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BIOMIMETIC SURFACE MODIFICATION OF DENTAL
IMPLANT FOR ENHANCED OSSEOINTEGRATION
RAJESWARI RAVICHANDRAN
(B.Tech, ANNA University)
A THESIS SUBMITTED
FOR THE DEGREE OF MASTER OF ENGINEERING
DIVISION OF BIOENGINEERING
NATIONAL UNIVERISTY OF SINGAPORE
2009
1
ACKNOWLEDGEMENT
I would like to express my sincere appreciation to those who have helped and contributed
to this thesis. I would like to thank Professor Michael Raghunath who has shown faith in
me and given me tremendous encouragement throughout my tenure. I would like to
express my sincere thanks to Professor Seeram Ramakrishna for his excellent supervision
and guidance throughout this project.
I would like to express my heartfelt gratitude to Dr Clarisse Ng and Dr Susan Liao, who
have provided unmatched guidance and support, throughout this project. I would also like
to thank Professor Casey Chan, Dr. Damian Pliza and Dr. Venugopal for giving me
invaluable advice, discussion, and suggestions. I would also like to thank all Prof
Seeram’s lab members for their assistance in the completion of this project.
I would like to thank the Division of Bioengineering and the Faculty of Dentistry for their
constant support.
Last but not the least I would like to thank my parents for their profound love and
support.
2
TABLE OF CONTENTS
ACKNOWLEDGEMENTS
TABLE OF CONTENTS
LIST OF FIGURES
LIST OF TABLES
LIST OF APPENDICES
LIST OF ABBREVIATIONS
Chapter 1: Introduction
1.1 Background
1
1.2 Clinical problems associated with osseointegration
2
1.3 Hypothesis and objectives
4
Chapter 2: Literature Review
2.1 Introduction
5
2.2 Surface modification techniques
6
2.2.1 Modification of scaffolds using surface adhesive molecules
2.2.2 Cell – substrate interaction
9
13
3
2.3 Tissue Engineering
2.3.1 Introduction
14
2.3.2 Nanofiber fabrication by electrospinning
14
2.3.3 Modifications of the electrospun nanofibers
19
2.3.4 Potential application of Mesenchymal stem cells for osseointegration
20
Chapter 3: Biomimetic surface modification of dental implant by advanced
electrospinning
3.1 Introduction
21
3.2 Materials and Methods
22
3.2.1 Mechanical Polishing/ etching
22
3.2.2 Pretreatment of Ti
22
3.2.3 Electrospinning of PLGA and PLGA/Collagen nanofibers on the Ti discs
23
3.2.4 Biomineralization using Calcium-Phosphate dipping method
25
3.2.5 Cell adhesion study
26
3.2.6 Surface characterization analysis
27
3.2.7 Surface roughness analysis
28
4
3.2.8 Fourier transform infrared spectroscopy (FT-IR) and X-ray photoelectron
spectroscopy (XPS)
28
3.2.9 Water contact angle measurement
29
3.2.10 Statistical analysis
29
3.3 Results and Discussion
29
3.3.1 Surface characterization analysis
29
3.3.2 Surface Roughness analysis
34
3.3.3 Fourier transform infrared spectroscopy (FT-IR)
35
3.3.4 Water contact angle measurement
38
3.3.5 X-ray photoelectron spectroscopy (XPS)
39
3.3.6 Cell culture analysis
40
3.4 Conclusion
46
Chapter 4: Mesenchymal stem cells proliferation and differentiation studies on the
modified implant surfaces
4.1 Introduction
47
4.2 Material and methods
48
4.2.1 Mesenchymal stem cells culture
48
4.2.2 Cell Morphology study
49
5
4.2.3 Cell Proliferation study
49
4.2.4 Alkaline phosphatase activity
50
4.2.5 Cell mineralization study
50
4.2.6 Statistical analysis
51
4.3 Results and Discussion
51
4.3.1 Cell Morphology study
51
4.3.2 Cell Proliferation study
56
