructing a tissue engineered replacement in
vitro [Saltzman, (2004), Blitterswijk , (2008), Ross , (1998) and
Kohane, (2008)].
Chapter 2 Theoretical Aspects
24
Definition
TE is defined as the interdisciplinary field applying the
principles and methods of engineering and life sciences to
fundamentally understand and develop biological substitutes to
restore, maintain or improve tissue functions [Saltzman, (2004)].
In basis, TE attempts to mimic the function of natural tissue.
Therefore, to optimize the development of functional biological
substitutes, the natural circumstances of the specific tissue have to
be fundamentally understood. Biological tissues basically consist
of cells, signaling systems and extracellular matrix (ECM). The
cells are the core of the tissue; however, cells can't function in the
absence of signaling systems and/or of the ECM. The signaling
system consists of genes that secrete transcriptional products when
differentially activated, and urges cues for tissue formation and
differentiation [Ross, (1998)].
The ECM is a meshwork‐like substance within the
extracellular space and supports cell attachment and promotes cell
proliferation [O'Brien, (2011) and David, (2008)].
TE approaches can generally be sub‐divided based on these three
phenomena, either studied single or combined [Saltzman WM,
(2004) and Blitterswijk and Thomsen (2008)]:
- Cell-based therapies
- Induction of tissue‐formation by soluble signaling factors
- And/or biocompatible support by an artificial ECM (scaffold)
Chapter 2 Theoretical Aspects
25
The driving force behind tissue engineering is the desire to avoid
these problems by creating biological substitutes capable of
replacing the damaged tissue.
Nowadays, damaged tissue can be replaced by xenografts,
allografts or autografts. A xenograft is a graft of tissue proceeding
from another species. Xenografts offer the advantage of
availability in a variety of shapes and sizes, but they also imply a
nonnegligible risk of immunological reactions and infections.
Allografts are grafts made of tissue from a human donor, usually
post-mortem. This tissue must be thoroughly sterilised in order to
avoid immunological reactions in the receiver and infections.
Their limitations include donor shortages and risks of infections as
mentioned above. Autografts are grafts made of tissue obtained
from the patient who receives the graft: a self-transplant of tissue
in other words. Autografts are in some way a gold standard
because they avoid most problems related to transfection and
rejection. They do involve significant donor site morbidity and
chronic donor shortages however. For example, in the case of
bone replacement with tissue from the iliac crest, patients often
complain of more pain in the hip area (iliac crest) than at the
implantation site [Brook, (2008)].
The idea behind tissue engineering is to create or engineer
autografts, either by expanding autologous cells in vitro guided by
a scaffold, or by implanting a cellular scaffold in vivo and
allowing the patient‟s cells to repair the tissue guided by the
Chapter 2 Theoretical Aspects
26
scaffold. In both cases, the scaffold should degrade in time with
tissue regeneration, so that once the tissue has matured the
scaffold no longer exists as such and the newly created tissue can
perform the function of the lost tissue Platel, et al (2009). This
approach avoids some of the drawbacks of the grafting techniques
discussed above. Namely, small number of cells is harvested from
the patient, thus avoiding the problems of tissue shortage and
donor-site morbidity. The cells are seeded into a scaffold which
will eventually degrade completely, thus eliminating the presence
of a foreign body at the implantation site and its consequent
chronic inflammation. Finally, the use of autologous cells avoids
problems of rejection and transfection (Fig. 2.1).
Fig. (2.1) Schematic diagram of the different phases in tissue
engineering, from scaffold fabrication and cell isolation to in vivo
implantation.
Chapter 2 Theoretical Aspects
27
Microorganisms that enter bone structures by spreading
from the bloodstream or surrounding tissues or by direct
contamination during trauma or surgery causes osteomyelitis. A
chronic osteomyelitis treatment protocol combines both surgical
removing of dead bone tissue and prolonged parenteral or oral
antimicrobial therapy [Makinen, et al (2005), Brady, et al (2008)
and Garcia et al, (2004)]. The efficiency of systemic
antimicrobial therapy is limited by poor drug accumulation in
bone tissue, an impaired local immune response, and changes in
bacterial growth rate, biofilm formation and intracellular location
of the pathogens. Thus systemic treatment should be continued for
at least six weeks, which causes serious side effects and makes
patient compliance difficult [Lazzarini, et al, (2005)]. The
production of implantable devices able to provide high levels of
antimicrobial agents for a prolonged time at the infection site and
with low level of side effects may improve the efficacy/safety
ratio of the therapeutic strategies [Kanellakopoulou, (2008) and
Lepretre, S., (2009)].
2.3. Bone and Bone Tissue Engineering
2.3.1. Bone Structure
Bone and connective tissue are the main building blocks of
the human skeletal system. Bone is made up of organic and
inorganic or mineral matter. The organic matter is concentrated in
the bone matrix, which consists mainly of 90% collagen fibers and
Chapter 2 Theoretical Aspects
28
other noncollageneous proteins [Shoichet, (2010), Hollister,
(2005) and Hutmacher, (2004)]. The mineral matter of bone is a
calcium phosphate called hydroxyapatite (HA): Ca10(PO4)6(OH)2.
The HA crystals are thought to occupy the spaces between the
collagen fibrils, although their exact shape is under discussion.
The mineral phase of bone acts as an ion reservoir and largely
determines the mechanical properties of bone. In fact, the
mechanical properties of bone result from the impregnation of the
soft organic matrix with the very hard and brittle HA [Zhang, et
al (2009), Dalton, (2009) and Mooney, et al (1996)].
Bone‟s function is both biomechanical and metabolic.
Biomechanically, bone acts to:
a) Maintain the shape of the skeleton,
b) Protect soft tissues in the cranial, thoracic and pelvic
cavities,
c) Transmit the forces of muscular contraction during
movement,
d) Supply a framework for bone marrow.
Metabolically, bone acts to:
a) serves as a reservoir for ions, especially calcium ions,
b) Contributes to the regulation of the extracellular matrix
composition.
Macroscopically, bone is made up of cortical and cancellous bone.
Cortical or compact bone is very dense and contains only
Chapter 2 Theoretical Aspects
29
microscopic channels. It forms the outer wall of bones and bears
most of the supportive and protective function of the skeleton.
Cortical bone represents 80% of the total bone mass in the human
body. Cancellous bone makes up the remaining 20% of bone mass
in the body. It consists of trabeculae which form an interconnected
lattice (Fig. 2.2). Cancellous bone can be found in vertebrae,
fracture joints, ends of long bones and in foetuses.
Mechanical properties of bone
The mechanical properties of bone can be measured by
testing whole anatomical units or specimens prepared to isolate
particular structural components. The mechanical properties of
cortical bone have been well documented. They can be measured
via traditional testing techniques such as: uniaxial compressive or
tensile testing,
Fig. 2.2 Three-dimensional reconstruction of a cross-section of a
long bone showing the cortical and cancellous regions. Adapted
from [Dehghani, and Annabi (2011)].
Chapter 2 Theoretical Aspects
30
or three or four-point bending. They can also be tested using
ultrasound techniques or micro and nanoindentation. Cortical bone
exhibits a high degree of anisotropy and values of mechanical
properties vary between animal species, bone location and testing
conditions, age and disease. Testing conditions, for example, may
vary between testing dry samples, testing wet samples at 37°C and
embedding them or not.
Table 2. 1: Mechanical properties of human cortical bone. From
[Shoichet, (2010), Sultana and Wang, (2008) and Gutsche, et al
(1996)].
Cortical Bone MPa ±SD Elastic Modulus range
(GPa )
Compression 200 ± 36 18.6 ± 28.8
Tensile Test 141 ± 28 7.1 - 28.2
Torsional Test 65 ± 9 /
Cancellous
Bone
Strength range
(MPa)
Elastic Modulus range
(MPa)
Compression 1.5 – 3.8 10-157
Measuring the properties of cancellous bone is far more complex
than in the case of cortical bone. The complexity is due to the
small dimensions of the individual trabeculae. It is speculated that
differences in moduli between cortical and cancellous bone are
entirely due to the bone mineral density. Thus, as can be seen in
Table 2.1, some authors find value of Elastic Modulus of
Chapter 2 Theoretical Aspects
31
cancellous bone as high as those for cortical bone [Lo, et al
(1996)].
2.3.2. Bone Tissue Engineering
Bone is a complex tissue with multiple cell phenotypes,
distinct tissue types, high vascularisation and which plays a very
demanding mechanical role. Furthermore, bone‟s structure is
highly anisotropic and it remodels itself along local stress field's
lines in order to optimize its properties Baker, et al (2009). Given
this scenario, tissue engineering, which would allow the body to
generate its own bone tissue, seems a sound approach to repair
bone. Despite the multiple functions bone has in the body.
Biomechanical role is the most compromised upon injury. Indeed,
the other bones in the body can compensate for the injured bone‟s
metabolic function, but if a bone broken or injured, it can no
longer support the load it is meant for, and the body remains
handicapped. Bone transplantation‟s aim is thus to restore the
biomechanical function of the injured bone.
Trabecular bone autografts are the gold standard in bone
transplantation. The high porosity of trabecular bone allows the
surrounding tissue to vascularise the graft in a matter of weeks and
grows new bone within months. Compact bone autografts offer
higher initial strength. Their vascularisation and tissue in-growth,
however, can only take place through the osteon canals, and
osteoclasts must resorb the bone in the graft before new bone can
be generated. Thus, the bone tissue engineering scaffold should
Chapter 2 Theoretical Aspects
32
ideally resemble trabecular bone‟s architecture, biochemistry and
mechanical properties [Hutmacher, et al (2004)].
Cells for bone tissue engineering should ideally be
autologous. The bone marrow is an extraordinary source for bone
regenerating cells, and many of the engineering problems
associated to their culture and expansion have been solved. The
necessary signals or soluble factors include bone morphogenetic
proteins and growth factors which promote bone growth. No
single material possesses all the criteria required for successful
bone grafting. One approach is to design composite materials that
combine the strengths of the parent phases and minimize their
drawbacks. The combination of polymeric and ceramic materials
could improve the mechanical properties of the material and
enhance its biological properties. These concepts will be
developed in detail in the following sections [Pham and Gault,
(1998), Hollister, (2005) and Rho, et al (1998)].
2.4. Scaffolds and its role in tissue engineering
A basic concept in tissue engineering is that the scaffold
performs as a transient architecture and is "foreign" to the natural
environment. In other words, it ideally disappears once its
function has been fulfilled, leaving behind a viable and functional
biological system [Wang, et al (2001)]. In the first consensus
Conference of the European Society for Biomaterials (ESB) in
1976, a biomaterial was defined as a “nonviable material used in a
Chapter 2 Theoretical Aspects
33
medical device, intended to interact with biological systems”;
however, the ESB‟s current definition is a “material intended to
interact with biological systems to evaluate, treat, augment or
replace any tissue, organ or function of the body” [Glimcher,
(1989)]. This subtle change in definition shows how the field of
biomaterials has evolved over the years, from the use of materials
that are merely interacting with the body to the materials that
actively modulate biological processes toward the goal of tissue
regeneration.
Three major classes of biomaterials which are generally
used for the preparation of scaffolds can be distinguished:
polymers and ceramics. Also hybrid systems (e.g., combination of
polymer and ceramic) can be used. Broadly speaking, the main
demands on scaffolds for TE applications are that they serve the
bulk mechanical and structural requirements of the target tissue,
and importantly, allow for tissue healing. In general, the
biomechanical properties of the construct should match those of
the surrounding tissue (e.g. relatively tough in bone, softer in
pliable tissues). Last but not least, the main key elements of
scaffolds are biocompatibility and biodegradation.
2.4.1 Biocompatibility of scaffolds
Biocompatibility is defined by Williams (1987), as “the
ability of a scaffold or matrix to perform as a substrate that will
support the appropriate cellular activity, including the facilitation
Chapter 2 Theoretical Aspects
34
of molecular and mechanical signaling systems, in order to
optimize tissue regeneration, without eliciting any undesirable
local or systemic responses in the eventual host” [Glimcher,
(1998)]. In fact, the key to understand biocompatibility is the
understanding of the mechanisms (chemical, biochemical,
physiological, physical or other types) which become operative
under the highly specific conditions associated with contact
between biomaterials and tissues of the body and what are the
consequences of these interactions [Glimcher, (1998)]. Several
natural as well as synthetic polymers with good biocompatibility
are known and FDA approved for certain applications within the
body and therefore frequently used for TE applications. However,
there is a need for improving the physical/chemical properties of
these polymers.
2.4.2 Biodegradability of scaffolds
The use of non-permanent scaffold materials that over time
are completely replaced by natural extracellular matrix is an
important theme in TE. The objective is to create a scaffold that
can persist in a robust state for sufficient time to allow for the
formation of new tissue, but that ultimately will degrade and be
replaced by this tissue Wainwright, et al (1976). Scaffolds
biodegradation was calculate in this work using the following
equation:
Chapter 2 Theoretical Aspects
35
Initial weight of the scaffold was noted as Wo and dry weight as
Wt.
Degradation % = (Wo – Wt)/Wo × 100 2.(1)
2.4.3 Preparation methods
The physical properties of scaffolds are very relevant with
respect to final application of the graft. The scaffold is meant to
provide the appropriate chemical, physical, and mechanical
properties required for cell survival and tissue formation [Baron,
(1996)].
Material chemistry together with processing route
determines to a large extent the maximum functional properties
that a scaffold can achieve as well as how cells interact with the
scaffold. Several requirements have been identified as crucial for
the production of tissue engineering scaffolds [Guo, (2001)]: the
scaffold should: (1) possess interconnecting pores of appropriate
scale to favor tissue integration and vascularization, (2) be made
from a material with controlled biodegradability and
bioresorbability so that regenerated tissue will eventually replace
the scaffold, (3) have appropriate surface chemistry to favor
cellular attachment, proliferation, and differentiation, (4) possess
adequate mechanical properties to match the intended site of
implantation and handling, (5) should not induce any adverse
response, and (6) be easily fabricated into a variety of shapes and
sizes. Bearing these requirements in mind, several approaches
Chapter 2 Theoretical Aspects
36
have been employed to fabricate scaffolds for TE applications that
can be divided into conventional and advanced methods of
scaffolds fabrication.
2.4.3.1 Conventional scaffold fabricating techniques
- Solvent casting/ particulate leaching (SCPL)
This is the oldest and still a commonly used technique to
fabricate scaffolds. This technique is based on the principle that
porogens (most commonly salt particles) are dispersed into a
polymer solution and that after evaporation of the solvent,
followed by solidification of the polymer, and dissolution of the
porogens, a highly (for high volume fractions of salt particles)
porous scaffold (also known as a foam) is created as depicted in
(Fig. 2.3). This technique is characterized by its simple operation
and adequate control over the pore size and the porosity is tailored
by the particle size and the amount of added salt particles,
respectively.
Chapter 2 Theoretical Aspects
37
Fig. (2.3) Schematic representation of solvent casting/particulate
leaching (SCPL) method. The SEM image illustrates the
morphology of a porous hydroxyapatite/PLGA scaffold obtained
using this method [Reilly, (1974)].
However, the distribution of the salt particles is often not
uniform within the polymer solution. This is because the density
of the liquid polymer solution and the solid salt are substantially
different, and the degree of direct contact between the salt
particles is not well controlled. As a result, the interconnectivity of
the pores in the final scaffold cannot be modulated well.
Moreover, the polymer solution and the salt particles are mixed in
such a way that salt particles tend to be wrapped completely by the
polymer solution. These wrapped salt particles cannot be easily
leached out with water. Thus, most porous scaffolds prepared by
SCPL method are limited to thickness of maximum 4-18 mm and
Chapter 2 Theoretical Aspects
38
contain a porosity of up to 90% with a pore size ranging from 5 to
600 μm Brodie, et al (2000).
- Melt molding
Melt molding is another method that uses the principles of
solvent casting and particulate leaching. A polymer powder is
mixed with hydrated gelatin microspheres placed into a Teflon
mold and heated above the polymer glass transition temperature.
Poly lactic glycolic acid (PLGA) is generally preferred over other
biodegradable polyesters such as poly lactic acid (PLA) or poly
glycolic acid (PGA) because it can be processed at low
temperatures. Elevated temperatures preclude the incorporation of
bioactive molecules and may result in structural changes in gelatin
that adversely affect its aqueous solubility. After heating, the
composite is placed in water, which dissolves the water-soluble
microspheres yielding a porous structure. This technique is similar
to SCPL in that the pore size and porosity are determined by the
porogen diameter and concentration, respectively. Also, like other
leaching methods, pores do not have uniform diameters and
incomplete porogen removal is probable. In favor of the method,
melt molding does avoid the use of (toxic) solvents.
- Gas foaming
This method to fabricate porous scaffolds was first
introduced by Mooney et al [Bertram and Swartz (1991)]. In this
technique, a foaming agent such as sodium bicarbonate is added
Chapter 2 Theoretical Aspects
39
into the polymer phase to generate a gas such as N2 or CO2 when
exposed to acidic solutions [Yaszemski, et al (1996)]. A porous
structure is formed when the dispersed particles are converted into
a gas due to the exposure to an acidic aqueous solution. During the
formation of the polymeric foam, however, the liquid phase tends
to drain downwards while the gas tends to move upwards, which
leads to the formation of inhomogeneous foam with a non-porous
bottom layer and highly porous top surface.
- Emulsification/ Freeze drying
Following this technique, first a polymer is dissolved into a
suitable, water-immiscible organic solvent and then a small
volume of water is added to the polymeric introduction solution
and the two liquids are mixed in order to obtain a w/o emulsion
(Fig. 2.4). Before macroscopic phase separation occurs, the
emulsion is cast into a mold and quickly frozen (i.e. by immersion
into liquid nitrogen). The frozen emulsion is subsequently freeze-
dried to remove the dispersed water and organic solvent, yielding
a solidified, porous polymeric structure [Davies, (2000)]. The
porosity percentage of the prepared scaffolds in this study was
calculated using the following equation:
P % = [(W1-W3)/ (W2-W3)] x100 2.(2)
- Phase separation
Phase separation is another means of scaffold processing
designed with the intent of incorporating bioactive molecules. A
Chapter 2 Theoretical Aspects
40
Fig. 2.4 Schematic representation of emulsion/freeze drying
technique. The SEM image illustrates the morphology of PCL
scaffolds obtained using this method [Doherty, et al (1991)].
liquid-liquid phase separation technique has been employed to
produce foams with the potential for drug delivery [Vacanti, et al
(2000) and Burgess and Hollinger, (1998)]. As an example, poly
l-lactic acid (PLLA) and solid naphthalene are mixed in a flask,
heated, and stirred to obtain a homogenous solution. The solution
is then poured or sprayed (using an atomizer) into a cooled mold
resulting in the formation of a polymer-rich and a polymer-poor
phase. Naphthalene is subsequently removed by vacuum drying.
The foam morphology and pore distribution depend on the
kinetics of phase separation. This technique creates scaffolds with
a relatively uniform pore distribution with diameters of 50-100 μm
Chapter 2 Theoretical Aspects
41
and porosity percentage up to 87% can be achieved depending on
the polymer concentration in the solution. However, the use of
organic solvents might have detrimental effects on cells.
2.4.3.2 Advanced scaffold fabricating techniques
There has been a growing realization of the importance of
three-dimensionality in engineered tissue constructs. This interest
is largely driven by considerations such as complex issues of
nutrients and oxygen delivery and waste removal in engineered
organs (i.e. need for vascularization) [Wang, et al (2001)].
Advanced mouldless manufacturing techniques, commonly known
as solid freeform fabrication (SFF), rapid prototyping (RP), or
more colloquially art to part technology have recently been used
for fabricating complex shaped scaffolds [Hench and Polak,
(2002)]. SFF builds parts by selectively adding materials, layer-
by-layer, as specified by a computer program.
Each layer represents the shape of a cross-section of the mold at a
specific level [Liu and Ma (2004)], (Fig. 2.5).
SFF today, is considered as an efficient way of reproducibly
generating scaffolds of desired properties on a large scale fig.
(2.5). Additionally, one of the potential benefits of SFF
technology is the ability to create parts with highly reproducible
architecture and compositional variation. Rapid prototyping
techniques (as a subgroup of SFF techniques) used in tissue
engineering field can be automated and integrated with imaging
Chapter 2 Theoretical Aspects
42
techniques to produce scaffolds that are customized in size and
shape, allowing tissue-engineered implants to be tailored for
specific applications or individual patients. RP methods can be
divided into different categories:
a) Systems based on laser technology: that either
photopolymerize a liquid resin (i.e., stereolithography (SLA)) or
sinter powdered materials (i.e., selective laser sintering),
b) Systems based on print technology: including printing a
chemical binder onto a powdered material 3D-printing) or directly
printing wax (wax printing), and
c) Systems based on extrusion (also defined as nozzle-based
systems [Cooper, et al (2004)]): that process a material either
thermally or chemically as it passes through a nozzle such as 3D-
bioplotting, fused deposition modeling, and precise extrusion
manufacturing.
Chapter 2 Theoretical Aspects
43
Fig. 2.5 Tissue engineering of patient-specific implant (e.g. bone graft)
via SFF technique. CT scan data of patient's bone defect (a) are used to
generate a computer based 3D model (b) which is then sliced into
layers using rapid prototyping (RP) software. This software controls a
dispensing system (c) to deposit the polymer in a layer-by-layer fashion
(d), resulting in a well-defined 3D-structre, which will be implanted
into patient's bone defect (e). Reproduced from Ref [Hench and
Polak (2002)].
