mention : chimie · vii 51 rp rapid prototyping 52 sbf simulated body fluid 53 sg-b sol-gel...

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


116

48 Fig. 4.16.b

SBF P ions concentrations after soaking of the prepared samples


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


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).


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


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


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).


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


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


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


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


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


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.


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


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


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


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,


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


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


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


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.


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


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


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)].


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)


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


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.


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


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


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)].


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


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


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


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


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:


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


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.


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


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


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


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


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


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.


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


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


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.


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.


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.


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,


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


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,


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


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


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)].


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.


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


57


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


58


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


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


60

Part 1

Preparation of 46S6 bioactive glass by

-Melting technique and

-Sol-gel method


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.


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.


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.


64

Part 2

Preparation of biocomposites using

polymer technique


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.


66

Part 3

Preparation of the composite scaffolds

by freeze drying technique


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 %)


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.


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


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.


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


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


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


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


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


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


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)


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.


80

Part 1

Results and discussions of 46S6

bioactive glass by

-Melting technique and

-Sol-gel method


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


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.


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).


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


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)


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.


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


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


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.


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


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


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


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


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


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.


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


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.


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)


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)


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


101

Part 2

Results and discussions of the prepared

biocomposites using polymer technique


102


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


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


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


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.


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.


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


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


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


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.


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


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


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


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


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.


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


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


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


119

Part 3

Results and discussion of the prepared

composite scaffolds by freeze drying

technique


120


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


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).


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


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.


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)


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)


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)


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).


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


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


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)


131

c) XRD of 1PVA:1MB before

and after immersion in SBF.

d) XRD of 1PVA:1SG-B before


e) XRD of 2PVA:1MB before


f) XRD of 2PVA:1SG-B before


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


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


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.


134

a) FTIR of 1PVA:2MB before


b) FTIR of 1PVA:2SG-B before


c) FTIR of 1PVA:1MB before


d) FTIR of 1PVA:1SG-B before


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


135

e) FTIR of 2PVA:1MB before


f) FTIR of 2PVA:1SG-B before


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


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.


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


138

b) Ca ions concentrations after soaking of PVA/SG-B Scaffolds in SBF


c) P ions concentrations after soaking of PVA/MB Scaffolds in SBF for


d) P ions concentrations after soaking of PVA/SG-B Scaffolds in SBF


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


139

e) Si ions concentrations after soaking of PVA/MB Scaffolds in SBF for


f) Si ions concentrations after soaking of PVA/SG-B Scaffolds in SBF


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


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).


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


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).


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


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


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.


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.


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


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)


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


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%


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%


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

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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.

mention : chimie · vii 51 rp rapid prototyping 52 sbf simulated body fluid 53 sg-b sol-gel...

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