4.3.3 Alkaline phosphatase activity
58
4.3.4 Cell mineralization study
61
4.4 Conclusion
70
Chapter 5: Conclusions and Recommendations
5.1 Conclusions
72
5.2 Recommendations
73
6
LIST OF FIGURES
Figure 1.1 A model Ti dental implant
3
Figure 2.1 Scaffold architecture affects cell binding and spreading
6
Figure 2.2 Schematic diagram of electrospinning set-up
18
Figure 3.1 Electrospinning set up
24
Figure 3.2 Electric field pattern a) 18kV at the needle tip and 10kV at the
ring electrode, b) 18kV at the needle tip and 14kV at the ring electrode
24
Figure 3.3 Biomineralization procedure
26
Figure 3.4 SEM images of a) untreated Ti, b) Ti after surface modification c) Ti
34
coated with PLGA nanofibers at 1000X magnification d) Ti coated with PLGA/Collagen
nanofibers at 1000X magnification e) Ti coated with PLGA nanofibers at 5000X magnification f)
Ti coated with PLGA/Collagen nanofibers at 5000X magnification g) Ti coated with
functionalized PLGA/Collagen nanofibers h) Ti coated with functionalized PLGA/Collagen
nanofibers
Figure 3.5 AFM image of pretreated Ti showing the surface roughness
35
Figure 3.6 FTIR results for a) cpTi treated and untreated, b) Ti6Al4V alloy treated
and untreated, c) PLGA and PLGA/Collagen nanofibers coated over the Ti surface.
37
Figure 3.7 XPS results showing the Ti2p peaks in the treated samples
39
7
Figure 3.8A Adhesion of hMSCs on the a) untreated cpTi implants, b) cpTi
42
implant coated with PLGA nanofibers, c) cpTi implant coated with PLGA/Collagen nanofibers,
d) cpTi implant coated with PLGA/HA, e) cpTi implant coated with PLGA/Collagen/HA
nanofibers at 500x
Figure 3.8B Adhesion of hMSCs on the a) untreated Ti6Al4V implants,
43
b) Ti6Al4V implant coated with PLGA nanofibers, c) Ti6Al4V implant coated with
PLGA/Collagen nanofibers, d) Ti6Al4V implant coated with PLGA/HA, e) Ti6Al4V implant
coated with PLGA/Collagen/HA nanofibers at 500x
Figure 3.9 Percentage attachment efficiency of hMSCs on cpTi and Ti6Al4V alloy
45
Figure 4.1 SEM images of the hMSC morphology on day 7 on a) untreated Ti, b)
53
Treated Ti coated with PLGA nanofibers, c) Treated Ti coated with PLGA/Collagen nanofibers,
d) Treated Ti coated with functionalized PLGA nanofibers, e) Treated Ti coated with
functionalized PLGA/Collagen nanofibers.
Figure 4.2 SEM images of the hMSC morphology on day 14 on a) untreated Ti, b)
54
Treated Ti coated with PLGA nanofibers, c) Treated Ti coated with PLGA/Collagen nanofibers,
d) Treated Ti coated with functionalized PLGA nanofibers, e) Treated Ti coated with
functionalized PLGA/Collagen nanofibers.
Figure 4.3 SEM images of the hMSC morphology on day 21 on a) untreated Ti, b)
55
Treated Ti coated with PLGA nanofibers, c) Treated Ti coated with PLGA/Collagen nanofibers,
d) Treated Ti coated with functionalized PLGA nanofibers, e) Treated Ti coated with
functionalized PLGA/Collagen nanofibers.
Figure 4.4 MTS assay for hMSC cells proliferation on a) cpTi based
57
scaffolds untreated, coated with PLGA nanofibers, coated with PLGA/Collagen nanofibers,
coated with PLGA/HA and coated with PLGA/Collagen/HA nanofibers b) Ti-6Al-4V based
scaffolds - untreated, coated with PLGA nanofibers, coated with PLGA/Collagen nanofibers,
coated with PLGA/HA and coated with PLGA/Collagen/HA nanofibers for day 7, 14 and 21. *
represents p≤ 0.05 statistical difference. Control refers to the Tissue Culture Plate (TCP); TiK
refers to Ti6Al4V alloy.