2.4.4 Scaffolds as drug delivery system.
Drug delivery is the process of administering an active
pharmaceutical ingredient in vivo to achieve a therapeutic effect in
the patient. Medication plays an important role in the medical
treatment. Most of the drugs would take their curative effect only
when their concentrations in the blood are above their minimum
Chapter 2 Theoretical Aspects
44
effective level. However, each kind of drug has its own biological
half-life and cannot maintain an effective concentration for a long
time. Merely increasing the dose of drug will extend itself into the
toxic response region, whereas taking the selected dose of drug for
several times during a period of time (e.g. three times a day) is not
convenient for the patient. In this case, drug controlled release
formulations and devices exhibit particular advantage because
they can maintain the desired drug concentration in blood for a
long period of time without reaching a toxic level or dropping
below the minimum effective level. The drug controlled release
system such as micelle, hydrogel and scaffolds has been described
in recent publications [Enrica, et al (2011), Brady, et al (2008)
and Lazzarini, et al (2005)]. Drug release mechanism from the
prepared scaffolds in this study was investigated using the following
equation:
Mt/ M∞= Ktn 2.(3)
Where, Mt / M∞ is fraction of drug released at time t, k is the rate
constant and n is the release exponent.
2.5. Biomaterials for tissue engineering applications
A biomaterial is a “material intended to interface with
biological systems to evaluate, treat, augment or replace any
tissue, organ or function of the body” [Schmidt and Baier,
(2000), Hench and Polak, (2002) and Griffith, (2002)].
Biomaterials have evolved during the past 50 years, and can now
Chapter 2 Theoretical Aspects
45
be considered “third-generation biomaterials”. Initially,
biomaterials were chosen because of their biological inertness, the
goal was to minimize the body‟s immune response to the foreign
material. Though this goal is still valid today, scientists have come
to understand that complete biological inertness is synonym to
non-recognition by the body. This lack of biological recognition is
often accompanied by fibrous tissue encapsulation and chronic
inflammation, which in turn compromise the mechanical
performance and long-term biocompatibility of the prosthesis.
The second-generation biomaterials were developed seeking
to tailor or enhance biological recognition in an attempt to
improve the biomaterial-body interface Second generation
biomaterials used bioactive components that could elicit a
controlled action and reaction in the physiological environment.
Two very typical examples of these components are synthetic
hydroxyapatite and Bioglass®. Both were used as porous
scaffolds, coatings or powders, and by the mid-80s these new
bioactive materials had attained clinical use for various dental and
orthopedic applications. The biomaterial-body interface problem
was also addressed by exploiting resorbable materials, thus
eliminating the interface all together. Resorbable polymers are the
main example of these resorbable materials, namely polylactic,
polyglycolic acid and polyvinyl alcohol which decompose
hydrolytically into H2O and CO2. They are used as sutures, screws
in orthopedics and in controlled-release drug-delivery systems.
Chapter 2 Theoretical Aspects
46
Third-generation biomaterials are being designed at present,
expanding the concept of biological recognition to specific
biological recognition. Thus, third generation biomaterials aim to
stimulate precise cellular responses: interaction with distinct
integrins, stimulation of cell differentiation or the activation of
certain genes. It is also important to emphasize that these
biomaterials are being designed. That is, third generation
biomaterials are no longer borrowed from existing materials and
adapted to a medical application. Instead, they are being designed
prior to their development. In this way, the properties of
bioactivity and resorbability are being combined to create
materials capable of helping the body repair itself better or faster
than it could do on its own. Typically, biomaterials can be divided
into: polymers, metals, ceramics and natural materials. Composite
biomaterials are created by combining two or more of these fields.
The material used in this thesis is a typical third-generation
composite biomaterial. A biodegradable polyvinyl alcohol
polymer has been combined with a resorbable bioactive glass in
order to create a composite material. This composite material has
then been shaped and processed into a scaffold and loaded with
ciprofloxacin drug for tissue engineering applications. The used
materials will be described in detail in the following sections.
Chapter 2 Theoretical Aspects
47
2.5.1Polymers for tissue engineering applications
Polymers have found widespread use in biomedical
applications for more than fifty years now. Polymers classify as
the largest class of biomaterials. They often present the advantages
of degradability and easy processability with respect to ceramic or
metallic biomaterials [Chiellini, et al (2003) and Pereira, et al
(2000)]. Both natural and synthetic polymers are used for medical
applications. Natural polymers can be of both plant and animal
origin. Some examples of natural polymers derived from plants
are cellulose, sodium alginate or natural rubber. Examples of those
derived from animals are collagen, or hyaluronic acid. Natural
polymers offer the advantage of biological recognition, which
reduces problems such as platelet adhesion, and indiscriminate
protein adsorption. This makes them ideal candidates for
cardiovascular tissue engineering, where these issues are crucial
[Yamaoka, et al (1995)]. They often require chemical or physical
pre-treatment, however, to enhance their material properties,
increase their resistance to enzymatic or chemical degradation, and
reduce immunogenicity. These treatments, cross-linking with
glutaraldehyde for instance, may have toxic effects and affect cell
growth. Natural polymers may also include pathogenic impurities
and in general offer low reproducibility.
Synthetic polymers, on the other hand, offer high
reproducibility and the possibility of large-scale production, as
well as controlled mechanical and biodegradability properties.
Chapter 2 Theoretical Aspects
48
They lack, however, biological activity and may be very
hydrophobic. Some synthetic polymers include; polyethylene
(PE), Polypropylene (PP), poly(ethylene terephtalate),
polytetrafluoroethylene (PTFE), the polyhydroxyester family :
polylactic acid (PLA) and polyglycolic acid (PGA),
polyhydroxybutyrate (PHB), copolymers of PHB and
hydroxyvalerate (PHBV), polycaprolactone (PCL), polyethylene
oxide (PEO), polyanhydrides, and polyorthoesters Tang, et al
(2007).
Polyvinyl alcohol:
Among several choices of polymers, poly(vinyl alcohol)
(PVA), a water-soluble polyhydroxy polymer with CH, CH2 and
OH as side group as showed in its chemical structure (fig. 2.6),
has been frequently explored as an implant material in biomedical
applications such as drug delivery systems, dialysis membranes,
wound dressing, artificial skin, cardiovascular devices and
Fig. 2.6 The chemical structure of PVA.
orthopedics and maxillofacial surgeries when it is combined with
ceramic because of its excellent mechanical strength,
Chapter 2 Theoretical Aspects
49
biocompatibility and non-toxicity [Chiellini, et al (2003) and
Pereira, et al (2000)]. Poly (vinyl alcohol) is one of the more
widely used polymers because of its excellent mechanical
properties. It is also biodegradable under suitable conditions
[Tang, et al (2007)].
Commercial PVA is a mixture of different types of
steroregular PVA structures (isotactic, syndiotactic, and atactic).
Its steroregularity and physical and chemical properties are highly
dependent on the preparation method used. Solubility of PVA in
water depends on the degree of hydrolysis and polymerization.
Usually, PVA with hydrolysis of 98.5 % or higher can be
dissolved in water at 70 ◦C, which is a common practice for
preparing this solution. The relative viscosity of aqueous PVA
solution within the range of 1–25 % and temperature range of 10–
80 ◦C can be expressed as a linear function of concentration and
molecular weight [Mano, et al (2007)].
2.5.2 Inorganic materials in orthopedic and maxillafacial
surgeries
The market for biomaterials based treatments in orthopedics
and maxillofacial surgeries is growing at a rapid rate. While
materials intended for implantation were in the past designed to be
„bio-inert‟, materials scientists have now shifted toward the design
of deliberately „bioactive‟ materials that integrate with biological
molecules or cells and regenerate tissues [Zijderveld, (2005) and
Chapter 2 Theoretical Aspects
50
Jell and Stevens, (2006)]. In the case of bone, materials should
preferably be both osteoinductive (capable of promoting the
differentiation of progenitor cells down an osteoblastic lineage),
osteoconductive (support bone growth and encourage the ingrowth
of surrounding bone), and capable of osseointegration (integrate
into surrounding bone). Many bone substitute materials intended
to replace the need for autologous or allogeneic bone have been
evaluated over the last two decades. In general, they consist of
bioactive ceramics, bioactive glasses, biological or synthetic
polymers, and composites of these [Tsigkou, et al (2007) and
Oonishi, (1995)]. The ideal basic premise, if following the tissue
engineering paradigm, is that the materials will be resorbed and
replaced over time by, and in tune with, the body‟s own newly
regenerated biological tissue.
A wide range of bioactive inorganic materials similar in
composition to the mineral phase of bone are of clinical interest,
e.g. tricalcium phosphate, HA, bioactive glasses, and their
combinations [Saravanapavan and Hench, (2001)]. Bioactive
glasses (Ca- and possibly P-containing silica glasses), for
example, when immersed in biological fluid, can rapidly produce
a bioactive hydroxyl carbonated apatite layer that can bond to
biological tissue. Furthermore, they can be tailored to deliver ions
such as Si at levels capable of activating complex gene
transduction pathways, leading to enhanced cell differentiation
and osteogenesis [Hamadouche, et al (2001) and Hench,
Chapter 2 Theoretical Aspects
51
(1991)]. The resorption rate of bioactive glasses and bioceramics
can be tailored with crystalline HA persisting for years following
implantation, while other calcium phosphates have a greater
capacity to be resorbed but less strength for sustaining load. The
brittle nature of bioactive inorganic materials means that their
fracture toughness cannot match that of bone and on their own are
not good for load-bearing applications [Xynos, et al (2000)].
Bioactive glass
The first bioactive material described was the glass composed of
SiO2, CaO, Na2O and P2O5 by [Xynos, et al (2000)]. Bioactive
glasses have been successfully used in various clinical
applications for over 10 years [Hench and Polak, (2002)]. Then
main feature of bioactive glasses is a well-known controlled
reaction in the physiological environment, leading to the
formation of a continuous interface connecting the tissue with the
implanted material.
Bioactive glasses bond to and integrate with living bone in the
body without forming fibrous tissue around them or promoting
inflammation or toxicity [Laczka, et al (2000)]. The high
reactivity of these glasses is the main advantage for their
application in periodontal repair and bone augmentation, since the
reaction products obtained from these types of glasses and the
physiological fluids lead to the crystallization of the apatite-like
phase, similar to the inorganic component of bones in vertebrate
species. In addition, degradation ionic products, especially silica
Chapter 2 Theoretical Aspects
52
species, have shown osteoinductive properties [Salinas, et al
(2000) and Martinez, et al (2000)]. Summarizing, from a
biological and chemical point of view, silica bioactive glasses
exhibit many of the properties associated with an ideal material for
grafting and scaffolding. This feature promoted new perspectives
for SiO2- based glasses as third-generation biomaterials for bone
tissue regeneration [Hench and Polak, (2002)]. Many factors
plays an important role in the surface interaction of glasses and
glass ceramics with the surrounding medium, such as chemical
composition, surface topography, pore size, volume and chemical
structure [Li, et al (1991) , Brinker and Scherer, (1990) and
Hench and West, (1990)]. Synthesis of bioactive glass can be
made by melting and sol-gel methods. The sol-gel is more tolerant
due to the smaller particle size produce due to its higher surface
area that leads to more binding to living tissue [Zarzycki, (1997)].
In the early 1990s bioactive glasses were for the first time
prepared by the sol–gel process [Avnir, et al (1997)]. Porous
bioglasses could be prepared from the hydrolysis and
polymerization of metal hydroxides, alkoxides and/or inorganic
salts. A wide bibliography, including excellent reviews, has dealt
with this synthesis method and application, explaining how sol–
gel chemistry offers a potential processing method for molecular
and textural tailoring [Coradin, et al (2006), Avnir, et al (2006) ,
Zhong and Greenspan, (2000) and Hamadouche, et al (2001)].
Contrarily to melt-derived bioactive glasses, sol–gel glasses are
Chapter 2 Theoretical Aspects
53
not prepared at high processing temperatures. In addition, and due
to the high surface area and porosity derived from the sol–gel
process, the range of bioactive compositions is wider, also
exhibiting higher bone bonding rates together with excellent
degradation/ resorption properties [Campostrini and Carturam,
(1996) and Pope, (1997)]. During the sol–gel process, the gelling
stage occurs around room temperature. Gels, aerogels, glasses,
dense oxides, etc., can be made by sol–gel processing (Fig. 2.7),
thus facilitating the incorporation of organic and biological
molecules within the network [Nieto, et al (2009)], or even cells
within silica matrices. Moreover, sol–gel processes can be
combined with supramolecular chemistry of surfactants, resulting
in a new generation of highly ordered mesoporous materials for
biomedical applications. Mesoporous materials are excellent
candidates for controlled drug delivery systems, and a great
research effort has been carried out in this topic during the last
years [Vallet-Reg, et al (2007) and Arcos and Vallet-Reg ,
(2010)].
Chapter 2 Theoretical Aspects
54
Fig. 2.7, Diagram of Sol–gel processing methods.
Oudadesse research group, studied many compositions of
bioactive glasses in the goal to determine qualitatively and
quantitatively the real limits between different areas in the Hench's
ternary diagram. The 46S6 compound is one of these compositions
[Mamai, et al (2008), Dietrich, et al (2008) and Oudadesse, et
al (2011)]. They found that the changes of SiO2, CaO and Na2O
amounts in the bioactive glasses induce modifications in the
physicochemical properties like the temperature of vitreous
transition (Tg) and other parameters necessary for the bioactive
glasses synthesis and consequently, modifications in their general
behavior.
Chapter 2 Theoretical Aspects
55
Ciprofloxacin drug
Ciprofloxacin (1-cyclopropyl-6-fluro-1, 4-dihydro-4-oxo-7-
(1-pipera Zinyl)-3- quinoline carboxylic acid) is a fluroquinolone
derivative, widely used in osteomyelitis because of its favorable
penetration and bactericidal effect on all the probable
osteomyelitis pathogens. Ciprofloxacin act by inhibiting the
bacterial enzymes DNA gyrase and enhance also bone formation
[Nayak and Sen, (2009), A. K, Nayak, et al (2011)].
Chapter 3 Materials & Methods
56
Chapter 3
Materials & Methods
This chapter outlines the experimental methods used in the
preparation of quaternary bioactive glass with 46S6 glass
composition of SiO2–CaO Na2OP2O5 (46 % SiO2, 24% CaO, 24 %
Na2O, 6 % P2O5 wt %) by two different methods melting molding
technique and sol-gel, preparation of bioactive composites using
polymer route technique. The preparation of 46S6 composite
scaffolds with Poly vinyl alcohol polymer loaded with
ciprofloxacin drug.
3.1. Materials
The used materials in this thesis were listed in table (3.1)
Table (3.1) The used materials.
Nomenclature Molecular formula Source
Tetraethylorthosilicate
(TEOS)
C8H20O4Si
M= 208.33 , ρ=0.932
Merck
Calcium nitrate
hydrated
Ca (NO3) 2.4H2O
M=236.15 ,
Minimum Assay 99%
Fisher Scientific
Chapter 3 Materials & Methods
57
Nomenclature Molecular formula Source
Sodium nitrite
purified
NaNO2 M= 69.00 ,
Minimum Assay 98%
Laboratory Rasayan
Sodium hydroxide
pellets purified
NaOH M=40.00,
Minimum Assay 97%
Laboratory Rasayan
Ammonium Di-
hydrogen Phosphate
NH4H2PO4 M=
115.03, Minimum
Assay 98%
Oxford Laboratory
reagent
Poly vinyl alcohol
(PVA)
(C2H4O)n M=67.000 QulaiKems
Poly vinyl
pyrrolidone (PVP)
(C6H9NO)n ,
M=40.000
ALDRICH
Acetic acid CH3CO2H El Nasr Pharmaceutical Chemicals Co.
Hydrochloric acid
(HCl)
HCl El- Ghonemy Group Co.
Calcium silicate Ca2SiO4
M=233-250
Alfa Aesar
Sodium Metasilicate penta hydrate
Na2SiO3·5H2O
M=212.1
SIGMA
Chapter 3 Materials & Methods
58
Nomenclature Molecular formula Source
Trisodium trimeta
phosphate
Na3P3O9
M=305.9
SIGMA
polyethylene glycol
(PEG)
M=20.000 Fluka
Concentrated Nitric
acid
HNO3, 55 wt % Egyptian company
for chemicals and
pharmaceuticals
Sodium chloride
(NaCl).
NaCl M=40
Fluka
Sodium hydrogen
carbonate
NaHCO3 M=84 ALDRICH
Potassium chloride
(KCl)
KCl M=74.55
ALDRICH
Di-potassium hydrogen phosphate trihydrate
K2HPO4 3H2O M=228.22
Fluka
Magnesium chloride
hexahydrate
MgCl2.6H2O
M=95.211
ALDRICH
Calcium chloride CaCl2 M=111 Fluka
Chapter 3 Materials & Methods
59
Sodium sulfate Na2SO4 M=142 ALDRICH
Tris-hydroxymethyl aminomethane: (Tris)
(HOCH2)3CNH2 ALDRICH
pH standard solution,
(pH 4, 7 and 9).
------- ALDRICH
Ciprofloxacin drug ---------- Nile pharmaceutical and chemical industries company
Phosphate Buffered
Saline (PBS)
Tablets Fluka
Chapter 3 Materials & Methods
60
Part 1
Preparation of 46S6 bioactive glass by
-Melting technique and
-Sol-gel method
Chapter 3 Materials & Methods
61
In this part we address the preparation methods of bioactive glass
(46S6) by two methods the melting technique and sol-gel method.
3.2. Methods
3.2.1. Preparation of Bioactive glass
a) Melting molding technique
Three starting materials were used, Calcium silicate, Sodium
Meta silicate penta hydrate pre heated at 200ºC/2h, and
trisodium trimeta phosphate , weighing and mixing of the
starting materials by mechanical mixer for 2h.
The batch melted in Rh-Pt crucible through the following firing
regime: heating up to 900ºC/1h with rate of 10ºC/min, firing at
1300ºC/3h with rate of 20ºC/min.
Pouring the melted glass in a pre heated molds at 500ºC as
shown in fig. (3.1),
The resulted glass was crushed and ground in mechanical agate
mortar and sieved to the grain size of less than 63 µm, the glass
given the code BM for simplicity.
Fig. (3.1), The melted 46S6 bioactive glass after removed from the molds.
Chapter 3 Materials & Methods
62
b) Sol-gel method
The sol-gel of the glass composition of the same glass
composition previously prepared by melting was performed as
demonstrated in Fig. (3.1b); initially,
Hydrolysis of 56 ml of tetraethoxysilane (TEOS) in 350 ml
of distilled water and 350 ml of ethanol at room
temperature.
pH was adjusted at 2 by nitric acid with continuous stirring
for 1 h,
Addition of 24.5 g of calcium nitrate hydrate to the above
solution continues stirring till dissolving.
Addition of 21.07 g of sodium hydroxide to the above
mixture (the previous mixture was named solution A), 10 g
of polyethylene glycol was dissolved in 400 ml of distillate
water at room temperature.
3.43 g of ammonium dihydrogen phosphate was added to
the PEG solution (this mixture was named solution B).
Furthermore, solution (B) was gradually added on solution
(A) with continuous stirring for overnight as shown in fig.
(3.2).
The resulted sol-gel was filtrated and washed with distillate
water for 3 times and with ethanol for 1 time using
centrifuge with 1650 rpm for 10 min.
Drying of the washed gel at 700C for overnight.
Chapter 3 Materials & Methods
63
Different sintering temperatures were applied on the dried powder
to assess the influence of heat treatment on the preparation of
46S6.
Fig. (3.2), Schematic diagram for preparation of 46S6 bioactive glass by solgel
method.
Chapter 3 Materials & Methods
64
Part 2
Preparation of biocomposites using
polymer technique
Chapter 3 Materials & Methods
65
This part includes the preparation of biocomposite in situe using
polymer rout technique.
3.3. Polymer Technique
Using of PVA or PVP
Dissolve 12.5 g of PVA, or PVP in distillate Water, continue
stirring for one hour at 80ºC, furthermore,
pH was adjusted at 2, addition of 40 mls of tetraethyl
orthosilicate continue stirring for one hour at room temperature,
Addition 17.5 g of calcium nitrate hydrate continue stirring till
complete dissolving, addition of 13.25g of sodium nitrite to the
above mixture,
Aaddition of 2.4 g ammonium dihydrogen phosphate to the
above mixture continue stirring till gel formation;
approximately for 1h,
Ultrasonication of the resulted precipitate for 1h, drying for
three days at 500ºC, the prepared samples were given the codes
PVA biocomposite and PVP biocomposite.
Chapter 3 Materials & Methods
66
Part 3
Preparation of the composite scaffolds
by freeze drying technique
Chapter 3 Materials & Methods
67
This part includes the composite scaffolds preparation using the
prepared 46S6 bioactive glass by melting technique or by sol-gel
method with PVA polymer through freeze drying technique.
3.4. Scaffolds preparation
Bioactive glass (MB or SG-B) / PVA scaffolds were prepared with
polymer concentrations (15%W/V). The prepared compositions
were prepared according to the proportions listed in table (3.2)
Table (3.2): The different compositions of the prepared composite
scaffolds.
Composition code MB SG-B PVA
PVA ------ ------ 100%
1PVA:2MB 66.5% ------ 33.5%
1PVA:1MB 50% ------ 50%
2PVA:1MB 33.5% ------ 66.5%
1PVA:2SG-B ------ 66.5% 33.5%
1PVA:1SG-B ------ 50% 50%
2PVA:1SG-B ------ 33.5% 66.5%
The composition 1PVA: 2BG (MB or SG-B) were loaded with
different drug concentrations (5, 10 and 20 wt %)
Chapter 3 Materials & Methods
68
Preparation method as follow:
PVA/BG-Cip composite scaffolds were prepared by employing
thermally induced phase separation technique (freeze drying) as
demonstrated in fig. (3.2).
Firstly, PVA (ALDRICH, Mwt= 67.000) was dissolved in
distilled water at 80oC for 2hr using a polymer concentration
of 15 wt%.
Three different concentrations of MB 33.5,50 and 66.5 wt%,
were added to the PVA solution and continue stirred for
overnight using a magnetic stirrer in order to break the BG
agglomerates and ensure a better (homogenous) distribution of
MB and SG-B particles in the composite scaffolds.
Three different concentrations of ciprofloxacin 5, 10 and 20
wt% were added to the above mixture continue stirred for 1hr
(scaffolds with the same composition was prepared without
drug loading as a control).
Scaffolds were casted in 24 well plates and kept at -18oC for
overnight,
and freeze dried for 24 hr then the scaffolds were removed
from the well plates and kept in the dissector for further
analysis as mentioned below.
Chapter 3 Materials & Methods
69
Fig. 3.3, Schematic diagram for PVA/BG-Cip preparation method.
3. 5. Preparation of Simulated Body Fluid
Two solutions were prepared
a) Ca- SBF solution
6.057g of Tris buffer were dissolved in 950 mls of distilled
water, addition of 0.5549 g of CaCl2 on the above solution.
Addition of 0.6095g of MgCl2.6H2O on the above mixture and the
pH was adjusted at 7.4 using HCl 6molar solution and the
temperature was adjusted at 37.5ºC throughout the whole
experiment. Complete the above solution by distilled water till
1000 mls.
PVA dissolving in distilled water at 80 оC
Addition of BG on the dissolved PVA continues stirring
The resulted mixture was casted in 24 well plates
Freezing at 18 оC for igh
Lyophilization at 56 оC
Drug addition on the above mixture continues
Chapter 3 Materials & Methods
70
b) P- SBF solution
6.057g of Tris buffer were dissolved in 950 mls of distilled
water, addition of 0.4566 g of KH2PO4.3H2O on the above
solution. 0.7056g of NaHCO3 was added on the above mixture.
0.4475g of KCl was added on the above mixture. 16.1061g of
NaCl was added on the pH was adjusted at 7.4 using HCl 6molar
solution and the temperature was adjusted at 37.5ºC throughout the
whole experiment. Complete the above solution by distilled water
till 1000 mls.
3.6. Characterizations techniques
3.6.1. Differential thermal analysis by (DSC)
The crystallization kinetics of the major phase was investigated
by differential scanning calorimetry (DSC). Isothermal
measurements were conducted on 10 mg of powder. The samples
were subjected to heating at 15 ºC min-1.
Fig. 3.4, DSC/TG instrument and some of its results.
Chapter 3 Materials & Methods
71
3.6.2 Elemental composition analysis (XRF)
The elemental composition of bioactive
glass particles was confirmed by X-ray
fluorescence spectroscopy (XRF)
(PW2404, PHILIPS).
3.6.3. Transmission electron microscope (TEM)
This technique is used to measure the morphology, surface
structure and size of samples.
Fig . 3.6, TEM device and some of its results.
The samples were suspended in acetone, dispersed
ultrasonically to separate individual particles, and one or two
drops of the suspension deposited onto holey carbon coated
copper grids. High resolution electron microscopic (HREM) and
bright field images were collected using a JEOL JEM-3010
transmission electron microscope operated at 300 kV.
Fig 3.5, XRF instrument
Chapter 3 Materials & Methods
72
3.6.4. Particle size distribution and charge by Zetasizer
The particle size and the charge of the prepared PVA and PVP
biocomposites with reference to MB bioactive glass were
determined using Zetasizer (nano ZS), through dynamic light
scattering technique.
Fig . 3.7, Zetasizer device and some of its results.
3.6.5. Morphological and microstructural properties
The microarchitecture of prepared powder and scaffolds was
assessed qualitatively using (a) scanning electron microscopy
(SEM) and (b) quantitatively using (c) mercury intrusion
porosimetry (MIP) and liquid displacement method.
3.6.5.a. Scanning Electron Microscope (SEM)
SEM analyses were performed on glass powder or thin piece of
scaffold sheared from the center using a sharp razor blade after
soaking in liquid nitrogen for 2 minutes. Scaffolds were observed
using (max. of 20 kV) SEM with gold palladium coating to avoid
damage of the polymer that could takes place by the beam, which
Chapter 3 Materials & Methods
73
can be prominent on these scaffolds that have very fine
microstructure.
Fig. 3.8, SEM device and some of its results.
3.6.5.b. Mercury Intrusion Porosimeter (MIP)
MIP was performed using (PORESIZER 9320
V2.08) to determine median pore diameter, and
percentage porosity.
3.6.5.c. Liquid displacement method
Scaffold samples were submersed in cyclohexan for 1 hr. The
volume of a scaffold immersed in the fluid is equal to the volume
of the displaced fluid, and we can calculate the porosity
percentage could be calculated using the equation follow:
P % = [(W1-W3)/ (W2-W3)] x100 3.(1)
Where W1: weight of the scaffold before immersion, W2: weight
of the scaffold after immersion and W3: weight after drying.
Fig. 3.9, MIP device
Chapter 3 Materials & Methods
74
3.6.6. Mechanical properties of the prepared scaffolds
Bones are often submitted to compression stress in the body. It has
been broadly accepted by the research community to perform
compression assays for evaluating
biomaterials for potential use as
bone repair. For that reason, the
mechanical behavior of the
composites was evaluated by
compression tests. Specimens
were evenly cut from the most
homogeneous region of the foam
to form blocks measuring 10 × 10
× 10mm3. These samples were positioned between parallel plates
using equipment EMIC DL 3000 and compressed with a
crosshead speed of 0.5mm·min−1 and a 1.0 kN load cell. At least
three samples (n = 3) of each hybrid system were measured and
the results were averaged. Compressive strength tests were carried
out to determine the effect of bioactive glass and the drug
concentrations on the mechanical strength of scaffolds.
3.6.7. Bioactivity Assessment
The bioactivity test of the prepared glass powder by both
methods was tested in SBF. 30 mg of glass powder was
submerged in 60 ml of SBF in incubator at 37◦C with 50 rpm of
oscillation for different time intervals. The SBF solution was
Fig. 3.10 Universal testing machine
Chapter 3 Materials & Methods
75
filtrated and powder was collected and washed. The washed
powder was dried at 50◦C and was subjected for the following
characterization techniques.
3.6.7.a. Phase analysis by X-ray diffraction (XRD)
X-ray diffraction (XRD) technique (PhilipsX’Pert-MPD system
with a CuK wavelength of 1.5418Å) was used to analyze the
structure of the prepared BG and the prepared composite
scaffolds. The diffractometer was operated at 40 kV and 30 mA at
a 2ɵ range of 10–70°employing a step size of 0.058/s.
Fig. 3.11, XRD device and example of its results.
3.6.7.b. Infrared studies
Fourier transformed infrared analysis (FTIR; Nicolet Magna-IR
550 spectrometer, Madison, Wisconsin) was performed to identify
the nature of the chemical bonds between atoms. The samples
were small pellets, of 0.5 cm diameter, obtained by pressing the
scaffolds powder with KBr.
10 20 30 40 50 60 70
MB
Inte
nsi
ty (
a.u
)
2
SG-B 70°C overnight
SG-B 200°C /2h
SG-B 300°C /2h
SG-B 400°C /2h
SG-B 500°C /2h
SG-B 600°C /2h
Chapter 3 Materials & Methods
76
Fig. 3.12, FTIR instrument and example of its results.
3.6.7.c. SEM coupled with EDX
The morphology of surfaces of scaffolds was studied by using
scanning electron microscopy (SEM) (Jeol JSM 6301). It is a
technique of morphological analysis based on the principle of
electron-matter interactions. To allow surface conduction, the
scaffolds were metalized by gold-palladium layer (a few µm of
thickness) before being introduced into the analysis room. Semi
quantitative chemical analysis on scaffolds surfaces after
immersion in SBF, covered by gold-palladium layer to allow
surface conduction, was performed by energy dispersive
spectroscopy (EDS) in Jeol JSM 6400.
3.6.7.d. Inductively coupled palsma -Optical - Emission
Spectrometry ICP-OES
The concentrations of (Ca, P and Si) elements after each soaking
time in SBF were measured by using ICP-OES. This method
offers a high sensitivity, less than 1µg/g depending on the
analyzed matrix and offers a high accuracy. The principle is based
on the determination of the amount of each element present in
4000 3500 3000 2500 2000 1500 1000 500
FTIR
(%)
MB
Si-oH Si-O-Si bending
wavenumber cm-1
SG-B 600°C/2h
C=OPO+SiO
2
SiO2
PO
Chapter 3 Materials & Methods
77
solution by analyzing the intensity of the radiation emitted at the
specific elemental frequency after the nebulisation of atoms.
Fig. 3.13, ICP-OES device and example of its results.
3.6.8. Drug loaded scaffolds In vitro degradation study
The degradation pattern of the composite scaffold was studied in
phosphate buffer saline (PBS) medium at 37 ◦C. groups of
scaffolds (3 scaffolds in each) were immersed in PBS and
incubated for up to 30 days. After each period time one of the
scaffolds was washed two times by distilled water to remove ions
adsorbed on the surface and was dried. Initial weight of the
scaffold was noted as Wo and dry weight as Wt. The degradation
of scaffolds was calculated using the following formula:
Degradation % = (Wo – Wt)/Wo × 100 3.(2)
3.6.9. Ciprofloxacin release behavior
Drug incorporation into the scaffolds was investigated by means
of XRD, FTIR and SEM coupled with EDS.
Phosphate buffer solution (PBS), pH 7.4 (10 ml), previously
heated at 37⁰C, was added to test tubes containing freshly prepared
0
20
40
60
80
0 2 5 7
MB SGB
Si Conc. (P
PM)
Time(days)
Chapter 3 Materials & Methods
78
scaffolds. The tubes were kept at 37⁰C with shaking (50
oscillations min-1) and, at pre-established times, 1 ml samples of
the release medium were taken and the drug concentration was
determined spectrophotometrically at 277 nm (Jenway 6705
UV/Vis, UK). The samples were replaced with fresh buffer in
order to keep constant volume of medium. All experiments were
carried out in triplicate. Ciprofloxacin release was monitored for
360 hr.
3.6.10. Mechanism of Ciprofloxacin release
Korsmeyer–Peppas model Peppas NA (2006) was used to find out
the mechanism of drug release from the investigated scaffolds:
Mt/ M∞= Ktn 3.(3)
Where, Mt / M∞ is fraction of drug released at time t, k is the rate
constant and n is the release exponent. In case of quasi-Fickian
diffusion the value of n < 0.5, Fickian diffusion n =0.5, non-
Fickain or anomalus transport n =0.5-1.0 and Case II transport n =
1.0.
Chapter 4 Results & Discussions
79
Chapter 4
Results & Discussion
This chapter includes the results and the discussion of:
the prepared bioactive glass by melting technique and sol-gel
method. The prepared biocomposites using polymer technique as
well fabricated composite scaffolds using the prepared bioactive
glass by the two methods previously mentioned with PVA
polymer through freeze drying technique.
Chapter 4 Results & Discussions
80
Part 1
Results and discussions of 46S6
bioactive glass by
-Melting technique and
-Sol-gel method
Chapter 4 Results & Discussions
81
This part includes the characterization of the prepared
46S6 bioactive glass (MB and SG-B) before immersion in SBF by
means of DSC/TG, XRD, FTIR, XRF, and TEM. The bioactivity
was in-vitro studied in SBF by XRD, FTIR, SEM and ICP-OES.
The citotoxcicity and cellular viability were tested by MTT assay.
4.1. Characterization of 46S6 bioactive glass prepared by
melting and sol-gel methods.
a) DSC/TG analysis
DSC/TG analysis was employed to determine the thermal
characteristics of the investigated powder samples as in fig. (4.1).
MB shows three characteristics peaks: at 559ºC , as temperature of
vitreous transition (Tg), at 727.6ºC, as temperature of
crystallization (Tc) and at 1235ºC as (Tf) temperature of fusion.
The thermal stability (Thermal stability is the stability of a
molecule at high temperatures; i.e. a molecule with more stability
has more resistance to decomposition at high temperatures) = (Tc -
Tg) = (727.6 - 559) = 168.
SG-B shows three characteristics peaks at 550.6 ºC as
temperature of vitreous transition (Tg), at 605.5ºC, as temperature
of crystallization (Tc) and at 835.8ºC, as (Tf) temperature of
fusion. The thermal stability = (Tc - Tg) = (605.5- 550.6) = 54.9,
which means that sol-gel method induces decreasing of 46S6
bioactive glass thermal stability Lefebvre, et al (2007) and
Matthew, et al (2011). Variation in thermal stability between the
Chapter 4 Results & Discussions
82
MB and SG-B is due to the difference in preparation methods
which disturbs the chemical stability of the SG-B bioactive glass.
This enhances its bioactivity and will be confirmed by XRD,
FTIR, SEM and ICP-OES.
Fig. 4.1, The thermal behaviour of SG-B with reference to MB.
Chapter 4 Results & Discussions
83
b.1) XRD analysis before immersion in SBF
Fig. (4.2-a) shows the XRD patterns of the 46S6 bioactive
glass prepared by melting technique (MB) at 1350°C and by sol-
gel method at different sintering temperatures from 70°C
overnight to 600°C during 2h. The characteristic peak of MB
presents a diffraction halo between 20° and 37° (2θ) with centre at
32°.
The prepared SG-B powder at 70°C overnight show two
amorphous phases with two diffraction halos, the first one is
between 11° and 20° (2θ), the second one is between 24° and 37°
(2θ) due to the presence of water and carbon in the glass structure.
Sintering temperatures 200°C and 300°C shows an amorphous
phase with little crystallinity due to the presence of carbon in the
glass structure. Sintering temperatures 400°C and 500°C shows an
amorphous phase similar to that of MB but with another small
diffraction halo between 40° and 50° (2θ) which means that the
formed phase has not achieved stability yet. On the other side, the
prepared BG by sol-gel method at 600°C, this degree greater than
the glass transition tempreture (550°C), and lesser than
crystallization tempreture (605.5 °C). These results are clarified
DSC/TG. The formed phase by this method is identical with the
formed amorphous by melting technique as reported before for
silicate glass XRD Oudadesse, et al (2011).
Chapter 4 Results & Discussions
84
Fig. 4.2-(a), XRD of SG-B at different temperatures with
reference to MB.
b.2) Influence of the sintering temperature on the prepared
powder by sol gel method
XRD was used to assess the influence of the sintering
temperature in the prepared powder by sol-gel method as shown in
fig. (4.2-b) . It illustrates different perpetrating temperatures SG-B
600ºC, 700ºC and 750ºC with reference to MB at 1350ºC.
Sintering temperature at 700ºC during 2h shows amorphous phase
but with small halo between 40º and 50º (2θ) which means that the
10 20 30 40 50 60 70
MB
Inte
nsi
ty (
a.u
)
2
SG-B 70°C overnight
SG-B 200°C /2h
SG-B 300°C /2h
SG-B 400°C /2h
SG-B 500°C /2h
SG-B 600°C /2h
Chapter 4 Results & Discussions
85
formed phase start its changes from amorphous phase to
crystalline phase as we can see at 750ºC during 2h. It shows two
crystalline phases, the first one is Na3 (Si2 PO8) phase with
characteristic peaks at (20.98º, 23.32º, 25.05º, 25.79º, 26.75º,
28.70º, 49.30º and 50.97º(2θ)) and the second one CaCO3 phase
with characteristic peaks at (17.74º, 21.29º, 25.47º, 28.84º, 30.05º,
31.67º, 49.11º and 50.05º(2θ)).The formation of the crystalline
phase decreases the bioactivity of the bioactive glass Ma, et al
(2010), Devis, et al (2011) and Holand, et al (1985), which
confirms that the ambient sintering temperature to produce 46S6
bioactive glass by sol-gel method is 600ºC/2h.
Fig. 4.2-(b), The Influence of the sintering temperature on SG-B
with reference to MB.
10 20 30 40 50 60 70
MBInte
nsity
(a.u
)
2
SG-B 600°C/2h
SG-B 700°C/3h
SG-B 750°C/3h
CaCO3
Na3(Si
2PO
8)
Chapter 4 Results & Discussions
86
c) FTIR before immersion of MB and SG-B in SBF solution
Fig. (4.3) shows FTIR spectra of the MB and SG-B at
600°C/2h and presents characteristic silicate absorption bands,
Pereira, et al (1994, 2005).
MB shows seven obvious bands, the first band at 467 cm−1
which is characteristic to Si–O–Si bending, the second at 600 cm−1
which is characteristic to phosphate group (PO43−), the third at 740
cm−1 which is characteristic to (PO), the fourth at 945 cm−1 which
is characteristic to SiO2 stretching band, and the fifth and the sixth
at 1045 cm−1 which is characteristic to phosphate group (PO2−)
with SiO2 stretching band and the seventh at 3460 cm−1 which is
characteristic to Si-OH. The prepared glass by sol-gel includes all
above mentioned bands which confirms that the prepared powder
by sol-gel method at 600°C during 2h is 46S6 bioactive glass
Martinez, et al (2000). So it is clear that using of different
preparation methods with the same composition do not affect on
the obtained phases and FTIR charts.
Chapter 4 Results & Discussions
87
Fig. 4.3, FTIR of SG-B and MB before immersion in SBF.
d) X-rays Fluorescence (XRF) analysis
The quantitative analysis of the prepared powder samples
MB and SG-B were determined by XRF analysis. MB chemical
composition was (44.042% SiO2, 27.71% CaO, 20.62 % Na2O,
6.31 % P2O5 wt %) and SG-B was (47.49% SiO2, 24.80% CaO,
20.58 % Na2O, 6.73 % P2O5 wt %) which indicate that the two
samples have approximately the same chemical composition as
presented in table (4.1) with small fractions of impurities (0.892 %
for MB and 0.406% for SG-B), which means that SG-B have
amounted impurities less than the half of those of MB. Therefore,
it's confirmed that the prepared 46S6 bioactive glass by sol-gel
method has a higher purity and should has better bioactivity and
4000 3500 3000 2500 2000 1500 1000 500
MB
Si-oH
Si-O-Si bending F
TIR
(%
)
wavenumber (cm-1)
SG-B 600°C/2h
C=OPO+SiO
2
SiO2
PO
Chapter 4 Results & Discussions
88
biocompatibility as it will be confirmed by the bioactivity and
biocompatibility tests and as reported before Pereira, et al (1994).
Table (4.1): The chemical analysis of MB and SG-B determined by XRF analysis.
e) Morphology and particle size of bioactive glass using TEM
TEM is a powerful tool for observing the morphology and
size of nanoparticles Fig. (4.4(a,b)) shows the TEM micrographs
of the prepared MB and SG-B respectively samples, fig. (4.4-a)
shows agglomerated sphere particles while Fig. (4.4-b) SG-B
shows highly homogenous nano-spheres ranging between 40-60
nm. These differences in the homogeneity and particle size are due
to the preparation methods. In most sol–gel procedures to
synthesize glasses, the sols are formed by the hydrolysis of low
molecular weight tetraethoxysilane (TEOS), using water in the
presence of a catalyst. In the hydrolysis reaction, the alkoxide
groups are replaced with hydroxyl groups. Siloxane bonds (Si–O–
Si) are then formed during subsequent condensation. Further
condensation leads to gelation which, after drying, forms a dry gel
Hench and West, (1990). The final size of the sol–gel derived
Sample
Name
SiO2 % CaO% Na2O% P2O5%
SG-B 47.49 24.8 20.58 6.73
MB 44.04 27.71 20.62 6.31
Chapter 4 Results & Discussions
89
powder depends mostly on the type of the catalysts used, which
affect on the pH of the solution and changes the relative rates of
hydrolysis and condensation reactions Brinker and Scherer,
(1990). One-step acid catalysis bioactive glasses require long
gelation times. This allows for the aggregation and growth of
colloidal particles in the solution, leading to final products with
microscale particle sizes Webster, et al (2001) and Zhong and
Greenspan, (2000). However, in this work, two-step Acid–base
catalysis was followed. The addition of sodium hydroxide, as a
second catalyst, to the sol that was initially catalyzed by nitric acid
was found to rais the rate of condensation and decrease gelation
time to few hours. The condensation rate is proportional to [OH-]
above the isoelectric point Brinker and Scherer, (1990). In this
study, sodium hydroxide was used as source of Na2O and for
gelation to provide an environment of a pH much higher than the
isoelectric point of silica Brinker and Scherer, (1990).
Therefore, the gelation time was shortened to 36 hours (overnight
stirring and overnight drying). In our study, fast gelation time of
the sols and the addition of ethanol as dispersant prevented the
growth of colloid particles during gelation. Therefore, glass
particles of less than 100 nm (40-60 nm) were successfully
prepared using the two-step acid–base catalysis.