Figure 4.5 ALP activity for hMSC cells on a) cpTi based
60
scaffolds - untreated, coated with PLGA nanofibers, coated with PLGA/Collagen nanofibers,
8
coated with PLGA/HA and coated with PLGA/Collagen/HA nanofibers b) Ti-6Al-4V based
scaffolds - untreated, coated with PLGA nanofibers, coated with PLGA/Collagen nanofibers,
coated with PLGA/HA and coated with PLGA/Collagen/HA nanofibers for day 7, 14 and 21. *
represents p≤ 0.05 statistical difference. Control refers to the Tissue Culture Plate (TCP); TiK
refers to Ti6Al4V alloy.
Figure 4.6 Quantitative data for Alizarin red staining on hMSC cells on a) cp Ti
63
scaffolds b) Ti-6Al-4V scaffolds for days 7, 14 and 21. * represents p≤ 0.05 statistical difference
Figure 4.7A Optical image of the ARS stained hMSCs on the cp Ti scaffolds
64 on day 7 a) untreated Ti, b) Treated Ti coated with PLGA nanofibers, c) Treated Ti coated
with PLGA/Collagen nanofibers, d) Treated Ti coated with functionalized PLGA nanofibers, e)
Treated Ti coated with functionalized PLGA/Collagen nanofibers.
Figure 4.7B Optical image of the ARS stained hMSCs on the cpTi
65
scaffolds on day 14 a) untreated cpTi, b) Treated cpTi coated with PLGA nanofibers, c) Treated
cpTi coated with PLGA/Collagen nanofibers, d) Treated cpTi coated with functionalized PLGA
nanofibers, e) Treated cpTi coated with functionalized PLGA/Collagen nanofibers.
Figure 4.7C Optical image of the ARS stained hMSCs on the cpTi scaffolds
66
on day 21 a) untreated cpTi, b) Treated cpTi coated with PLGA nanofibers, c) Treated cpTi
coated with PLGA/Collagen nanofibers, d) Treated cpTi coated with functionalized PLGA
nanofibers, e) Treated cpTi coated with functionalized PLGA/Collagen nanofibers.
Figure 4.8A Optical image of the ARS stained hMSCs on the Ti6Al4V
67
scaffolds on day 7 a) untreated Ti6Al4V, b) Treated Ti6Al4V coated with PLGA nanofibers, c)
Treated Ti6Al4V coated with PLGA/Collagen nanofibers, d) Treated Ti6Al4V coated with
functionalized PLGA nanofibers, e) Treated
Ti6Al4V coated with functionalized
PLGA/Collagen nanofibers.
Figure 4.8B Optical image of the ARS stained hMSCs on the Ti6Al4V
68
scaffolds on day 14 a) untreated Ti6Al4V, b) Treated Ti6Al4V coated with PLGA nanofibers, c)
Treated Ti6Al4V coated with PLGA/Collagen nanofibers, d) Treated Ti6Al4V coated with
functionalized PLGA nanofibers, e) Treated Ti6Al4V coated with functionalized PLGA/Collagen
nanofibers.
Figure 4.8C Optical image of the ARS stained hMSCs on the Ti6Al4V
69
scaffolds on day 21 a) untreated Ti6Al4V, b) Treated Ti6Al4V coated with PLGA nanofibers, c)
Treated Ti6Al4V coated with PLGA/Collagen nanofibers, d) Treated Ti6Al4V coated with
9
functionalized PLGA nanofibers, e) Treated Ti6Al4V coated with functionalized PLGA/Collagen
nanofibers.
10
LIST OF TABLES
Table 2.1 Different types of implant surface modifications and their surface
7
roughness and contact angle.
Table 2.2: Various fabrication techniques along with their advantages and
12
disadvantages
Table 2.3 commonly used polymers and their properties
Table 2.4. Factors that affect the electrospinning process and fiber morphology
Table 3.1 Optimization of electrospinning parameters by varying the time and
15
16
32
concentration for PLGA nanofibers
Table 3.2 Optimization of electrospinning parameters by varying the time and
32
concentration for PLGA/Collagen nanofibers
Table 3.3 Water contact angle measurements for treated and untreated cpTi and
38
Ti6Al4V alloy
Table 3.4 Water contact angle measurements for PLGA and PLGA/Collagen
39
nanofibers
Table 3.5: Average number of cells adhered to the Ti samples
44
11
LIST OF APPENDICES
Appendix A: Optical image of hMSC morphology cultured on TCP
Appendix B: FESEM EDX results showing cell mineralization.