Chapter 4 Results & Discussions
90
Fig. 4.4-(a), TEM of MB. Fig. 4.4-(b), TEM of SG-B.
f) Bioactivity Assessment
f.1) XRD after immersion of MB and SG-B in SBF at
different periods
Fig. (4.5) shows the XRD patterns for MB and SG-B before
and after immersion in SBF at different time intervals (2, 5 and 7
days). Two days after immersion in SBF solution, two
characteristic amorphous peaks of Hydroxyapatite (HA) at
25.80°and 31.79° were observed for SG-B powder. The intensity
and degree of crystallinity of these peaks increases with the
increasing of the soaking time as we can see after five and seven
days of immersion, also it is notable that appearance of new peak
at 45.142° with intensity higher of synthetic HA which is due to
the over lapping of this peak with the peak of Rhombohedral
calcium phosphate, also the peak that appears at 56.462° is
characteristic for Rhombohedral calcium phosphate as reported
before Ra´mila, et al (2002) and Mami, et al (2008). On the
Chapter 4 Results & Discussions
91
other hand for MB after two and five days of immersion we can't
observe any characteristic peaks for HA, but for seven days we
just observe two characteristic peaks for HA at 25.80°and 31.79
which indicate that the prepared glass by sol-gel is much more
reactive and bioactive than that prepared by melting as it will be
confirmed later by ICP-OES results.
Fig. 4.5-(a), XRD for MB before and after immersion in SBF for 2, 5 and 7 days.
20 30 40 50 60 702
Synthetic HA
MB 2 days
In
ten
sit
y (
a.u
)
(211)(002)
MB 5 days
MB 7 days
Chapter 4 Results & Discussions
92
f.2) FTIR of MB and SG-B before and after immersion in SBF
for different time intervals.
Fig. (4.6(a,b)) shows the FTIR for MB and SG-B before and
after immersion in SBF at different times (2, 5 and 7 days). After
two days of immersion we can recognize the difference between
MB and SG-B, for SiO2 band at 945 cm-1 it can be noted in MB
after two days and even after seven days we can see it with low
intensity. In the other hand it can't be noted for SG-B after two
days of immersion , but after 7 days, the spectrum is quite similar
to that of hydroxyapatite except two bands located at 1620 and
3423cm-1. These absorptions are characteristic of the presence of
Fig. 4.5-(b), XRD for SG-B before and after immersion in SBF for 2, 5 and 7 days
20 30 40 50 60 70
Inte
nsit
y (a
.u)
2
Synthetic HA
SG-B 2 days
Rh
-Ca P
ho
sp
hate
(224)(111)
(200)(002)
(211) (203)(231)
SG-B 7 days
SG-B 5 days
Chapter 4 Results & Discussions
93
water related to the hygroscopic feature of the formed apatite. The
OH band at 3561 cm-1 is included in the H–O–H band at 3423 cm-
1. The high water content is probably due to the presence of strong
nucleophilic groups such as: P–OH or Ca–OH, which favor the
adsorption of the ambient humidity Ra´mila, et al (2002) and
Mami, et al (2008). Also the intensity increases for wavenumbers
at 600 and 740 cm-1 which are characteristic wavenumbers of
PO43- and PO. These bands assigned to crystalline calcium
phosphate [Vallet-Regı, et al (1999) and Garcia, et al (2004)].
However, for MB there is no notable changes for these bands,
which confirm the formation of HA layer on the surface of SG-B
faster than on MB confirmed by XRD results.
Fig. 4.6-(a) FTIR for MB before and after immersion in SBF for 2, 5 and 7 days.
4000 3500 3000 2500 2000 1500 1000 500
FT
IR (
%)
wavenumber (CM-1)
MB before immersion
O-H C=OSiO
2
Si-O-Si bending
MB After 2 days
MB After 5 days
C–O PO+SiO
2
MB After 7 days
Chapter 4 Results & Discussions
94
f.3) SEM evaluation before and after immersion in SBF at
different periods
Fig. (4.7(a,b)) shows the SEM images for MB and SG-B
before and after immersion in SBF at different times (2, 5 and 7
days). For both MB and SG-B the images before immersion show
a homogeneous amorphous bulk with almost uniform particles
size. SG-B shows a small particle size less than MB (less than 100
nm). These results are in agreement with previous studies Vallet-
Regı, et al (1999) and Brunski, et al (1996). After immersion in
SBF for two days there is no notable precipitation observed for
MB but on the other hand for SG-B we can observe a huge layer
of nano-precipitated HA, which accumulate with the time to
Fig. 4.6-b: FTIR for SG-B before and after immersion in SBF for 2, 5 and 7 days
4000 3500 3000 2500 2000 1500 1000 500
FT
IR(%
)
Wavenumbers (Cm-1)
SG-B before immerssion
SG-B after 2 days
SG-B after 5 days
O-H C=OC–O
PO+SiO2
PO
PO4-3
Si-O-Si bending
SG-B after 7 days
Chapter 4 Results & Discussions
95
forming a multi layer of HA as we can see after seven days. This
is coinside and supports the results of XRD and FTIR.
Fig. 4.7, SEM micrographs; a and b for MB and SG-B before immersion in SBF, c, e and g for sample MB after immersion in SBF for 2,5, 7 days and d, f and h for sample SG-B after immersion in SBF for 2,5, 7 days.
Chapter 4 Results & Discussions
96
f.4) Chemical reactivity investigation using ICP-OES
Fig. (4.8(a, b and c)) shows the ICP diagram for the ions
concentration in SBF solution for both MB and SG-B glass
powders before and after immersion in SBF at 2, 5 and 7 days
periods. When the MB glass powder was introduced in the SBF
solution, a rapid increase of Ca ions in the solution (from 130 ppm
to 262.3 ppm) was observed; due to it is fast release of Ca ions
into the SBF. While at the same time the P ions in the SBF
solution was rapidly decrease, which delays the precipitation of
HA layer on MB glass powder because of its small surface area as
reported before Oudadesse, et al (2007 and 2011), Dietrich, et
al (2009) and Mami, et al (2008). On another side for SG-B glass
powder after two days of immersion in SBF, Ca ions fig. (4.8 (a))
increase in the solution (from 130 ppm to 199.3 ppm). However,
it's not so rapidly as in the case of MB glass powder. After five
days the Ca ions slowly decrease (from 199.3 ppm to 182.6 ppm)
and it increases again very slowly (from 182.6 ppm to 196 ppm)
after seven days, which means that there is a inverse process
between the release of calcium ions into the SBF and consuming
of the calcium ions from the SBF solution in the formation of HA
layer which is due for large surface area of SG-B. This leads to
rapid formation of HA layer on the surface of SG-B faster than
MB as confirme previously by XRD, FTIR and SEM.
These changes in the ionic concentration demonstrate the
dissolution/precipitation process, i.e., the dissolution of Ca, P and
Chapter 4 Results & Discussions
97
Si fig. (4.8 (c)), from the SG-B, and the subsequent precipitation
of Ca–P crystals from the medium, which became supersaturated
by the dissolution of Ca and P fig. (4.8 (b)). In detail, Hench and
co-workers had porposed series of reactions; the exchange of
alkali ions, such as Ca2+ and Na+ with H3O+, the attack of hydroxyl
ions (OH–) present in the medium through the silica network
structure to form silanol groups (Si–OH), and through which the
precipitation of Ca2+ and PO4 3– also CO3– takes place, followed by
the crystallization of HA Dietrich, et al (2009). But with special
note, when compared to melt-derived glasses, wherein the increase
of Ca concentration usually continues for several days to weeks,
the SG-B exhibited a inverse ions dissolution/precipitation process
as short as a few days in the Ca concentration of the medium. This
was mainly attributed to the large surface area afforded by the
nanoscale SG-B as reported before Mami, et al (2008), which
resulted in precipitation of Ca and P ions from the SBF onto SG-B
powder in the same time of the dissolution of these ions into SBF.
Chapter 4 Results & Discussions
98
Fig. 4.8-(a), Ca ions concentration after 2, 5 and 7 days of immersion in SBF.
Fig. 4.8-(b), P ions concentration after 2, 5 and 7 days of immersion in SBF.
100
150
200
250
300
0 2 5 7
SG-B MBCa Conc. (P
PM)
Time(days)
20
22
24
26
28
30
32
0 2 5 7
SG-B MB
P Conc. (P
PM)
Time(days)
Chapter 4 Results & Discussions
99
g) Cytotoxicity and cellular viability
The cell viability, bioactivity and cytotoxicity were assessed
by the MTT assay and the results are presented in histogram Fig.
(4.9). The cells treated with different concentrations of the
prepared samples showed relatively good cell viability compared
to the control used as reference. The prepared bioactive glass by
sol-gel method shows a little tendency for cell growth at higher
concentrations than those made by melting technique owing to
their higher reactivity due to its smaller particle size and higher
surface area resulted from shortened gelatin time and thermal
stability difference. The obtained results from cell proliferation are
in the same line with the obtained results from bioactivity test
Fig. 4.8-(c), Si ions concentration after 2, 5 and 7 days of immersion in SBF.
0
10
20
30
40
50
60
70
80
0 2 5 7
MB SG-BSi Conc. (P
PM)
Time(days)
Chapter 4 Results & Discussions
100
confirming the high performance of SG-B as a biomaterial
Oudadesse, et al (2011), Enrica, et al (2011) and Webster, et al
(1999 and 2000).
Fig. 4.9, The MTT assay of MB and SG-B.
100
250
305 314
100
256
325
377
0
50
100
150
200
250
300
350
400
450
0 300 600 1200SGB and MB concentration(µg /ml)
MB SGB
% of control
Chapter 4 Results & Discussions
101
Part 2
Results and discussions of the prepared
biocomposites using polymer technique
Chapter 4 Results & Discussions
102
This part includes the characterization of the prepared
biocomposites before immersion in SBF by means of DTA/TG,
XRD, FTIR, DLS and Zeta potential. The bioactivity was in vitro
studied in SBF by SEM and ICP-OES.
4.2 Characterization of polymer technique for Composites Preparation a) DTA/TG analysis
Fig. (4.10) corresponds to the thermogravimetric analysis
(TG) and differential thermal analysis (DTA) of the prepared
samples MB, PVA biocomposite and PVP biocomposite. The total
weight loss was apparently increased in the order of MB, PVA
biocomposite and PVP biocomposite from 2.8%, 50% and 52%,
respectively after heating up to 1400 ºC. This can be attributed to
the effect of polymer.
For sample MB the DTA /TG curve shows a total 6.8%
weight reduction, which can be divided in two main weight losses,
at 100 and 400 °C. They are due respectively to the departure of
free water and –OH groups. The small apparent weight loss at 610
°C may be attributed to the onset of crystallization. DTA shows an
endothermic effect at Tg1 = 550 °C caused by the glass transition,
followed by an exothermic band beginning at Tc1 = 610 °C. These
two events were already well identified in the literature El
Ghannam, et al (2001), Chatzistavrou, et al (2004) and Maria,
et al (2004). A second small endothermic effect is observed at Tg2
= 850 °C. This event had already been observed Chiellini, et al
Chapter 4 Results & Discussions
103
(2003) and Lefebrve, et al (2007). Finally, melting takes place in
the 1180– 1200°C range. Two endothermic bands (maximum
signal respectively at Tm1 = 1270 ºC and Tm2 = 1350°C) may be
attributed to the melting of two different crystalline phases.
Fig. (4.10), DTA/TG of samples (MB, PVA biocomposite and PVP
biocomposite).
0 200 400 600 800 1000 1200 1400 1600
-5
-4
-3
-2
-1
0
1
0
2
4
6
8
100 200 400 600 800 1000 1200 1400 1600
-60-50-40-30-20-10
010
0
2
4
6
8
100 200 400 600 800 1000 1200 1400 1600
-60-50-40-30-20-10
01020
0
2
4
6
8
10
MB
Hea
t F
low
(W
)
T (0C)
TG DTA
TG DTA
PVA biocomposite
TG DTAPVP biocomposite
Chapter 4 Results & Discussions
104
For samples PVA biocomposite and PVP biocomposite the
second weight loss accompanied by an endothermic peak of DTA
appeared from 200 to 500 ºC corresponding to the loss of lattice
water partly and/or hydroxyl ions (loss until 400 ºC), according to
other authors Nordstrom, et al (1990) and Mayer, et al (1997)
and eventually to carbonate decomposition. Also, the
accompanied weight loss of the second stage could be ascribed to
the polymer decomposition matrix. This peak appeared as a small
peak at 370 ºC then to 380 ºC for PVA biocomposite and PVP
biocomposite, respectively. This means that the presence of
polymer preserves the structure of bioactive glass and prevents the
removal of lattice OH to a higher temperature. This confirmed that
PVA biocomposite and PVP biocomposite have a polymer
structure dissociated and introduced CO2 in the synthesized
samples.
b) XRD before immersion in SBF Fig. (4.11(a,b)) represent the x-rays diffraction of the
prepared biocomposite with reference to 46S6 bioactive glass
prepared by melting (MB) and polymer alone before immersion in
simulated body fluid (SBF). Fig. (4.11-a) represents XRD pattern
of samples PVA biocomposite with PVA and MB XRD patterns.
While fig. (4.11-b) represents XRD pattern of PVP biocomposite
with PVP and MB.
MB shows an amorphous phase which indicated by
amorphous hump ranging from 25-35°. All polymers shows a semi
Chapter 4 Results & Discussions
105
crystalline phase ranging 17-21°, while biocomposites shows a
glassy phases with crystalline peaks which are characteristic to
Rhombohedral calcite (CaCO3). This is due to the combination
between the polymer and the glass. Also we can note that the main
characteristic peak of polymer was diminished and shifted to be
ranging from 12 to 20º, which may be due to the formation of new
phase characteristic to bioactive glass /polymer composites
Bellucci, et al (2011).
Fig. (4.11-a), XRD patterns of PVA biocomposite with reference to PVA and MB.
Chapter 4 Results & Discussions
106
c) FTIR before immersion in SBF
Fig. 4.12 (a and b) represent the FTIR of the prepared
biocomposite with reference to MB bioactive glass and PVP
polymer. FTIR spectra of the prepared MB, shows four obvious
bands, the first band at 467 cm−1 which is characteristic of Si–O–
Si stretching, the second at 600 cm−1 which is characteristic of
phosphate group (PO4−3), the third at 945 cm−1 which is
characteristic of Si-OH bond stretching, and the fourth at 1045
cm−1 which is characteristic of phosphate group (PO−2) Mamai, et
al (2008), Dietrich, et al (2008) and Oudadesse, et al (2011).
Fig. (4.11-b), XRD patterns of PVP biocomposite with reference to PVP and MB.
Chapter 4 Results & Discussions
107
FTIR spectra of the used polymers shows, band at 1165
cm−1 which is characteristic to stretching, crystalline C-O, band at
1670 cm−1 which is characteristic to bending (water molecular) O-
H, band at 2960 cm−1 which is characteristic to stretching CH,
band at 3440 cm−1 which is characteristic to stretching O-H and
adsorbed water. FTIR spectra of the prepared biocomposites
exhibits a strong band at 2850–2950 cm−1 attributed to alkyl
stretching mode (νCH). FTIR spectrum of the bioactive glass
shows the bands related to Si–O–Si asymmetric and symmetric
stretching modes. They are observed at 1080 and 450 cm−1,
respectively. The vibrational band at 950 cm−1 has been credited to
the presence of silanol groups (Si–OH) usually found in silica
synthesized via sol–gel method. FT-IR spectrum of hybrid with
composition of polymer/bioactive glass is identified by major
vibration bands, (Si–O–Si, at 1080 and 450 cm−1) Boccaccini, et
al (2010) and Almeida, et al (1990). In addition, the band at 950
cm−1 associated with the Si–OH vibrational mode remains as a
shoulder. Polymer-derived hybrid samples have also presented
broad bands in the frequency ranging from 3000 to 3650 cm−1
attributed to both contributions of hydroxyls polymer and silanols
groups of bioglass. In the range 1500–900 cm−1 there is a
superposition of the bands derived from the bioactive glass and the
polymer components Dietrich, et al (2008) and Almeida, et al
(1990). Also, typical phosphate group bands at ʎ = 1000–1220
cm−1 (PO2−, PO3
2−) and ʎ = 960 cm−1 (PO43−). The FTIR spectra of
Chapter 4 Results & Discussions
108
samples containing phosphorus showed a weak signal related to a
doublet at approximately ʎ = 565 and ʎ = 600 cm−1, which is
associated with the stretching vibrations of phosphate groups
related to the presence of crystalline phosphates in the glasses
Coates, et al (2000).
Fig. 4.12-(a), FTIR of PVA biocomposite with PVP and MB.
4000 3500 3000 2500 2000 1500 1000 500
MB
Wavenumbers (Cm-1)
PVA
PVA Biocomposite
Si-O-Si bending PO+SiO2
C–O stretchingC–O (cyclic)
C=O +C=NCH
NH-OH
FT
IR (
%)
Si-OH
Chapter 4 Results & Discussions
109
Fig. 4.12-(b), FTIR of PVP biocomposite with PVP and MB
d) Dynamic light scattering (DLS) and zeta potential
Fig. (4.13-a) shows, particles size distribution by intensity curves
for samples MB, PVA biocomposite and PVP biocomposite with
particle size distribution ranging from 600 nm to 850 nm, 500 nm
to 750 nm and from 100 nm to 600 nm respectively. These results
indicate that the presence of polymer in both biocomposite
samples reduces their particle size in comparison with MB sample
Utsel, et al (2012).
Fig. (4.13-b) shows, zeta potential distribution for samples
MB, PVA biocomposite and PVP biocomposite. MB exists in the
4000 3500 3000 2500 2000 1500 1000 500
MB
FT
IR (
%)
Wavenumber (Cm-1)
PVP
Si-OH
PVP Biocomposite
CHC–O stretching
NH-OH C=O and C=N.
C–O (cyclic)
PO+SiO2 Si-O-Si
Chapter 4 Results & Discussions
110
left side of zero axes, but in the other hand both samples PVA
biocomposite and PVP biocomposite are exists in the both sides of
zero axis, which means that MB sample carries a negative charge
on its surface, and presence of polymers in samples PVA
biocomposite and PVP biocomposite makes this samples have
both negative and positive charges Kulkarni and Wunder,
(2011).
Fig. (4.13-a), DLS of PVA biocomposite and PVP biocomposite
with reference to MB.
Chapter 4 Results & Discussions
111
Fig. (4.13-b), Zeta potential of PVA biocomposite and PVP
biocomposite with reference to MB.
e) Bioactivity Assessment
e.1) FTIR before and after immersion in SBF
Two periods were measured by FT-IR (5 and 14) days. Figures
(4.14) (a, b and c) represent FTIR results of sample MB, PVA
biocomposite and PVP biocomposite, respectively, after
immersion periods of (5 and 14) days with reference to their
curves before immersion. All FT-IR results shows an increase in
the intensity of beaks 1050 and 3460 cm-1 due to an increase in
PO2- and Si-OH groups which indicates the precipitation of Ca-P
layer on the surface of the prepared samples as it will be
Chapter 4 Results & Discussions
112
confirmed by SEM and ICP-OES results and as reported before
Boccaccini, et al (2003) and Maquet, et al (2004).
Fig. 4.14-(a), FTIR of MB before and after soaking in SBF.
Fig. 4.14-(b), FTIR of PVA biocomposite before and after
soaking in SBF.
4000 3500 3000 2500 2000 1500 1000 500
FTI
R (%
)
Wavenumbers (Cm-1)
MB before
after 5 days
Si-OH CHCH C=O Si-OH
PO2- Si-O-Si
after 14 days
4000 3500 3000 2500 2000 1500 1000 500
C–O (cyclic)
FTIR
(%)
Wavenumbers (Cm-1)
PVA biocomposite before
CHC=O +C=N
after 5 days
Si-OH
C–O stretching
PO2- Si-O-Si
after 14 days
Chapter 4 Results & Discussions
113
Fig. 4.14-(c), FTIR of PVP biocomposite before and after
soaking in SBF.
e.2) SEM before and after immersion in SBF
Figures (4.15-a, 4.15-b and 4.15-c) represents the surface
morphology of the prepared sample before immersion in SBF. MB
exhibits an amorphous bulk with almost uniform particles size. On
the other hand the presence of polymer in biocomposites samples
prefers the formation of the huge agglomerations of combined two
shapes of particles, spheres and plates ranging in nanosize, and
this is attributed to Ostwald ripening process that occurs in the
solution which leads to a typical oriented attachment process
especially on using PVP. The morphology transformation of
synthetic bioactive glass and oriented attachment process would
be accelerated under the reaction condition. Because of the effect
4000 3500 3000 2500 2000 1500 1000 500
C=O +C=NPO2
-FT
IR (%
)
Wavenumbers (Cm-1)
PVP biocomposite before
after 5 days
after 14 days
Si-OH CH C–O stretching Si-O-Si
Chapter 4 Results & Discussions
114
of the strong van der Waals attraction, the inorganic phase
(spheres like structure) tend to aggregate together with the
polymer and form side by side spheres-plates like structure along
the c axis Mamai, et al (2008), Dietrich, et al (2008) and
Oudadesse, et al (2011).
Figures (4.15-d) and (4.15-e) show the SEM results for sample
MB after immersion in SBF for 5 and 14 days, respectively, shows
a huge precipitation of crystalline hydroxyl apatite (HA). Figures
(4.15-f), (4.15-g), (4.15-h) and (4.15-i), show SEM results,
respectively, samples PVA biocomposite and PVP biocomposite
after 5 and 14 days of immersion in SBF. Amorphous layer of
calcium phosphate were precipitated on the surface of
biocomposites samples due to the fact that the presence of
polymer had delay the formation of crystalline (HA) on the
biocomposite surfaces Verrier, et al (2004), Day, et al (2005),
Jiang, et al (2005) and Jaakkola, et al (2004), as it will be
discussed and confirmed by ICP-OES results
Chapter 4 Results & Discussions
115
Fig. (4.15), SEM images for a) MB before soaking in SBF, b)
PVA biocomposite before soaking in SBF, c) PVP biocomposite
before soaking in SBF, d) MB after 5 days of soaking in SBF, e)
PVA biocomposite after 5 days of soaking in SBF, f) PVP
biocomposite after 5 days of soaking in SBF, g) MB after 7 days
of soaking in SBF, h) PVA biocomposite after 7 days of soaking
in SBF and i) PVP biocomposite after 7 days of soaking in SBF.
Chapter 4 Results & Discussions
116
e.3) Ions concentrations in SBF by ICP-OES
The solutions of the bioactivity tests were analyzed using (ICP-
OES) spectroscopy in order to determine the elemental
concentrations of Ca, Si and P ions as a function of immersion
time as demonstrated in fig. (4.16(a, b and c respectively). The
silicon concentration in the SBF increases rapidly from a value of
(0 ppm) to approximately (35 ppm). This release of silicon ions
indicates the first stage of dissolution by breaking up of the outer
silica layers of the network. The solid silica dissolves in the form
of monosilicic acid Si (OH)4 to the solution resulting from
breakage of Si–O–Si bonds and formation of Si–OH (silanols) at
the glass solution interface according to the following formula.
Si–O–Si + H2O → Si–OH + HO–Si
The curves in fig. (4.16a) presents the evolution of the
concentration of calcium shows 3 stages. A transitional stage in
which the concentration of calcium decreased between the 5th day
and the 14th day (95 PPM), then increases up to (120 PPM)
between the 15th day and the 21th day. Finally, a decrease is
observed for time between 21 and 28 days again to reach (96
PPM), reaching value of (137 PPM). Concurrent with the increase
in silicon, this result was indicated for MB sample while we note
the same result for biocomposites samples with slight differences,
Ca ions decrease till (96 PPM) which is due to presence of
polymer, also a slow consuming of P ions from SBF, therefore, we
still have P ions in SBF about (10 PPM) , which means that the
Chapter 4 Results & Discussions
117
presence of polymer in this samples delays the ions leakage from
biocomposites into SBF, due to presence of two types of charges
on the surface of biocomposites samples which is confirmed by
Zeta potential which lead to delaying in apatite layer formation on
their surfaces, as confirmed by SEM after immersion in SBF
Laczka, et al (2000), Salinas, et al (2000), Martinez, et al (2000)
and (2010).
Fig. 4.16-(a), SBF Ca ions concentrations after
soaking of the prepared samples for different periods.
0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 3090
100
110
120
130
1400 2 4 6 8 10 12 14 16 18 20 22 24 26 28 30
708090
100110120130
0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 30708090
100110120130
Ca Io
ns C
on
cen
trati
on
s (
PP
M)
Time (days)
MB
PVA biocomposite
PVP biocomposite
Chapter 4 Results & Discussions
118
Fig. 4.16-(b), SBF P ions concentrations after soaking
of the prepared samples for different periods.
Fig. 4.16-(c), SBF Si ions concentrations after soaking
of the prepared samples for different periods.
0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 3005
101520253035 0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 3005
1015202530
0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 3005
1015202530
P Io
ns C
once
ntra
tions
(PPM
)
Time(days)
MB
PVA biocomposite
PVP biocomposite
0 5 10 15 20 25 300
10
20
30
40
50
60 0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 300
10
20
30
40
500 2 4 6 8 10 12 14 16 18 20 22 24 26 28 30
0
10
20
30
40
50
Si I
on
s C
on
cen
trat
ion
s (P
PM
)
Time (days)
MB
PVA biocomposite
PVP biocomposite
Chapter 4 Results & Discussions
119
Part 3
Results and discussion of the prepared
composite scaffolds by freeze drying
technique
Chapter 4 Results & Discussions
120
This part includes the characterization of the prepared
composites scaffolds using MB or SG-B with PVA polymer
through freeze drying technique before immersion in SBF by
means of SEM, Hg-porosimeter, compressive strength by
universal testing machine, XRD and FTIR. The bioactivity was in-
vitro studied in SBF by XRD, FTIR, SEM coupled with EDX and
ICP-OES. The drug incorporation into the prepared scaffolds was
assessed by XRD, FTIR and SEM coupled with EDX. The
biodegradation rate and the drug release behavior were evaluated
in PBS.
4.3 Scaffolds Results
4.3. BG/PVA scaffold with and without drug
This part is the results and the discussion of the prepared bioactive
glass (by melting or by sol –gel methods)/ PVA polymer before
and after loading of Ciprofloxacin drug.
4.3.1. Morphological and microstructural properties
The effect of the pore size of BG particles (µm and nm-size) and
the drug presence on the properties of PVA/BG scaffolds, which
are being developed for tissue engineering applications was
studied as mention before Yazdanpanah, et al (2012). The
morphology of the prepared scaffolds is presented in fig. (4.17); in
which we can observe that all the prepared scaffolds have wide
range of interconnected pores including macro, micro and
Chapter 4 Results & Discussions
121
nanopores as it also confirmed by mercury porosimeter. PVA
scaffold shows highly interconnected pores with smooth pore
walls. As the glass content increases the porosity decreases and
the pore walls becomes thicker. Among several processing
techniques, the freeze drying method was chosen since it could
provide easy control of the pore structure Misra, et al (2007). The
co-existence of macropores and micropores is not only favorable
for the ingrowth of cells and new tissue but also beneficial to the
exchange of nutrients and metabolic waste Wong, et al (2008).
The porosity percentage for the prepared scaffolds was determined
by MIP and liquid displacement methods and there was no
significant difference between the two methods as it is
demonstrated in table (4.2). It is also noted that the particle size of
the used BG and the drug is affecting on the average pore diameter
and the array of the internal microstructural of the prepared
scaffolds Misra, et al (2007).
Chapter 4 Results & Discussions
122
Table (4.2), Porosity percentage and pore diameter of the samples
measured by Hg porosimeter and liquid displacement techniques.
Sample
nomenclature
pore
diameter
range
(4V/A)
Porosity %
µm nm MIP Liquid displacement
Without
drug
With drug
5% 10% 20%
PVA 136 6.2 88.14 85.45 72 69 66
1PVA:2MB 145 6.3 46.18 41.31 72 72 72
1PVA:2SG-B 110 6.3 46.68 60.5 69.9 68.3 67.57
4.3.2. Mechanical properties
The mechanical behavior of the prepared scaffolds was
characterized by determining the compressive strength before and
after drug incorporation. The PVA alone exhibit low compressive
strength as shown in fig. (4.18). In the produced scaffolds, a
marked change could be observed, as the amount of glass and drug
concentration increased the compressive strength increase. The
incorporation of SG-B into PVA polymer enhances the
compressive strength more than those incorporated with MB, due
to their small particle size and large surface area which results in
great attachment of SG-B particles to the polymer matrix as its
Chapter 4 Results & Discussions
123
obvious from SEM images (figure 4.17) and as reported before
Nallaa, et al (2003) and Julian, et al (2009).
Fig. 4.17, SEM images for a) PVA scaffold, b) 1PVA:2MB
scaffolds, c) PVA loaded with 20% of drug, d) 1PVA:2MB loaded
with 20% of drug e) 1PVA:2SG-B and f) 1PVA:2SG-B loaded
with 20% of drug scaffolds with magnifications of Χ15and Χ100.
Chapter 4 Results & Discussions
124
Fig. (4.18), The compressive strength of the prepared
scaffolds before and after drug incorporation.
4.3.3. XRD before immersion in SBF
The XRD result from the pure bioactive glass is as expected. It did
not show the presence of any crystalline phase, being totally
amorphous. On the other hand, the XRD patterns from both
samples of pure PVA have shown some diffraction bands. Hence,
it has being identified as a semi-crystalline structure due to the
superior concentration of hydroxyl groups. The XRD curve for
PVA/BG can be directly verified the sum up of both contributions
from PVA with semi-crystalline structure and amorphous phase of
BG María, et al (2007). It can be noted for PVA/MB scaffold,
one peak at approximately crystalline peak at 2θ of 19.8◦ (0 2 0)
and two peaks for PVA/SG-B at 2θ of 26.6◦ (113) and 33.64◦
(131), which indicated some degree of crystallinity on the
0
100
200
300
400
500
600
700
0 wt% 5wt% 10 wt% 20 wt%
PVA 1PVA:2MB 1PVA:2SG-B
Ciprofloxacin concentrations wt%
Compressive strength (M
Pa)
Chapter 4 Results & Discussions
125
biopolymer network which diminish with increase of the glass
content. That would be a typical XRD pattern for the scaffold
showing contribution from all components in the system as shown
in fig. (4.19) Hutmacher, (2000).
Fig. 4.19-(a), XRD of PVA and PVA/MB scaffolds before
immersion in SBF.
20 30 40 50 60 70
Inte
nsi
ty (
a.u
)
2
PVA
MB
2PVA:1MB
1PVA:1MB
1PVA:2MB(020)
Chapter 4 Results & Discussions
126
Fig. 4.19-(b), XRD of PVA and PVA/SG-B scaffolds before
immersion in SBF.
4.3.4. FTIR before immersion in SBF
The contribution of each and every component on the final
produced scaffold network was confirmed by FTIR as shown in
fig. (4.20). Hence, the broad band observed from 3200 to 3550
cm−1 in the PVA spectra assigned to hydroxyls (νOH) stretching
due to the strong hydrogen bond of intramolecular and
intermolecular type Tao, et al (2004) and Andrade, et al (2006).
Also, the strong band at 2870–2950 cm−1 was attributed to alkyl
stretching mode (νCH). The bands ranging from 1710 to 1750
20 30 40 50 60 70
Inte
nsi
ty (
a.u
)
2
PVA
SG-B
2PVA:1SG-B
1PVA:1SG-B
1PVA:2SG-B
(113)(131)
Chapter 4 Results & Discussions
127
cm−1 and 1200 to 1275 cm−1 arise due to the stretching vibration
of carbonyl (νC=O) and ester, respectively, from the vinyl acetate
group found in partially hydrolyzed PVA polymer . Some other
bands which can be found related to PVA are located at 1410–
1460 cm−1 assigned to δ(CH)CH2; 1200–1270 cm−1 of group
ν(C–O)–C–OH; 820–850 cm−1 from alkyl chain backbone
Mansur, et al (2004) and Mansur and Costa, (2008). In an
analogous analysis, the FTIR spectrum of the BG presented the
bands related to Si–O–Si asymmetric and symmetric stretching
modes at approximately 1100 cm−1 and 800 cm−1, respectively
Julian, (2009). There is an overlapping of the bands in the range
from 900 to 1500 cm−1 derived from the bioactive glass and the
PVA components Shin and Kim, (2001) and Mami, et al
(2008). It is worth noting that the composite formation leads to the
broadening of the bands related to vinyl acetate copolymer, that
almost disappear as a consequence of the hydrogen bonds
involving C=O groups and silanol groups in silicate networks
Oudadesse, et al (2011) and Superb, et al (2008).
Chapter 4 Results & Discussions
128
Fig. 4.20-(a), FTIR of PVA and PVA/MB scaffolds before
immersion in SBF.
4000 3500 3000 2500 2000 1500 1000 500
2PVA:1MB
FT
IR (
%)
Wavenumber (Cm-1)
PVA
MB
CHPO+SiO
2
SiO2
PO4-3
O-H C=OC–O (cyclic) PO
1PVA:1MB
1PVA:2MB
Chapter 4 Results & Discussions
129
Fig. 4.20-(b) FTIR of PVA and PVA/SG-B scaffolds before
immersion in SBF.
4.3.5. Bioactivity Assessment
a) XRD after immersion in SBF.
The XRD of the prepared scaffolds after soaking in SBF for
different time intervals are demonstrated in fig. (4.21). The
calcium phosphate layer formed on the surface PVA/MB is not
crystallized as the same with PVA/SG-B after three weeks of
immersion in the SBF. The kinetic of calcium phosphate phase on
4000 3500 3000 2500 2000 1500 1000 500
FT
IR (
%)
Wavenumber (Cm-1)
PVA
C=O
1PVA:2SG-B
1PVA:1SG-B
2PVA:1SG-B
SG-B
SiO2
CH O-H C–O (cyclic)
PO+SiO2 PO PO4
-3
Chapter 4 Results & Discussions
130
PVA/SG-B is faster than those of PVA/MB as documented before
Luo, et al (1999). Indeed after 2 days of immersion, the peaks of
crystallization related to the layer of HA formed on the surface of
PVA/SG-B starts to appear and intensity increase progressively
versus the time of immersion and BG content. After 21 days of
soaking in SBF, the XRD pattern show peaks with maximum at
about 32◦. These peaks corresponding respectively to (211), (310)
and (203) reticular plan and highlight the apatite like layer Hench,
et al (1971) and Oudadesse, et al (2009).
a) XRD of 1PVA:2MB after
immersion in SBF.
b) XRD of 1PVA:2SG-B after
immersion in SBF.
20 30 40 50 60 70
Inte
nsit
y (
a.u
)
2
Before
After 2days After week
After 2weeks
Synthetic HA
After 3weeks
After Month
(221)(022) (222)(211)
20 30 40 50 60 70
Inte
nsit
y (
a.u
)
2
Before
After 2days
After 15 days
After 7 days
Synthetic HA
After 21 days
After 30 days
(022)(221)
(222)(211)
Chapter 4 Results & Discussions
131
c) XRD of 1PVA:1MB before
and after immersion in SBF.
d) XRD of 1PVA:1SG-B before
and after immersion in SBF.
e) XRD of 2PVA:1MB before
and after immersion in SBF.
f) XRD of 2PVA:1SG-B before
and after immersion in SBF.
Fig. 4.21 (a, b, c, d, e and f), XRD of the prepared scaffolds before
and after soaking in SBF.
20 30 40 50 60 70
Inte
ns
ity (
a.u
)
2
Before
After 2days
After week
After 2weeks
After 3weeks
After Month
Synthetic HA
(221)
(211) (222)
20 30 40 50 60 70
(222)
Inte
nsit
y (a
.u)
2
Before
After 2 days
After 7 days
After 15 days
After 21 days
After 30 days
(211)
(221)
Synthetic HA
20 30 40 50 60 70
Inte
ns
ity (
a.u
)
2
Before
After 2 days
After week
After 2weeks
After 3weeks
Synthetic HA
After Month
(221)(222)(203)
(211)
20 30 40 50 60 70
Inte
nsit
y (
a.u
)
2
Before
After 2 days
After 7 days
After 15 days
After 21 days
(221)(211)
After 30 days
(222)
synthetic HA
Chapter 4 Results & Discussions
132
b) FTIR after immersion in SBF
Fig. (4.22) The IR spectrum of synthetic hydroxyapatite is used as
references to evaluate the structural evolution and the bioactivities
of the prepared scaffolds Hench, et al (2006). After soaking in
SBF solution, the initial characteristic bands of PVA/BG
biocomposite are modified strongly because of the interfacial
reactions scaffolds and the SBF. Consequently, the spectra of
these biomaterials reveal new bands.
In detail, the spectrum of PVA/BG biocomposite shows
three new well-defined phosphate bands at 565, 603 and 1039 cm-1
after 2 days of soaking in physiological solution for PVA/MB
scaffolds. They are assigned to stretching vibrations of PO43-
group in phosphate crystalline phases. On the other hand,
PVA/SG-B scaffold has the same bands with low intensity due to
great bounding of SG-B with PVA which result in slow reaction
rate between PVA/SG-B scaffolds and SBF. This result confirms
the formation of a calcium phosphate layer; this spectrum is quite
similar to that of hydroxyl apatite except two bands located at
1620 and 3423 cm-1. These bands are characterstic of the presence
of water related to the hygroscopic feature of the formed apatite.
In addition, the carbonate band at 1420 cm-1 is also observed. This
band attributes to a stretching vibration of the C-O liaisons in
carbonate groups. The presence of carbonate bands indicates the
formation of a layer of carbonated hydroxyapatite on the surface
of PVA/BG biocomposite. The obtained results highlight the rapid
Chapter 4 Results & Discussions
133
formation of apatite layer on the surface of PVA/BG
biocomposite. In addition, PVA/BG scaffolds reveal three Si-O-Si
bands at 470 cm-1 (bending vibration), 799 cm-1 (bending
vibration) and 1075 cm-1 (stretch vibration). These confirm the
presence of a silica gel Hench, et al (1996). The appearance of
apatite mineral and a silica gel indicate the interactions between
the scaffolds and SBF as described by Hench et al. This
mechanism could be explained through the following steps:
(a) Rapid exchange of protons H3O+ from the SBF with Ca2+ ,
Na+ ions in bioglass to form the Si-OH groups,
(b) Loss of soluble silica as Si(OH)4 by breaking of Si-O-Si
bridging links and subsequent formation of surface silanol groups
in the process,
(c) Condensation and repolymerization of surface silanols to
form SiO2-rich surface layer,
(d) Migration of Ca2+ and PO43- through the surface silica-rich
layer and formation of a Ca-P rich layer on the surface of
biocomposite,
(e) Incorporation of OH-, CO32- from the solution and
subsequent crystallization of the Ca-P layer to form HCA
Oudadesse, et al (2009 and 2011) and Superb, et al (2008).
The obtained results confirm the bioactivity of PVA/BG
biocomposite. Especially, they highlight the positive effect of BG
particle size and BG bounding strength with PVA which controls
the formation rate of well crystallized apatite layer on its surface.
Chapter 4 Results & Discussions
134
a) FTIR of 1PVA:2MB before
and after immersion in SBF.
b) FTIR of 1PVA:2SG-B before
and after immersion in SBF.
c) FTIR of 1PVA:1MB before
and after immersion in SBF.
d) FTIR of 1PVA:1SG-B before
and after immersion in SBF.
4000 3500 3000 2500 2000 1500 1000 500
C=OC–O (cyclic)
PO+SiO2
POPO4
-3
Synthetic HA
After Month
After 3weeks
After 2weeks
After week
After 2 days
FT
IR (
%)
Wavenumber (Cm-1)
Before
O-H
4000 3500 3000 2500 2000 1500 1000 500
PO PO4-3
Synthetic HA
After 30 days
After 21 days
After 15 days
After 7 days
After 2days
FT
IR (
%)
Wavenumber (Cm-1)
Before
O-H C=OC–O (cyclic)
PO+SiO2
4000 3500 3000 2500 2000 1500 1000 500
O-H
C=OC–O (cyclic)
PO+SiO2 PO
PO4-3
Synthetic HA
After Month
After 3weeks
After 2weeks
After week
After 2days
FT
IR (
%)
Wavenumber (Cm-1)
Before
4000 3500 3000 2500 2000 1500 1000 500
O-H C=O
Synthetic HA
After 30 days
After 21 days
After 15 days
After 7 days
After 2days
FT
IR (
%)
Wavenumber (Cm-1)
Before
C–O (cyclic)
PO+SiO2 PO4
-3
Chapter 4 Results & Discussions
135
e) FTIR of 2PVA:1MB before
and after immersion in SBF.
f) FTIR of 2PVA:1SG-B before
and after immersion in SBF.
Fig. 4.22 (a, b, c, d, e and f), FTIR of the prepared scaffolds before
and after soaking in SBF.
c) SEM with EDX after immersion in SBF
Three compositions of the prepared scaffolds (PVA, 1PVA:2MB
and 1PVA:2SG-B) have been under investigated by SEM coupled
with EDX. fig. (4.23), evaluates the surface changes of these
scaffolds after soaking in SBF for 21 days. This scaffolds had
exhibit excellent bioactivity and high fracture toughness. The
hydroxy apatite crystals formed with condensed manure on the
surface of the biocomposite scaffolds but the surface of PVA
scaffold is not changed yet. Incorporation of PVA with BG
induces a great modification to PVA bioactivity. SEM analysis
suggested the existents' of strong molecular interaction between
each type of BG particles and PVA network, causing BG to be
4000 3500 3000 2500 2000 1500 1000 500
After 2weeks
After week
After 2days
FT
IR (
%)
Wavenumber (Cm-1)
Before
PO4-3
Synthetic HA
After Month
After 3weeks
C=OC–O (cyclic)
PO+SiO2
POO-H
4000 3500 3000 2500 2000 1500 1000 500
O-H C=O C–O (cyclic)
PO+SiO2 PO
PO4-3
Synthetic HA
After 30 days
After 21 days
After 15 days
After 7 days
After 2days
FT
IR (
%)
Wavenumber (Cm-1)
Before
Chapter 4 Results & Discussions
136
dispersed uniformly in the composite scaffolds. The presence of
Ca, P, Na and Cl elements on the surface of the prepared
composite scaffolds were determined by EDX. The phosphocalcic
ratio Ca/P after 21 days of immersion in SBF is nearly equal to the
stoichiometric apatite Oudadesse, H (2011), Mami, M (2008)
and Bellucci, D (2011).
Fig. 4.23( a, b and c), SEM image of the prepared scaffolds after
immersion in SBF for 21 days.