12
LIST OF ABBREVIATIONS
AFM
atomic force microscopy
ECM
extracellular matrix
HFP
1,1,1,3,3,3-hexafluoro-2-propanol
kDa
unit of 1000 Dalton
PBS
phosphate buffered saline
PLGA poly(lactic acid)-co-poly(glycolic acid)
SEM
scanning electron microscopy
w
weight
v
volume
XPS
x-ray photoelectron spectrometry
avg
average
13
SUMMARY
The introduction of dental implants has changed the way dentists approach the
replacement of missing teeth. The clinical success of dental implants is related to their
osseointegration, which is a property virtually unique to titanium and has enhanced the
science of joint replacement techniques. Generally, the time between implant placement
and implant loading ranged from 3 months in the mandible to 6 months in the maxilla, for
machined surfaces.
However, the trend towards a shorter healing time is largely driven by consumer demands
as many patients are unhappy waiting long periods of time for their prosthesis. In order to
achieve rapid osseointegration, it is necessary that the implant surface has an improved
capture ratio which will provide a critical number of mesenchymal stem cells (MSCs)
necessary for successful bone integration [63].
We have proven that the fabrication of a nanofibrous scaffold offers the possibility to
optimize stem cell capture as well as cell adhesion and proliferation, as the nanofibers
mimic the ECM matrix. It is our hypothesis that this improved capture ratio will provide
a critical number of MSCs necessary for successful bone integration. Thus the healing
time can be reduced, leading to enhanced initial osseointegration.
In this study, we have proven the feasibility of creating a nanotextured surface on
titanium by using a simple acid/alkali treatment. The surface roughness can be tailored by
modifying the etching/ polishing procedures. Besides we have demonstrated that the cell
14
adhesion can be increased by coating the titanium surface with nanofibers. This is
because the nanofibers mimic the natural ECM and hence improve cell attachment.
Through our advanced electrospinning set-up we have achieved more fiber deposition at
a shorter interval of time than conventional electrospinning. Moreover we have shown
that the adhesion efficiency of the human bone marrow derived MSCs was the maximum
on the biomineralized PLGA/Collagen nanofibers coated Ti compared to the other
samples. Furthermore, incorporation of biomolecular cue like collagen and nano-HA
have enhanced the cell proliferation, osteogenic differentiation and cell mineralization.
To our knowledge, dental implant using functionalized nanofibers as a surface
modification is a novel idea to enhance osseointegration using the bone regeneration
concept.
15
Chapter 1
Introduction
1.1 Background
In the past 20 years, the number of dental implant procedures has increased steadily
worldwide, reaching about one million dental implantations per year [1]. Dental implants
are useful for restoration of oral function, including mastication and speech, as well as for
aesthetic improvement in patients with tooth loss. The clinical success of dental implants
is related to their early osseointegration. The other implant related to the early
osseointegration is total joint replacement, which is an effective treatment for relieving
pain and restoring range of motion. Osseointegration may be defined as the direct
structural and functional connection between living bone and the surface of a loadbearing artificial implant, typically made of titanium. It is a property virtually unique to
titanium and hydroxylapatite, and has enhanced the science of medical bone, and joint
replacement techniques. As long as implants are positioned correctly and infection is
avoided, they will generally last for many years. Geometry and surface topography are
crucial for the short- and long-term success of the implants. These parameters are
associated with delicate surgical techniques, a prerequisite for a successful early clinical
outcome. High success rates for dental implants are reported in healthy patients with
good bone quality. In the future, with an aging population, more patients may be
considered for dental implants; osseointegration of dental implants under less than
optimal circumstances and reduced bone healing quality may then be encountered. In
such cases, enhanced bone formation around the implant would be an important criterion.
16
This may be achieved by implant coatings that are able to interact actively with the
surrounding tissues.