Chapter 4 Results & Discussions
137
d) Evaluation of elemental concentrations in SBF
The change of ions concentrations in SBF was demonstrated in
fig. (4.24). For P and Si ions they take the same behavior for all
the prepared scaffolds with little difference in their amount in the
SBF. Which is due to the limit of the integrate combination
between BG and PVA. This little difference is according to
bounding and incorporation of BG into PVA. The BG particle size
is affecting on the amount of P and Si in the SBF as it's confirmed
by XRD, FTIR and SEM with EDX. The ions concentration of Ca
was found to be completely different for each composition of
scaffolds. This is much believed to be according to the glass
content in the scaffolds and the particle size of the used BG as
they in turn changes the porosity and the degradation rate in the
SBF Oudadesse, et al (2011), Mami, et al (2008) and Bellucci,
et al (2011).
a) Ca ions concentrations after soaking of PVA/MB Scaffolds in SBF for
different time intervals.
10
60
110
160
210
260
0 2 7 14 21 30
Ca ions c
onc. (P
PM)
Time (days)
PVA 2PVA:1MB
1PVA:1MB 1PVA:2MB
Chapter 4 Results & Discussions
138
b) Ca ions concentrations after soaking of PVA/SG-B Scaffolds in SBF
for different time intervals.
c) P ions concentrations after soaking of PVA/MB Scaffolds in SBF for
different time intervals.
d) P ions concentrations after soaking of PVA/SG-B Scaffolds in SBF
for different time intervals.
10
60
110
160
210
0 2 7 14 21 30
Ca ions c
onc. (P
PM)
Time(days)
PVA 2PVA:1SGB
1PVA:1SGB 1PVA:2SGB
0
10
20
30
0 2 7 14 21 30
P ions c
onc. (P
PM)
Time (days)
PVA 2PVA:1MB
1PVA:1MB 1PVA:2MB
0
5
10
15
20
25
30
0 2 7 14 21 30
P ions c
onc. (P
PM)
Time (days)
PVA 2PVA:1SGB
1PVA:1SGB 1PVA:2SGB
Chapter 4 Results & Discussions
139
e) Si ions concentrations after soaking of PVA/MB Scaffolds in SBF for
different time intervals.
f) Si ions concentrations after soaking of PVA/SG-B Scaffolds in SBF
for different time intervals.
Fig. 4.24 (a, b, c, d, e and f) ICP-OES analysis of the bioactivity
solution.
0
10
20
30
40
50
60
70
0 2 7 14 21 30
Siionsconc.(PPM)
Time (days)
PVA 2PVA:1MB
1PVA:1MB 1PVA:2MB
0
10
20
30
40
50
60
70
0 2 7 14 21 30
Siionsconc.(PPM)
Time (days)
PVA 2PVA:1SGB
1PVA:1SGB 1PVA:2SGB
Chapter 4 Results & Discussions
140
4.3.6. Ciprofloxacin incorporation
The success of incorporation of ciprofloxacin into PVA and
PVA/BG scaffolds was confirmed by XRD, FTIR and SEM
coupled with EDX.
a) XRD analysis before and after drug loading
Figures (4.25-a) and (4.25-b) represent the XRD for PVA and
PVA/BG scaffolds with and without the drug. Ciprofloxacin has
specific sharp crystal peaks while PVA, MB, SG-B and PVA/BG
have broad peaks. When ciprofloxacin was entrapped into the
scaffold matrix, its sharp crystal peaks were overlapped with the
noise of the surrounded polymer and disappeared indicating that
ciprofloxacin was successfully entrapped into the scaffold matrix
system and formation of a new solid phase for ciprofloxacin with
low crystallinity Unnithan, et al (2012), Wang, et al (2007),
Sahoo, et al (2012), Rodrı´guez-Tenreiro, et al (2004) and
Nayak, et al (2011).
Chapter 4 Results & Discussions
141
10 20 30 40 50 60 70
Inte
nsity
(a.u
)
2
PVA
SG-B
1PVA:2SG-B
Ciprofloxacin
PVA 20% Cip
1PVA:2SG-B 20% Cip
a) XRD of 1PVA:2MB with 20% of ciprofloxacin.
b) XRD of 1PVA:2SG-B with 20% of ciprofloxacin.
Fig. 4.25, XRD of the prepared scaffolds before and after
drug loading.
10 20 30 40 50 60 70
Inte
nsity
(a.u
)
2
PVA
MB
1PVA:2MB
Cip
1PVA:2MB 20% Cip
PVA 20% Cip
Chapter 4 Results & Discussions
142
b) FTIR spectra before and after drug loading
The FTIR for ciprofloxacin loaded scaffolds are demonstrated in
figures (4.26-a) and (4.26-b). The FTIR spectrum of ciprofloxacin
shows one prominent characteristic band between 3500 and 3450
cm-1, which was assigned to stretching vibration of OH groups
Another band at 3000- 2950 cm-1 represent alkene and aromatic C-
H stretching, mainly υ=C-H was demonstrated. The 1950 to 1450
cm-1 region exhibited FTIR absorption from a wide variety of
double-bonded functional groups. The band at 1750 to 1700 cm-1
represented the carbonyl C=O stretching i.e., υC=O. The band
between 1650 and 1600 cm -1 was assigned to quinolones. The
band from 1450 to 1400 cm-1 represented υC-O and at 1300 to
1250 cm-1 suggested bending vibration of O-H group which
proved the presence of carboxylic acid. A strong absorption band
between 1050 and 1000cm-1 was assigned to C-F group. The FTIR
for the PVA scaffolds loaded with ciprofloxacin indicate the
presence of new bands at 3522, 1744, and 1473.52 cm-1 when
compared with that for non-medicated scaffold due to the presence
of ciprofloxacin. These bands were indicated also for PVA/BG
scaffolds loaded with ciprofloxacin beside another band at 1088
cm-1 with high intensity due to combination of drug with glass
particles into the polymer matrix. A shorter band appeared in the
region of 1500–1200 cm-1 that could be ascribed to the hydrated
bonds with ciprofloxacin molecules Sunitha, et al (2010) and
Kesavan, et al (2010).
Chapter 4 Results & Discussions
143
The FTIR spectra indicate that, although a physical interaction
between the drug and the scaffold components occurs with both
PVA/BG scaffolds, the interaction is notably more intense with
PVA/SG-B than PVA/MB. This is probably because PVA/SG-B
has a greater content of pendant hydroxyl groups that are more
accessible for establishing hydrogen bonds with the drug Wang,
et al (2007) and Sahoo, et al (2012).
a) FTIR of 1PVA:2MB with 20% of ciprofloxacin.
4000 3500 3000 2500 2000 1500 1000 500
1PVA:2MB
FTIR
(%)
Wavenumber (Cm-1)
PVA
MB
PO+SiO2
SiO2
PO4-3O-H C=O
C–O (cyclic) PO
Cip
Cip.
Cip.
1PVA:2MB 20% Cip
PVA 20% Cip
Chapter 4 Results & Discussions
144
b) FTIR of 1PVA:2SG-B with 20% of ciprofloxacin.
Fig. 4.26 FTIR of the prepared scaffolds before and after drug
loading.
Reaction mechanism between PVA and ciprofloxacin
Scheme 4.1 suggests the reaction mechanism between scaffolds
and ciprofloxacin. The PVA could react with the drug (CIP) in
multi-position, to form the cross linking bridge using on PVA the
four active centers and for the drug the active sites A,B and C by
the condensation reaction mechanism. All probabilities are
possible. The FTIR spectra indicate that, although a physical
interaction between the drug and the scaffold components occurs
with PVA/BG scaffolds, the interaction is probably because
PVA/BG has a greater content of pendant hydroxyl groups that are
more accessible for establishing hydrogen bonds with the drug
Wang, et al (2007) and Sahoo, et al (2012).
4000 3500 3000 2500 2000 1500 1000 500
FTIR
(%)
Wavenumber (Cm-1)
PVA
C=O
1PVA:2SG-B
SG-B
SiO2
O-H C–O (cyclic)PO+SiO
2 PO PO4
-3
Cip.
Cip.
1PVA:2SG-B 20% Cip
Ciprofloxacin
PVA 20% Cip
Chapter 4 Results & Discussions
145
Fig. (4.27), Reaction mechanism PVA and ciprofloxacin.
c) SEM coupled with EDX
The SEM image of the drug shows rod shape crystals and its EDX
indicate the presence of F and Cl elements which are the main
components of the drug as demonstrated in fig. (4.28). SEM
images for the cross-section of scaffold loaded with the
ciprofloxacin reveal the rod shape of ciprofloxacin crystal in the
scaffold matrix system Nayak, et al (2011) and Puga, et al
(2012). Also the EDX confirm the presence of F and Cl elements
in the scaffolds loaded with ciprofloxacin. Therefore, XRD, FTIR
and SEM coupled with EDX indicate and confirm the success
incorporation of ciprofloxacin into PVA and PVA/BG scaffolds.
Chapter 4 Results & Discussions
146
Fig. (4.28), SEM image and EDX of a) ciprofloxacin , b) PVA
20% ciprofloxacin , c) 1PVA:2MB 20% ciprofloxacin and d)
1PVA:2SG-B 20% ciprofloxacin.
Chapter 4 Results & Discussions
147
4.3.7. Scaffolds Degradation
Biodegradation rate of the prepared scaffolds with and without
drug was investigated in PBS at different time intervals as shown
in fig. (4.29). PVA scaffold exhibits higher degradation rate
(100% after 2 days) than those of PVA/BG with and without drug
scaffolds. Ideally, in tissue engineering, a scaffold is usually
intended to temporary fill a defect, while gradually degrading as
neo-tissue is formed. In due course, the scaffold is replaced by
new bone tissue Gomes, et al (2008). After implantation, the
scaffold interacts with the tissue fluids, up taking them at some
extent, starting the degradation process Zhang, et al (2011). A
relative low degradation rate is much favorable for cell attachment
and differentiation. Furthermore, increases of the glass amount in
the scaffold decreases the degradation rate due to the fact that
incorporation of inorganic filler into polymer matrix decreases the
porosity as confirmed by mercury porosimeter and liquid
displacement methods and as documented before Wu, et al (2012)
and Peter, et al (2010). Porosity decrease lead to decreases of the
exposed area from the scaffold to the PBS. This decreasing
prolong the consumed time for biodegradation, giving more time
for cells attachment and proliferation. SG-B bioactive glass
relatively decreases the biodegradation rate of the prepared
scaffold than MB due to the great bounding ability to PVA matrix
Peter, et al (2010). Presence of the drug relatively decreases the
biodegradation rate of the prepared scaffold than BG due to the
Chapter 4 Results & Discussions
148
great bounding ability of ciprofloxacin to PVA matrix as
confirmed by XRD and FTIR.
a) Biodegradation rate PVA/MB before and after drug loading
b) Biodegradation rate PVA/SG-B before and after drug loading
Fig. 4.29 (a and b), Biodegradation rate of the prepared
scaffolds before and after drug loading.
0
20
40
60
80
100
120
2 7 15 21 30
1PVA: 2MB 1PVA: 2MB 20% Cip
PVA 20% Cip PVA
Mass loss %
Time (days)
0
20
40
60
80
100
120
2 7 15 21 30
1PVA:2SG-B 1PVA:2SG-B 20% Cip
PVA 20% Cip PVA
Mass loss %
Time (days)
Chapter 4 Results & Discussions
149
4.3.8. Release behavior of ciprofloxacin
The release behavior of ciprofloxacin from the prepared scaffolds
is presented in fig. (4.30). Considering the hydrophilic molecule,
ciprofloxacin is expected to exhibit burst release from the
investigated system. The release behavior for ciprofloxacin from
the investigated scaffold seemed to be in a sustained release
profile with Korsmeyer–Peppas model as indicated by its higher
r2-values. Furthermore, the release of ciprofloxacin from the
investigated scaffolds obeyed quasi-Fickian diffusion mechanism
(n-values less than 0.5). This mechanism is based on hydrolysis
that indicates the polymer is hydrated, swell and then the drug
diffuses through the swollen matrix system to the exterior, which
ultimately slows down the kinetic release.
Generally, incorporation of glass into scaffolds resulted in
faster amount of drug released than released from the PVA-based
scaffold. A possible explanation for this observation is that the
glass particles have equipped a huge part of the polymer matrix
which leads to less compact structure causes higher and faster
drug release. Using higher percentage of ciprofloxacin (20 %) was
affected by the particle size of the incorporated glass as we can
note that 1PVA:2SG-B scaffold has relatively high drug release
profile compared with that for 1PVA:2MB scaffold, which could
be explained due to the great surface area that provided by SG-B
nanoparticles causing fast ciprofloxacin release from the polymer
matrix. Release of the drug from PVA/BG scaffolds containing
Chapter 4 Results & Discussions
150
20% ciprofloxacin was faster than those containing 10% and 5%
ciprofloxacin. This is due to that as the drug concentration
increase free drug particles are not attached to the matrix causing
faster release. Wang, et al (2007), Sunitha and Kumar, (2010)
and Thakre and Choudhary, (2011).
The structure morphology of ciprofloxacin loaded scaffolds
during the immersion. Fig. (4.31), explains the kinetic release of
ciprofloxacin from scaffolds. These micrographs show that all
samples have porous network structure which is responsible for
their swelling. Macroscopically, all the samples appeared
transparent.
a)
10
15
20
25
30
35
40
45
50
55
0 100 200 300 400
Cumulative re
lease (%
)
Time (hr)
PVA 5%
PVA 10%
PVA 20%
Chapter 4 Results & Discussions
151
b)
c)
Fig. (4.30), The cumulative ciprofloxacin release for a) PVA
scaffolds loaded with 5,10 and 20% Cip, b) 1PVA:2MB
scaffolds loaded with 5,10 and 20% and c) 1PVA:2SG-B
scaffolds loaded with 5,10 and 20%.
0
10
20
30
40
50
60
70
80
90
0 100 200 300 400
Cumulative re
lease (%
)
Time (hr)
1PVA:2MB 5% Cip
1PVA:2MB 10% Cip
1PVA:2MB 20% Cip
15
25
35
45
55
65
75
85
95
105
0 100 200 300 400
Cumulative re
lease (%
)
Time (hr)
1PVA:2SG-B 5%
1PVA:2SG-B 10%
1PVA:2SG-B 20%
Chapter 4 Results & Discussions
152
Fig. (4.31), SEM of the prepared scaffolds after soaking in PBS.
Conclusions
153
Conclusion
From the obtained results and discussion, it can be concluded that:
Nanobioactive quaternary glass system 46S6 has been
prepared by modified sol-gel (acid-base reaction) method at
600°C with particle size ranging between 40-60 nm and a
decrease in the gelation time as confirmed by DSC/TG , XRF,
TEM, XRD and FTIR.
The formation of apatite layer over the sol-gel prepared
glass was faster than in melting bioactive glass after
immersion in simulated body fluid as valid by results of XRD,
FTIR SEM coupled with EDX and ICP-OES.
Cell viability by MTT test confirmed the effectiveness of
the prepared bioactive glass nanopowder SG-B as a bone
replacement material
The prepared biocomposite samples that have been
fabricated at low temperature have both phases of bioactive
glass (46S6) with the same concentration of constituent within
the polymer matrix as verified by XRD and FTIR.
DTA of the prepared materials has confirmed that presence
of polymer had effect on the thermal behavior of biocomposite
samples. The incorporation of bioactive glass into the polymer
phase was reviled by SEM analysis.
Conclusions
154
Dynamic Light Scattering confirmed that the presence of
polymer reduces the particle size of biocomposite samples in
comparison with MB sample; also, the presence of polymer
affects their electrical behavior as confirmed by zeta potential.
SEM, FT-IR and ICP-OES proved that the presence of
polymer in biocomposite samples had delay the ions leakage
from bioactive glass, which indicate that the prepared
composites can be used in achieving of controllable
bioactivity, by controlling the BG/ polymer ratio. Also this can
be used as a drug delivery system.
The PVA/BG biocomposite scaffolds loaded with
ciprofloxacin with well interconnected pore structure and
appropriate porosity were fabricated via freeze drying
technique as confirmed by the results of SEM, MIP and liquid
displacement method this was assured for orthopedic and
maxillofacial surgeries.
The bioactivity of the prepared scaffolds is affected by the
glass particle size and glass/polymer ratio as proved by results
of XRD, FTIR SEM coupled with EDX and ICP-OES.
The addition of drug to scaffold provided by advantageous
of mechanical properties. Meanwhile preserving the porosity
without affecting of the drug efficiency.
Conclusions
155
The physicochemical properties, biodegradation rate and
bioactivity of the prepared scaffolds could be controlled by
regulation the glass contents and the drug concentrations
Drug loaded scaffolds with ciprofloxacin exhibit a good
drug delivery system with sustained drug release pattern. The
presence of glass particles in the drug loaded scaffolds affects
the drug release behavior as confirmed by UV-
Spectrophotometer analysis.
References
156
References
[1]Almeida, R. M. and Pantano, C. G.: Structural
investigation of silica gel films by infrared spectroscopy. J
App Phy, 1990; 68, no. 8: 1-8.
[2]Andrade G., Barbosa-Stancioli E.F., Piscitelli
Mansur A.A., et al.: J. Biomed. Mater. 2006; 1: 221–34.
[3]Andrade, G., Barbosa-Stancioli, E. F., Mansur, A. A.
Piscitelli, W Vasconcelos, L. and Mansur H. S.: Design
of novel hybrid organic-inorganic nanostructured
biomaterials for immunoassay applications. Biomed Mater,
2006; 1, no. 4: 221–34.
[4]Arcos D. and Vallet-Reg M. Sol–gel silica-based
biomaterials and bone tissue regeneration, Acta
Biomaterialia 2010; 6: 2874–2888.
[5]Avnir D and Braun S. Biochemical aspects of sol-gel
science and technology: New York: Springer-Verlag; 1996;
11-16.
[6]Avnir D, Coradin T, Lev O, Livage J. Recent bio-
applications of sol–gel materials. J Mater Chem 2006;
16:1013–1030.
[7] Avnir D, Klein Lisa C, Levy D, Schubert U, Wojcik
AB. Organo silica sol–gel materials. In: Apeloig Y,
References
157
Rappoport A, editors. The chemistry of organosilicon
compounds part 2. Chichester: Wiley & Sons; 1997; 28-35.
[8]Badylak SF. The extracellular matrix as a biologic
scaffold material. Biomaterials, 2007; 28(25):3587‐3593.
[9]Baker, S. C.; Rohman, G. r.; Southgate, J.; Cameron,
N. R., The relationship between the mechanical properties
and cell behavior on PLGA and PCL scaffolds for bladder
tissue engineering. Biomaterials 2009; 30, (7): 1321-1328.
[10]Baron R. Anatomy and Ultrastructure of Bone. In:
Favus MJ, editor. Primer on Metabolic and Bone Diseases
and Disorders of Mineral Metabolism. Philadelphia:
Lippincott-Raven, 1996: 3-9.
[11]Batal H.A. El, Azooz M.A., Khalil E.M.A., Soltan
Monem A. and Hamdy Y.M., J. Mater Chem Phys
2003;80: 599–609.
[12]Bellucci, D., Cannillo, V. a, et al.: Macroporous
Bioglass®-derived scaffolds for bone tissue regeneration.
Ceram Int 2011; 37: 1575–85.
[13]Bertram JE and Swartz SM. The 'law of bone
transformation': a case of crying Wolff? Biol Rev Camb
Philos Soc 1991; 66(3):245-273.
[14]Blitterswijk CAV and Thomsen P. Tissue
engineering. 1st ed. Amsterdam; Boston:
Elsevier/Academic Press, 2008.
References
158
[15]Boccaccini AR and Maquet V. Bioresorbable and
bioactive polymer/Bioglass(R) composites with tailored
pore structure for tissue engineering applications. Compos
Sci Technol 2003; 63:2417–2429.
[16]Boccaccini R., Erol Melek , Stark Wendelin J.:et al.
,Polymer/bioactive glass nanocomposites for biomedical
applications: A review, Composites Science and
Technology, 2010; 70: 1764–1776.
[17]Boutin, P., Arthroplastie totale de la hanche par
prostheses en aluminine fritte, Rev. Chir. Orthop. 1972;
58: 230–46.
[18]Brady, RA., Leid, JG. and Calhoun, JH.:
Osteomyelitis and the role of biofilms in chronic infection.
FEMS Immunol Med Microbiol 2008; 52:13–22.
[19]Brey EM, King TW, Johnston C, McIntire LV,
Reece GP, Patrick Jr CW. A technique for quantitative
three-dimensional analysis of microvascular structure.
Microvasc Res 2002; 63:279-94.
[20]Brinker C.J. and Scherer G.W.: The Physics and
Chemistry of Sol–Gel Processing. Academic Press. Inc.,
San Diego, CA, USA, 1990; 45-52.
[21]Brinker CJ, Scherer GW. Sol–gel science. San
Diego, CA: Academic Press; 1990; 60-67.
References
159
[22]Brodie E. McKoy, Yuehuei H. An. and Richard J. et
al, Mechanical Testing of Bone and the Bone-Implant
Interface First Edition ed. Boca Raton: CRC Press, 2000;
89-105.
[23]Brook, I.: Microbiology and management of joint and
bone infections due to anaerobic bacteria. J Orthop Sci
2008; 13:160–9.
[24]Brunski J.: An Introduction to Material in Medicine.
London: Academic Press; 1996; 36-37.
[25]Burgess EA, Hollinger JO. Options for Engineering
Bone. In: Patrick CW, Mikos AG, Mc.Intre L, editors.
Frontiers in Tissue Engineering. Oxford: Elsevier Science
Ltd., 1998: 383-399.
[26]Caixia Xu, Peiqiang Su, Xiaofeng Chen, et al.
Biocompatibility and osteogenesis of biomimetic Bioglass-
Collagen- Phosphatidylserine composite scaffolds for bone
tissue engineering. Biomaterials 2011; 32: 1051-1058.
[27]Campostrini R, and Carturam G. Immobilisation of
plant cells in hybrid sol–gel material. J Sol-Gel Sci Technol
1996; 7:87–97.
[28]Cassell OC, Hofer SO, Morrison WA, Knight KR.
Vascularisation of tissue- engineered grafts: the regulation
of angiogenesis in reconstructive surgery and in disease
states. Br J Plast Surg 2002; 55:603-10.
References
160
[29]Chang Ch, Peng N, He M, et al. Fabrication and
properties of chitin/hydroxyapatite hybrid hydrogels as
scaffold nano-materials. Carbohydrate Polymers. 2013; 91:
7-13.
[30]Chatzistavrou X., T. Zorba, E. Kontonasaki, K.
Chrissafis, et al. J Phys Stat Sol (a) 201 2004; 5: 944–951.
[31]Chengtie Wua, Yufeng Zhang, Yinghong Zhou, et
al., A comparative study of mesoporous glass/silk and non-
mesoporous glass/silk scaffolds: Physiochemistry and in
vivo osteogenesis, Acta Biomaterialia. 2011; 7: 2229–2236.
[32]Chiellini E., Corti A., D’antone S., Solaro R.,
Biodegradation of poly (vinyl alcohol) based materials,
Prog. Polym. Sci. 2003; 28:963–1014.
[33]Chu TM, Warden SJ, Turner CH, Stewart RL.
Segmental bone regeneration using a load-bearing
biodegradable carrier of bone morphogenetic protein-2.
Biomaterials 2007; 28:459e67.
[34]Coates J., in: R.A. Meyers (Ed.), John Wiley and Sons
Ltd., Chichester, 2000, 10815–10837.
[35]Coates J.: Encyclopedia of analytical chemistry:
interpretation of infrared spectra, a practical approach,” in
Encyclopedia of Analytical Chemistry, R. A. Meyers, Ed.,
2000; JohnWiley & Sons, Chichester, UK, 10815–37.
References
161
[36]Cooper SL, Visser SA, Hergenrother RW, et al.: and
editors. Biomaterials Science. San Diego, London: Elsevier
Academic Press, 2004; 67-79.
[37]Coradin T, Boissière M, Livage J. Sol–gel chemistry
in medicinal science. Curr Med Chem 2006; 13:99.
[38]Dalton, P. D.; Woodfield, T., et al.: Snapshot:
Polymer scaffolds for tissue engineering. Biomaterials
2009; 30, (4), 701-702.
[39]David F, W., On the mechanisms of biocompatibility.
Biomaterials 2008; 29, (20), 2941-2953.
[40]Davies, J. E. Bone Engineering First Edition ed.
Toronto Ontario, Canada: Em squared incorporated, 2000;
78-85.
[41]Dehghani, F.; Annabi, N., Engineering porous
scaffolds using gas-based techniques. Current Opinion in
Biotechnology 2011; 22, (5): 661-666.
[42]Bellucci D, Cannelloni V, et al.: Calcium and
potassium addition to facilitate the sintering of bioactive
glasses, Materials Letters. 2011; 65:1825-1827.
[43]Dietrich E. Oudadesse H., Lucas-Girot A., Mami M.
In vitro bioactivity of melt-derived glass 46S6 doped with
magnesium. Journal of Biomedical Materials Research
2009; 88: 1087-1096.
References
162
[44]Dietrich E., Oudadesse H., Lucas-Girot A. et al,
Effects of Mg and Zn on the surface of doped melt-derived
glass for biomaterials applications. Applied Surface
Science 2008; 255: 391–395.
[45]Doherty PJ, Williams RL, Williams D, Lee AJC,
editors. Biomaterial-Tissue Interfaces: Second Consensus
Conference on Definitions in Biomaterials, Chester 1991.
Amsterdam: Elsevier, 1992;74-79.
[46]Ducheyne, P., Bioglass coatings and Bioglass
composites as implant materials, J. Biomed. Mater. Res.
1985; 19:273–91.
[47]El Ghannam A., Hamazawy E and Yehia A., J
Biomed Mater Res 2001; 55: 387– 398.
[48]El-Kady A. M., Ali Ashraf F., Fabrication and
characterization of ZnO modified bioactive glass
nanoparticles, Ceramics International, 2012; 38:1195–
1204.
[49]El-Kady A. M., Ali Ashraf F., Rizk A. Rizk, et al.
Synthesis, characterization and microbiological response of
silver doped bioactive glass nanoparticles. Ceramics
International, 2012; 38: 177–188.
[50]Enrica Saino, Stefania Grandi, Eliana Quartarone,
et al. In vitro calcified matrix deposition by human
References
163
osteoblasts onto a zinc-containing bioactive glass.
Eurupean cells and materials.2011; 21:59-72.
[51]Fuchs JR, Nasseri BA, Vacanti JP. Tissue
engineering: A 21st century solution to surgical
reconstruction. Annual Thorac Surgery 2001; 72:577‐591.
[52]Gaalen, S.M.V., Kruyt, M.C. and Meijer, G.J.:
Tissue engineering of bone, in: C.v. Blitterswijk, J.
Sohier (Eds.), Tissue Engineering, Academic Press,
Elsevier, UK, 2008; 555–606.
[53]Gaalen, S.M.V., Kruyt, M.C. and Meijer, G.J.:
Tissue engineering of bone, in: C.v. Blitterswijk, J. Sohier
(Eds.), Tissue Engineering, Academic Press, Elsevier, UK,
2008; 555–606.
[54]Garcia C., Cere S., et al.: Bioactive coatings prepared
by sol–gel on stainless steel 316L. J Non- Cryst .Solids
2004; 348:218-224.
[55]Garima T. and Bikramjit B. A porous hydroxyapatite
scaffold for bone tissue engineering: Physico-mechanical
and biological evaluations. Ceramics International 2012;
38: 341–349.
[56]Glimcher MJ. Mechanism of calcification: role of
collagen fibrils and collagen phosphoprotein complexes in
vitro and in vivo. Anat. Rec.1989; 224(2):139-153.
References
164
[57]Glimcher MJ. The Nature of the Mineral Phase in
Bone: Biological and Clinical Implications. In: Avioli LV,
Krane SM, editors. Metabolic Bone Disease and Clinically
Related Disorders. St. Louis: Academic Press, 1998: 23-50.
[58]Gomes, M.E., Azevedo, H.S., Moreira, A.R., Ellä,
V., Kellomäki, M. et al.: Starchpoly(- caprolactone) and
starch-poly(lactic acid) fibre-mesh scaffolds for bone tissue
engineering applications: structure, mechanical properties
and degradation behavior. J Tiss Engi Rege Medici. 2008;
2: 243–252.
[59]Greenspan, D. C., and Hench, L. L. (1976), Chemical
and mechanical behavior of Bioglass-coated alumina, J.
Biomed. Mater. Res. Symp. 1976; 7:503–509.
[60]Griffith LG. Emerging design principles in
biomaterials and scaffolds for tissue engineering. Ann N Y
Acadi Sci 2002; 961:83-95.
[61]Guo XE. Mechanical Properies of Cortical Bone and
Cancellous Tissue. In: Cowin SC, editor. Bone Mechanics
HANDBOOK. Boca Raton: CRC Press LLC, 2001; 10-23.
[62]Gutsche, A. T.; Lo, H.; Zurlo, J.; Yager, J.; Leong,
K. W., Engineering of a sugar-derivatized porous network
for hepatocyte culture. Biomaterials 1996; 17, (3): 387-393.
[63]Hamadouche M, Meunier A, Greenspan DC,
Blanchat C, Zhong JP, LaTorre GP, et al. Long term in
References
165
vivo bioactivity and biodegradability of bulk sol–gel
bioactive glasses. J Biomed Mater Res 2001; 54:560–566.
[64]Hamadouche, M.; Meunier, A.; Greenspan, D.C.;
Blanchat, C.; Zhong, J.P, et al.: Key Engineering
Materials, 2001; 192-195: 593-596.
[65]Hench LL and Polak JM. Third-Generation
Biomedical Materials. Science 2002; 295:1014-1017.
[66]Hench LL and West JK. The sol–gel process. Chem
Rev 1990; 90:33–72.
[67]Hench LL. Bioceramics – from concept to clinic. J
Am Ceram Soc 1991; 74:1487–510.
[68]Hench, L. L., and Polak, J. M., Science 2002;
295:1014.
[69]Hench, L. L., The story of bioglass. J Mater Sci:
Mater Med 2006; 17: 967-978.
[70]Hench, L.L. and West, J.K.: Biological applications
of bioactive glasses, Life Chem Rep 1996, Vol. 13, pp.
187-241.
[71]Hench, L.L., Splinter, R.J., Allen, W.C. and
Greenlee, T.K.: Bonding mechanisms at the interface of
ceramic prosthetic materials. J Biomed Mater Res 1971,
5(6), pp. 117-141.
[72]Hoffman AS. Hydrogels for biomedical applications.
Adv Drug Deliv Rev 2002; 43:3–12.
References
166
[73]Holand W., Vogel W., Naumann K., Gummel J.:
Interface reactions between machinable bioactive glass-
ceramics and bone, J. Biomed. Mater. Res. 1985; 19: 303–
12.
[74]Hollister, S. J., Porous scaffold design for tissue
engineering. Nat Mater 2005; 4, (7): 518-524.
[75]Holmer, P., and Nielsen, P. T., Fracture of ceramic
femoral heads in total hip arthroplasty, J. Arthroplasty,
1993; 8(6):567–571.
[76]Hutmacher, D. W., Scaffold design and fabrication
technologies for engineering tissues - State of the art and
future perspectives. Journal of Biomaterials Science,
Polymer Edition 2001; 12, (1): 107-124.
[77]Hutmacher, D. W.; Sittinger, M, et al.: Scaffold-
based tissue engineering: Rationale for computer-aided
design and solid free-form fabrication systems. Trends in
Biotechnology 2004; 22, (7): 354-362.
[78]Hutmacher, DW.: Scaffolds in tissue engineering
bone and cartilage. Biomate 2000; 21:2529–2543.
[79]Jaakkola T, Rich J, Tirri T, Narhi T, Jokinen M,
Seppala J, et al.: In vitro Ca-P precipitation on
biodegradable thermoplastic composite of poly ([epsilon]-
caprolactone-co-lactide) and bioactive glass (S53P4).
Biomaterials 2004; 25:575–81.
References
167
[80]Jazrawi, L. M., Kummer, F. J., and DiCesare, P. E.,
Alternative bearing surfaces for total joint arthroplasty, J.
Am. Acad.Orthop. Surg. 1998; 6(4): 198–203.
[81]Jell, G., and Stevens, M. M., J. Mater. Sci.: Mater.
Med. 2006; 17: 997.
[82]Jiang G, Evans ME, Jones I, Rudd CD, Scotchford
CA, Walker GS. Preparation of poly (e-
caprolactone)/continuous bioglass fibre composite using
monomer transfer moulding for bone implant.
Biomaterials 2005; 26:2281–2288.
[83]Julian, R. J.: New trends in bioactive scaffolds: The
importance of nanostructure. J Eur Cer Soci. 2009; 29:
1275–1281.
[84]Kaisa Laattalaa, Reeta Huhtinena, Mervi Puskaa, et
al., Bioactive Composite for Keratoprosthesis Skirt,
Journal of the Mechanical Behavior of Biomedical
Materials. 2011; 4: 1700-1708.
[85]Kanellakopoulou, K.: Local treatment of
experimental seudomonas aeruginosa osteomyelitis with a
biodegradable dilactide polymer releasing ciprofloxacin.
Antimicrob Agents Chemother 2008; 52:2335–9.
[86]Karageorgiou, V. and Kaplan, D. Porosity of 3D
biomaterial scaffolds and osteogenesis. Biomat. 2005; 26:
5474–5491.
References
168
[87]Kesavan, S. and Alamelu Bai, S.: Effect of
surfactant on the release of ciprofloxacin from gelatin
microspheres. J. ARS Pharmaceu Ars Pharm, 2010;51 n
1: 1-16.
[88]Kneser U, Schaefer DJ, Polykandriotis E, Horch
RE. Tissue engineering of bone: the reconstructive
surgeon’s point of view. J Cell Mol Med 2006; 10:7-19.
[89]Kohane, D. S.; Langer, R., Polymeric biomaterials in
tissue engineering. Pediatric Research 2008; 63, (5): 487-
491.
[90]Korsmeyer, R. W., R. Gurny, et al. "Mechanisms
of solute release from porous hydrophilic polymers." Int J
Pharm. 1983; 15: 25-35.
[91]Kretlow, J. D., and Mikos, A. G., Tissue Eng2007;
13: 927.
[92]Laczka M., Cholewa-Kowalaska K., Laczka-
Osyczka A., Tworzydlo M., Turyna B., Gel-derived
materials of a CaO–P2O5–SiO2 system modified by boron,
sodium, magnesium, aluminum, and fluorine compounds,
J. Biomed. Mater. Res. 2000; 52: 601–612.
[93]Lanza RP, Langer RS, Vacanti J. Principles of
tissue engineering. 2nd ed. San Diego, CA: Academic
Press, 2000; 65-73.
References
169
[94]Laurencin, C.T. et al: Tissue engineering:
orthopedic applications. Annu. Rev. Biomed. Eng. 1999; 1:
19–46.
[95]Lazzarini, L., Lipsky, BA. and Mader, JT.:
Antibiotic treatment of osteomyelitis: what have we
learned from 30 years of clinical trials. Int J Infect Dis
2005; 9:127–138.
[96]Lefebvre L., Chevalier J., Gremillard L., et al;
Structural transformations of bioactive glass 45S5 with
thermal treatments. Acta Materialia. 2007; 55: 3305–
3313.
[97]Lepretre, S.: Prolonged local antibiotics delivery
from hydroxyapatite functionalised with cyclodextrin
polymers. Biomate 2009; 30:6086–6093.
[98]Li HY, Chang J. pH-compensation effect of
bioactive inorganic fillers on the degradation of PLGA.
Compos Sci Technol 2005; 65:2226–2232.
[99]Li R, Clark AE, Hench LL. An investigation of
bioactive glass powders by sol– gel processing. J Appl
Biomater 1991; 2:231–239.
[100]Liu X and Ma PX. Polymeric scaffolds for bone
tissue engineering. Ann Biomed Eng 2004; 32(3):477-
486.
References
170
[101] Liu, C. Z., and Czernuszka, J. T., Mater. Sci.
Technol. 2007; 23: 379.
[102]Lo, H.; Kadiyala, S.; Guggino, S. E.; Leong, K.
W., Poly (L-lactic acid) foams with cell seeding and
controlled-release capacity. Journal of Biomedical
Materials Research 1996; 30, (4): 475-484.
[103]Luo, P.: Methods of synthesizing hydroxyapatite
powders and bulk materials. United States Patent 1999;
502-530.
[104]Ma J., Chen C.Z., Wang D.G., et al.: Influence of
the sintering temperature on the structural feature and
bioactivity of sol–gel derived SiO2–CaO–P2O5 bioglass,
Ceram.Int.2010;36:1911–1916.
[105]Makinen, TJ., Veiranto, M. and Lankinen, P.: In
vitro and in vivo release of ciprofloxacin from
osteoconductive bone defect filler. J Antimicrob
Chemother 2005; 56:1063–8.
[106]Mamai M, Oudadesse H, Dorbez-Sridi R, et al,
Synthesis and in vitro characterization of melt drieved 47S
CaO–P2O5–SiO2–Na2O bioactive glass 2008; 52 ,3: 121-
129.
[107]Mami M., Lucas-Girot A. Oudadesse H., et al,
Investigation of the surface reactivity of a sol–gel derived
References
171
glass in the ternary system SiO2–CaO–P2O5. Applied
Surface Science, 2008; 254: 7386–7393.
[108]Mami, M., Oudadesse, H. and Doebez-Sridi, R.:
Synthesis and in vitro characterization of melt derived 47S
CaO–P2O5–SiO2–Na2O bioactive glass. Ceram − Silik
.2008; 52 (3):121-129.
[109]Mano JF, Silva GA, Azevedo HS, Malafaya PB,
Sousa RA, Silva SS, et al. Natural origin biodegradable
systems on tissue engineering and regenerative medicine:
present status and some moving trends. J R Soc Interface
2007; 4:999–1030.
[110]Mansur H.S., Or´efice R.L., et al.: J. Polymer
2004; 45: 7193–7202.
[111]Mansur H.S., Or´efice R.L., Vasconcelos W.L.,
Lobato Z.P., Machado L.J., J. Mater. Sci.: Mater. Med.
2005; 16: 333–340.
[112]Mansur, H.S. and Costa, H. S.: Nanostructured
poly(vinyl alcohol)/bioactive glass and poly(vinyl
alcohol)/chitosan/ bioactive glass hybrid scaffolds for
biomedical applications,”. Chem Engi J . 2008; 137, no. 1:
72–83.
[113]Maquet V, Boccaccini AR, Pravata L, et al.: Porous
poly([alpha]-hydroxyacid)/Bioglass® composite scaffolds
References
172
for bone tissue engineering. I: Preparation and in vitro
characterisation. Biomaterials 2004; 25:4185–94.
[114]Maria V.-R., Jos´e Maria G.-C., et al.: Solid State
Chem. 2004; 32: 1– 31.
[115]María, C., Gutiérrez, Zaira Y., Carvajal, G. and
Jobbágy M.: Poly (vinyl alcohol) Scaffolds with Tailored
Morphologies for Drug Delivery and Controlled Release.
Adv Funct Mater. 2007; 17: 3505–3513.
[116]Martinez A., Izuierdo-Barba I., Vallet-Regi M.,
Bioactivity of a CaO–SiO2 binary glasses system, Chem.
Mater. 2000; 12: 3080–88.
[117]Matthew D, Predicting bioactive glass properties
from the molecular chemical composition: Glass transition
temperature. Acta Biomaterialia. 2011; 7 : 2264–2269.
[118]Misra, SK., Nazhat, SN, et al.: Fabrication and
characterization of biodegradable Poly (3-hydroxybutyrate)
composite containing bioglass. Biomacromolecules 2007;
8:2112-2119.
[119]Mooney, D. J., Baldwin, D. F.; Suh, N. P.; et al.,
Novel approach to fabricate porous sponges of poly(d,l-
lactic-co-glycolic acid) without the use of organic solvents.
Biomaterials 1996; 17, (14), 1417-1422.
[120]Mouriño, V. and Boccaccini, A. R. J R Soc Interface
2010; 7:209.
References
173
[121]Nayak, A. K. and Sen, K. K.: Hydroxyapatite-
ciprofloxacin minipellets for bone-implant delivery:
Preparation, characterization, in-vitro drug adsorption and
dissolution studies. Int J Drug Dev Res 2009; 1: 47–59.
[122]Nayak, A. K., Laha, B., et al.: Development of
hydroxyapatite-ciprofloxacin bone-implants using Quality
by design. Acta Pharm. 2011; 61: 25–36.
[123]Nieto A, Areva S, Wilson T, et al , Cell viability in a
wet silica gel. Acta Biomater 2009; 5:3478–3487.
[124]Nishio, K., Neo, M., Akiyama, H., et al, Effects of
apatite and Wollastonite containing glass-ceramic powder
and two types of alumina powder in composites on
osteoblastic differentiation of bone marrow cells, J.
Biomed. Mater. Res. 2001; 55(2):164–176.
[125]O'Brien, F. J., Biomaterials & scaffolds for tissue
engineering. Materials Today 2011; 14, (3), 88-95.
[126]Oonishi, H., et al., In 8th International Symposium on
Ceramics in Medicine. (eds.), Elsevier, Tokyo, 1995; 137.
[127]Oudadesse H, Mamai M, Dorbez-Sridi R, et al,
Study of the Bioactivity of Various Mineral Compositions
of Bioactive Glasses. Bioceramics Development and
Applications 2011; 1: 1-3.
[128]Oudadesse H., Derrien A. C., Lucas-Girot A., et.
al; Calcification mechanism and bony bonding studies of
References
174
calcium carbonate and composite aluminosilicate/calcium
phosphate applied as biomaterials by using radioactivation
methods. J. Radioanalytical and Nuclear Chemistry, 2007;
274, No.2: 421–428.
[129]Oudadesse H., Dietrich E., Gal Y. L., et al. Apatite
forming ability and cytocompatibility of pure and Zn-doped
bioactive glasses. Biomed. Mater. 2011; 6: 35006 - 35015.
[130]Oudadesse, H., Bui, X. V. and Yann L.: Chitosan
Effects on Bioactive Glass for Application as Biocopmosite
Biomaterial. Int J of biolo and biomed engi . 2011; 5: 49-
56.
[131]Oudadesse, H., Mami, M., Doebez-Sridi, R., Pellen,
P., Perez, F., Jeanne S., Chauvel-Lebret D., Mostafa A.
and Cathelineau G., Study of various mineral compositions
and their bioactivity of bioactive glasses. Biocera 2009; 22:
379-382.