1.2 Clinical problems associated with osseointegration:
There are two types of responses exhibited by the body after implantation. The first type
involves the formation of a soft fibrous tissue around the implant. This fibrous tissue does
not ensure proper osseointegration and leads to the clinical failure of the dental implant.
The second type of bone response is related to direct bone–implant contact without an
intervening connective tissue layer. This is the desired response after implantation. From
the clinical point of view, during osseointegration, two factors play an important role:
primary stability (mechanical stability) and secondary stability (biological stability after
bone remodelling). Primary stability is the mechanical stability of the implant as soon as
the implant is placed into the bone. It gradually decreases in the bone remodelling
process. Secondary stability involves the formation of new bone with the implant after
bone remodelling. Primary stability is fully replaced by secondary stability when the
healing process is completed. However, at one point, the implant stability decreases
during the stability conversion, a process also called the ―dip‖. Many implant failures
occur during this period, and this period seems to be critical to the successful integration
of the implant [2].
17
Figure 1: A typical Ti dental implant
18
1.3 Hypothesis and Objectives:
Hypothesis
This project is to develop a surface modification system for dental implant using
electrospun nanofiber and biomineralization to fabricate a biomimetic substrate. We
hypothesized that both substrate topographical and biochemical cues promote
mesenchymal stem cells (MSCs) adhesive behaviors, followed by proliferation and
differentiation, which are crucial for enhanced osseointegration.
Objectives:
Modify the implant surface to produce nanotextured topography.
Develop a nanofibrous coating from biodegradable synthetic polymers and/or
natural polymers to mimic extracellular matrix.
Functionalization of the nanofiber by biomineralization.
Evaluate adhesion, proliferation of MSCs on the modified implant surface.
Investigate osteogenic differentiation and mineralization of MSCs on the
modified implant surface.
19
Chapter 2
Literature Review
2.1 Introduction
The native Extracellular Matrix (ECM) consists of nano- to micro- structured fibers
(proteins and proteoglycans). This hierarchical organization presents a defined
environment with nano-scale intermolecular binding interactions that will affect the
morphological and functional development of the cells. Recent studies have shown the
importance of nano-textured implant surface for tissue engineering applications [3]. Cells
that were cultured on micro-size fibrous scaffolds were flattened and the cells spread as if
they were cultured on flat surfaces (Figure 2.1) [4]. Scaffolds with nano-scale
architectures have larger surface area to adsorb proteins and present many more binding
sites to cell membrane receptors would be more biomimetic to support better cell-matrix
interactions [4]. Thus the presentation of suitable topographical cues is an important
aspect to consider when designing tissue engineered scaffolds.
20
Figure 2.1 Scaffold architecture affects cell binding and spreading [4]
2.2. Surface modification techniques
To generate topographical cues on the implant surface, in order to enhance
osseointegration process, several surface modification techniques have been tried as
shown in the Table 2.1. The nanostructured surfaces of nanometallic and nanoceramic
materials have several advantages compared to the conventional surfaces. These include,
(i) they possess greater surface roughness resulting from both decreased grain size and
possibly decreased diameter of surface pores, (ii) enhanced surface wettability due to
greater surface roughness and (iii) greater numbers of grain boundaries. There are a
number of physical and chemical techniques that can be used for the surface modification
or activation of an implant surface. Among these methods, chemical modifications seem
to be relatively simple and inexpensive. Hence it is widely used. There have been various
techniques tried out in the past to improve the surface roughness of the implant like
plasma treatment, acid-etching and heat treatment. For example, the TPS (titanium
21
plasma sprayed) surfaces used by Straumann recommended a healing period of 12 weeks
[5] and this was reduced to 6 to 8 weeks with the introduction of the SLA (sand blasted,
acid etched) surface [6]. The differences in the contact angle and the surface roughness of
the implant surface owing to the various surface modification techniques were shown in
Table 2.1
Table 2.1 Different types of implant surface modifications and their surface roughness
and contact angle.
Type of implant
Surface
roughness Contact
(μm)
cpTi (Commercially pure Ra = 0.22 ± 0.01
angle References
(°)
55.4 ± 4.1
[7], [8]
Ti)
Ti6Al4V
Ra = 0.23 ± 0.01
56.3 ± 2.7
[7], [8]
TPS
Ra = 7.01 ± 2.09
n.d.