[132]Oudadesse, H., Mostafa, A. and Bui X. V.: Physico-
chemical assessment of biomimetic nano-
hydroxyapatite/polymer matrix for use in bony surgery. Int
J of biolo and Biomed Engi. 2011; 5: 103-110.
[133]Oyanagi Y and Matsumoto A, J Coll Sci 1962;
17:426.
[134]Park J.B: Biomaterials Science and Engineering.
New York, Plenum Publishing Corp, 1984; 15-23.
References
175
[135]Peniche C, Argu¨elles-Monal W, Peniche H, Acosta
N. Chitosan: an attractive biocompatible polymer for
microencapsulation. Macromol Biosci 2003; 3:511–520.
[136]Peppas NA, Hilt JZ, Khademhosseini A, Langer R.
Hydrogels in biology and medicine: from molecular
principles to bionanotechnology. Adv Mater 2006;
18(11):1345–1360.
[137]Pereira A.P.V., Wander L.V., et al.: Novel
multicomponent silicate–poly (vinyl alcohol) hybrids with
controlled reactivity, J. Non- Cryst. Solids 2000; 273: 180–
185.
[138]Pereira M., Clarck A.E., et al.: Journal of
Biomedical Materials Research, 1994; 28: 693-698.
[139]Pereira M., Clarck A.E., et al.: Journal of Materials
Synthesis and Processing, 1994; 2, n. 3: 189-195.
[140]Pereira M., Jones J.R., et al.: J. Adv. Appl. Ceram.
2005; 104 (1) : 35–42.
[141]Peiro, A., Pardo, J., Navarrete, R., et al.: Fracture
of the ceramic head in total hip arthroplasty:Report of two
cases, J. Arthroplasty, 1991;6(4):371–374.
[142]Peter M, Binulala N.S., Nair S.V.et al., Novel
biodegradable chitosan–gelatin/nano-bioactive glass
ceramic composite scaffolds for alveolar bone tissue
engineering. Chemical Engineering Journal, 2010; 158:
References
176
353–361.
[143]Peter M, Binulal N.S., Soumya S., et al.,
Nanocomposite scaffolds of bioactive glass ceramic
nanoparticles disseminated chitosan matrix for tissue
engineering applications.Carbohydrate Polymers 2010 ; 79:
284–289.
[144]Pham, D. T.; Gault, R. S., A comparison of rapid
prototyping technologies. International Journal of Machine
Tools and Manufacture 1998; 38, (10-11): 1257-1287.
[145]Platel, R. H.; Hodgson, Place, E. S, et al., Synthetic
polymer scaffolds for tissue engineering. Chemical Society
Reviews 2009; 38, (4), 1139-1151.
[146]Pope Edgard JA. Bioartificial organs I: silica gel
encapsulated pancreatic islets for the treatment of diabetes
mellitus. J Sol-Gel Sci Technol 1997; 8:635–639.
[147]Poursamar S. A, Azami M, et al.: Controllable
synthesis and characterization of porous polyvinyl
alcohol/hydroxyapatite nanocomposite scaffolds via an in
situ colloidal technique. Colloids and Surfaces B:
Biointerfaces 2011; 84: 310–316.
[148]Puga A M., Rey-Rico A, Magariños B, Alvarez-
Lorenzo C, et al.: Hot melt poly-e-
caprolactone/poloxamine implantable matrices for
sustained delivery of ciprofloxacin. Acta Biomaterialia
References
177
2012; 8:1507–1518.
[149]Ra´mila A., Balas F., Vallet-Regı M. ´, Synthesis
routes for bioactive sol–gel glasses: alkoxides versus
nitrates, Chem. Mater. 2002; 14: 542–548.
[150]Ramakrishna S, Mayer J, Wintermantel E, Leong
KW. Biomedical applications of polymer-composite
materials: a review. Compos Sci Technol 2001; 61:1189–
1224.
[151]Reilly DT, Burstein AH, Frankel VH. The elastic
modulus for bone. J Biomech 1974; 7(3):271-275.
[152]Rho JY, Kuhn-Spearing L, Zioupos P. Mechanical
properties and the hierarchical structure of bone. Med. Eng.
Phys. 1998; 20(2):92-102.
[153]Robert J K and Ryan K R, Effects of hydroxyapatite
reinforcement on the architecture and mechanical
properties of freeze-dried collagen scaffolds. J. The
mechanical behavior of biomedical materials.2012; 7: 41-
49.
[154]Rodrı´guez-Tenreiro, C., Alvarez-Lorenzo, C.,
Concheiro, A., et al.: Characterization of cyclodextrin-
carbopol interactions by DSC and FTIR. J Therm Anal
Calorim. 2004; 77:403–411.
[155]Rom´an J., PadillS. a, et al.:, J. Chem. Mater. 2003;
15:798–806.
References
178
[156]Ross JM. Cell-Extracellular Matrix Interactions. In:
Patrick CW, Mikos AG, Mc.Intre L, editors. Frontiers in
Tissue Engineering. Oxford: Elsevier Science Ltd., 1998:
15-27.
[157]Sahoo, S., Charkaborti, Ch. K. et al.: Qualitative
analysis of a ciprofloxacin / HPMC mucoadhesive
suspension. Int J Pharma and Bio Sci , 2012; 3: 558-76.
[158]Salinas A.J., Roman J., Vallet-Regi M., Oliveira
J.M., Correia R.N., Fernandes M.H., In vitro bioactivity
of glass and glass-ceramics of the 3CaOP2O5–CaOSiO2–
CaOMgO2SiO2 system, Biomaterials 2000; 21 :251–257.
[159]Saltzman WM. Tissue Engineering: principles for the
design of replacement organs and tissues. 1st ed. Oxford:
Oxford University Press, 2004; 253-267.
[160]Saravanapavan, P and Hench, L.L. Key
Engineering Materials, 2001; 192-195: 609-612.
[161]Schmidt CE, Baier JM. Acellular vascular tissues:
natural biomaterials for tissue repair and tissue engineering.
Biomaterials 2000; 21(22):2215-2231.
[162]Shahrabi S, Hesaraki S, Moemeni S et al.:
Structural discrepancies and invitro nanoapatite formation
ability of sol–gel derived glasses doped with different bone
stimulator ions, Ceramics Internationa. 2011; 37: 2737-
2746.
References
179
[163]Shi C, Zhu Y, Ran X, Wang M, et al.: Therapeutic
potential of chitosan and its derivatives in regenerative
medicine. J Surg Res 2006; 133:185–92.
[164]Shin, S.-H. and Kim, H.-I.: Contribution of
hydrogen bond and coupling reaction improvement in
compatibility of organic polymer/silica nanocomposites,”.
Ind. & Engi Chem Res., 2001; 7: 147–152.
[165]Shoichet, M. S., Polymer scaffolds for biomaterials
applications. Macromolecules 2010; 43, (2): 581-591.
[166]Utsel S, Carlmark A, Pettersson T, et al.: Synthesis,
adsorption and adhesive properties of a cationic
amphiphilic block copolymer for use as compatibilizer in
composites. European Polymer Journal 2012; 48: 1195–
1204.
[167]Stahl A, Wenger A, Weber H, et al.: Bi-directional
cell contact-dependent regulation of gene expression
between endothelial cells and osteoblasts in a three-
dimensional spheroidal coculture model. Biochem Biophys
Res Commun 2004; 322:684-692.
[168]Sokolsky-Papkov M., K. Agashi, A,Olaye, K
Shakesheff, et al.: J. Polymer carriers for drug delivery in
tissue engineering. Adv Drug Deliv Rev 2007; 59:187–206.
[169]Sultana, N. and Wang, M., Fabrication of
HA/PHBV composite scaffolds through the emulsion
References
180
freezing/freeze-drying process and characterisation of the
scaffolds. Journal of Materials Science: Materials in
Medicine 2008; 19, (7): 2555-2561.
[170]Suchanek W. and Yoshimura M. Processing and
properties of hydroxyapatite-based biomaterials for use as
hard tissue replacement implants. J. Materials
Research 1998; 13 , 94-117.
[171]Kulkarni S and Wunder S L.: Moduli of ordered
polymer composites prepared by colloidal crystallization of
nano- and micro-SiO2 spheres in crosslinked methacrylate
resins. Mechanics of Materials. 2011; 43:643-653.
[172]Sunitha , A. and Kumar, S.: Study on the effect of
polymers on the release rate of drug from ciprofloxacin
hydrochloride microspheres .J Pharmaceu Cosmetolo
2010;1(1): 1-8.
[173]Sunitha, A. and Kumar, S.: Study on Effect of
Solvents & Nonsolvents on Microspheres of Ciprofloxacin:
Coacervation Phase Separation. J Adv Sci Res, 2010; 1(2):
24-33.
[174]Superb, K. M., Dirk, M. et al.: Comparison of
nanoscale and microscale bioactive glass on the properties
of P (3HB)/Bioglass composites. Biomat. 2008; 29: 1750-
1761.
[175]Swapnika R., Upasana M., Sridhar R., Russell R,
References
181
et al., Assessing viscoelastic properties of chitosan
scaffolds and validation with cyclical tests. Acta
Biomaterialia 2012; 8: 1566–1575.
[176]Takasu A, Aoki K, Tsucyia M and Okada M, J Appl
Polym Sci 1999; 73:1171.
[177]Tang Y.-F., Du Y.-M., Hu X.-W., Shi X.-W.,
Kennedy J.F., Rheological characterization of a novel
thermosensitive chitosan/poly (vinyl alcohol) blend
hydrogel, Carbohydr. Polym. 2007; 67 (4): 491–499.
[178]Tao, W., Mahir, T., et al.: Selected properties of pH-
sensitive, biodegradable chitosan–poly (vinyl alcohol)
hydrogel, Polym. Int. 2004; 53: 911–918.
[179]Tay, B. K., Patel, V. V., and Bradford, D. S.,
Calcium sulfate- and calcium phosphate-based bone
substitutes: Mimicry of the mineral phase of bone, Orthop.
Clin. North Am. 1999; 30(4):615–623.
[180]Thakre, Y. M. and Choudhary, M. D.: Synthesis,
characterization and evaluation of derivative of
Ciprofloxacin (1-cyclopropyl-6-fluoro-4-oxo-7-[4-(phenyl
carbonyl) piperazin-1-yl]-1, 4-dihydroquinoline-3-
carboxylic acid) and their complexes. J Chem Pharm Res,
2011; 3(5):390-8.
[181]Tharanathan RN, Kittur FS. Chitin – the
undisputed biomolecule of great potential. Crit Rev Food
References
182
Sci 2003; 43(1):61–87.
[182]Tsigkou, O., et al., J. Biomed. Mater. Res. A 2007; 80:
837.
[183]Unnithan, A. R., Barakat, N. A.M. and Tirupathi
Pichiah, P.B., Wound-dressing materials with antibacterial
activity from electrospun polyurethane–dextran nanofiber
mats containing ciprofloxacin HCl. Carbohyd Poly,
2012;90 : 1786–1793.
[184]Vacanti CA, Bonassar LJ, Vacanti JP. Structural
Tissue Engineering. In: Lanza RP, Langer R, Vacanti
JP, editors. Principles of Tissue Engineering. San Diego:
Academic Press, 2000; 671-682.
[185]Vacanti CA. History of Tissue Engineering and a
Glimpse into Its Future. Tissue Engineering 2006;
12(5):1137‐1142.
[186]Vallet-Reg M, Balas F, Arcos D. Mesoporous
materials for drug delivery. Angew Chem Int Ed 2007;
46:7548–58.
[187]Vallet-Reg M. Revisiting ceramics for medical
applications. Dalton Trans 2006; 44:5211–5220.
[188]Vallet-Regı M. ´, Romero A.M., Ragel C.V.,
LeGeros R.Z., XRD, SEM EDX, and FTIR studies of in
vitro growth of an apatite-like layer on sol–gel glasses, J.
Biomed. Mater. Res. Part A 1999; 44: 416–421.
References
183
[189]Wainwright SA, Biggs WD, Currey JD, Gosline
JM. Mechanical Design in Organisms. First Edition ed.
Princeton, New Jersey: Princeton University Press, 1976;
33-37.
[190]Wanderley dos S R, Marivalda M P et al.: Analysis
of bioactive glasses obtained by Sol-Gel Processing for
Radioactive Implants. Materials research. 2003; 6:123-127.
[191]Wang M. Developing bioactive composite materials
for tissue replacement. Biomaterials 2003; 24:2133–2151.
[192]Wang X, Bank RA, TeKoppele JM, et al.: The role
of collagen in determining bone mechanical properties. J.
Orthop. Res. 2001; 19(6):1021-1026.
[193]Wang, Q., Zedong, Du Y. et al.: Controlled release of
ciprofloxacin hydrochloride from chitosan/polyethylene
glycol blend films. Carbohyd. Polym.2007; 69:336–343.
[194]Wang, T., Turhan, M. et al.: “Selected properties of
pH-sensitive, biodegradable chitosan-poly (vinyl alcohol)
hydrogel,” Poly Int., 2004; 53, no. 7: 911–918.
[195]Webster T.J., Ergun C., Doremus R.H., et. al;
Enhanced osteoclast-like cell functions on nanophase
ceramics, Biomaterials, 2001;22 :1327.
[196]Webster T.J., Ergun C., Doremus R.H., Siegel
R.W., Bizios R., J. Biomed.Specific proteins mediate
enhanced osteoblast adhesion on nanophase ceramics.
References
184
Mater. Res. 2000; 51:475.
[197]Webster T.J., Siegel R.W., Bizios R.: Osteoblast
adhesion on nanophase ceramics. Biomaterials. 1999;
20:1221-1227.
[198]Williams D.F., Tissue-Biomaterial Interactions, J.
Mater. Sci., 1987;22:3421–45.
[199]Wong Sh., Baji, A. et al.: Effect of specimen
thickness on fracture toughness and adhesive properties of
hydroxyapatite-filled polycaprolactone. Compos: Part A,
2008; 39: 579–587.
[200]Wu, F., Liu, Ch., et al.: Fabrication and properties of
porous scaffold of magnesium phosphate/polycaprolactone
biocomposite for bone tissue engineering. App Surf Sci,
2012; 258:7589–7595.
[201]Wu Q, Zhang X, et al.: Fabrication and
characterization of porous HA/b-TCP scaffolds
strengthened with micro-ribs structure. Materials Letters.
2013 ; 92 : 274–277.
[202]Xin L, Mohamed N R, et al.: Oriented bioactive
glass (13-93) scaffolds with controllable pore size by
unidirectional freezing of camphene-based suspensions:
Microstructure and mechanical response. Acta
Biomaterialia 2011; 7:406–416.
[203]Xynos ID, Edgar AJ, et al.: Ionic products of
References
185
bioactive glass dissolution increase proliferation of human
osteoblasts and induce insulin-like growth factor II mRNA
expression and protein synthesis. Biochem Biophys Res
Commun 2000; 276:461–465.
[204]Xynos ID, Hukkanen MVJ, Batten JJ, et al.:
Bioglass ®45S5 stimulates osteoblast turnover and
enhances bone formation in vitro: implications and
applications for bone tissue engineering. Calcif Tissue Int
2000; 67:321–329.
[205]Yamaoka T., Tabata Y., et al.: Comparison of body
distribution of poly (vinyl alcohol) with other water-soluble
polymers after intravenous administration, J. Pharmaceut.
Pharmacol. 1995; 47:479–486.
[206]Yaszemski MJ, Payne RG, Hayes WC, et al.:
Evolution of bone transplantation: molecular, cellular and
tissue strategies to engineering human bone. Biomaterials
1996; 17:175-185.
[207]Yazdanpanah, A., Reza, K. et al.: Enhancement of
the fracture toughness in bioactive glass-based
nanocomposites with nanocrystalline forsterite as advanced
biomaterials for bone tissue engineering applications.
Ceram Inter 2012; 38:5007-5014.
[208]Yildirim, M. S., Hasanreisoglu, U., et al.:. J Oral
Rehabil. 2005; 32:518.
References
186
[209]Zarzycki J. Past and present of sol–gel science and
technology. J Sol-Gel Sci Technol 1997; 8:1–6.
[210]Zhang, F., Chuanglong H. et al.: Fabrication of
gelatin–hyaluronic acid hybrid scaffolds with tunable
porous structures for soft tissue engineering. Int J Biolo
Macromolec, 2011; 48: 474–481.
[211]Zhang, P., Hong, Z., Yu, T., et al.: In vivo
mineralization and osteogenesis of nanocomposite scaffold
of poly (lactide-co-glycolide) and hydroxyapatite surface-
grafted with poly(l-lactide). Biomaterials 2009; 30, (1): 58-
70.
[212]Zhang Q, Jiang Y, Zhang Y, et al. Effect of porosity
on long-term degradation of poly (ε-caprolactone) scaffolds
and their cellular response. J. Polymer Degradation and
Stability. 2013;98 : 209-218.
[213]Zhengmao Li, Xiaofeng Chen, Cai Lin, et al. The in
vitro bioactive of sol–gel bioactive glass powders with
three-dimensional lamellar structure. Advanced Powder
Technology, 2012; 23:13–15.
[214]Zhong J and Greenspan DC. Processing and
properties of sol–gel bioactive glasses. J Biomed Mater Res
2000; 53:694–701.
[215]Zijderveld, S. A., et al. Int. J. Oral Maxillofac.
Implants 2005; 20: 431-432.
Résumé:
Ce travail de thèse est basé sur la préparation de verres bioactifs (BG) par différents procédés tels que la fusion, la voie sol-gel et le scaffolds. Le Poly Vinyl Alcohol (PVA) a été associé aux verres élaborés dans un système quaternaire (BG) 46S6 par les procédés cités (fusion, sol-gel et sacffolds). Différents paramètres intervenant dans les synthèses des verres bioactifs ont été étudiés, nous citons à titre d’exemple : la température, le pH, la taille des particules, le rapport Polymère / verres, la microstructure, la porosité et la biodégradation. Les caractéristiques thermiques des verres élaborés ont été également déterminées après chaque synthèse par analyse thermique différentielle (DSC/TG, DTA/TG). Ainsi, la température de fusion, la température de transition vitreuse et la température de cristallisation ont été élucidées. Ces caractéristiques thermiques changent lorsque la composition chimique du verre est modifiée. A ce titre, les compositions chimiques ont été étudiées par Fluorescnece (XRF) et Inductively Coupled Plasma-Opticale Emission Spectroscopy (ICP-OES) après chaque synthèse pour s’assurer de la pureté des verres bioactifs élaborés et destinés à des applications médicales. Plusieurs techniques physico chimiques d’analyses (DRX, MEB, MET, FT-IR, XRF, ICP-OES) ont été mises en œuvre pour déterminer les propriétés physico chimiques de nos verres bioactifs avant et après expérimentations « in vitro ». Le nano composite Polymère-Verres scaffolds que nousavons obtenu présente des particules de tailles comprises entre 40 et 61 nm et une porosité d’environ 85%. La biodégradation des verres scaffolds décroît lorsque la teneur en verre scaffolds dans le nano composite croît. Les expérimentations « in vitro » montrent qu’après immersion de ces nano composites dans un liquide physiologique synthétique (SBF), une couche d’apatite (phosphate de calcium) se forme à leur surface. L’épaisseur de la couche formée dépend clairement de la taille des particules et du rapport polymère / verre scaffolds.
Mots clés : Biomatériaux, verres bioactifs, nanocomposite, fusion, sol-gel, scaffolds, caractérisation physico-chimique, réactivité chimique, bioactivité, biodégradation Summary:
The aim of the present work is the preparation of Bioactive Glass (BG) 46S6 by different techniques. Fabrication of composite scaffolds by using of Poly Vinyl Alcohol (PVA) and quaternary BG (two methods melting and sol-gel) with different ratios to the prepared scaffolds was carried out. Different factor affecting the final properties of the prepared composite scaffolds were investigated in this study, such as; temperature of treatment, BG particle size, polymer/glass ratio, microstructure, porosity, biodegradation, bioactivity, and drug release. The thermal behavior of the prepared bioactive glass by sol-gel and melting techniques were identified using Differential Scanning Calorimetric/Thermo Gravimetric (DSC/TG) or Differential Thermal Analysis/Thermo Gravimetric (DTA /TG). The elemental composition of the prepared bioactive glasses was determined by X-rays Fluorescence (XRF) to confirm that the prepared bioactive glasses have the same elemental compositions and high purity for biomedical applications. The particle size of the prepared bioactive glass was determined by Transmission Electron Microscopic (TEM). Nano-bioactive glass could be obtained by modified sol-gel and the obtained particle size ranged between 40 to 61 nm. The prepared bioactive glass by both applied methods has the same amorphous phase and all identified groups as well as. The porous scaffold has 85% porosity with a slight decrease by increasing the glass contents. The degradation rate decreased by increasing of glass content in the prepared scaffolds. The bioactivity of the prepared composite scaffolds was evaluated by XRD, FTIR, SEM coupled with EDX and Inductively Coupled Plasma-Optical Emission Spectroscopic (ICP-OES). It has been observed that after soaking in Simulated Body Fluid (SBF), there was an apatite layer formed on the surface of the prepared samples with different thickness depending on the glass particle size and polymer/glass ratio. Key words: Biomaterials, bioactive glass, nanocomposite, melting, sol-gel, scaffolds, physic- chemical characterization, chemical reactivity, bioactivity, biodegradation.