[7]
SLA
Sa = 1.15 ± 0.05
138.3 ± 4.2
[9]
Modified SLA
Sa = 1.16 ± 0.04
0
[9]
Plasma-sprayed HA coating
Ra = 1.06 ± 0.21
57.4 ± 3.2
[10], [11]
Mitsuru Takemoto et al., compared HCl–Alkali-heat treatment, alkali and heat treatment
and water–acid–alkali treatments [12]. He demonstrated that dilute HCl treatment
22
effectively removed sodium from the sodium titanate layer of alkali-treated porous
titanium and contributed to the formation of the titania layer on the surface of porous
bioactive titanium. Furthermore, the HCl–Alkali-heat treated implants possessed a more
complex surface when compared to other treatments, which may have been caused by an
etching effect of the dilute HCl treatment. The results of this study indicated that
chemistry and topography were related to material-induced osteoinduction as the dilute
HCl treatment was considered to give both chemical (titania formation and sodium
removal) and topographic (etching) effects on the titanium surface [12]. Timothy et al.,
adopted porous bone metal implant strategy to improve implant fixation, as it allows for
the ingrowths of bone and also reduces the Young’s modulus of the implant material to
better match that of bone. Besides bone ingrowths it also reduces the risks associated
with the bone resorption due to stiffness mismatch [13].
It was demonstrated that the treatment of Ti with a NaOH solution followed by heat
treatment at 873 K forms a crystalline phase of sodium titanate layer on the Ti surface
resulting in improved adhesion of apatite coating prepared by incubation in simulated
body fluid (SBF). The authors concluded that the released sodium ions from the sodium
titanate layer caused the formation of Ti–OH groups that react with the calcium ions from
the SBF and form calcium titanate, which then could act as nucleation sites for apatite
crystal formation [14, 15].
Lewandowska et al., characterized the chemical composition and morphology of titanium
surfaces exposed to acidic, alkaline or polymer solutions. It was found that there were
large differences in the morphology of Ti pretreated with different procedures whereas
23
only minor differences in the chemistry of the surfaces. In all the cases TiO2 being the
principle chemical component [16].
The Ti metal spontaneously forms a protective TiO2 layer in the atmosphere. When the Ti
implant is inserted into the human body, the surrounding tissues directly contact the TiO2
layer on the implant surface. The surface characteristics of the TiO2 layer determine the
biocompatibility of Ti implant. Therefore, it is important to use appropriate surface
modifications to increase the biocompatibility of the Ti implant for long-term clinical
applications. Several chemical etching agents like sodium hydroxide, hydrogen peroxide
and hydrofluoric acid have been used to improve the TiO2 layer, which is responsible for
the excellent corrosion resistance of the implant. In the body, however, mechanical
friction and chemical influences might lead to rupture or weakening of the TiO2 layer,
leading to a corrosion processes and the formation of wear debris in such regions [17].
Meanwhile, Nishiguchi et al., compared the bone-bonding ability of alkali- and heattreated titanium with that treated in NaOH without subsequent heat treatment. It was
concluded that the NaOH-treated titanium without heat treatment had no bone-bonding
ability due to its unstable reactive surface layer. He also demonstrated that soaking the
implant in NaOH solution stimulated the bone ingrowths onto the surface of the implant
[18].
2.2.1 Modification of the implant surface using surface adhesive molecules
In native tissues, ECM presents their adhesion proteins such as laminin, collagen,
fibronectin, and vitronectin to effect cell attachment through the binding between integrin
24
receptors on the cell surfaces. Therefore much work is done to enhance the
biocompatibility of polymeric tissue engineered scaffolds to create a biochemical-like
environment on the biomaterial surfaces [3].
Biomolecules such as adhesive proteins like collagen, RGD peptides, fibronectin and
growth factors like basic fibroblast growth factor and epidermal growth factor that can be
easily recognized by the cells can be coupled onto the biomaterials to induce biorecognition mechanisms of the interaction of cells and polymeric biomaterial scaffolds.
These modifications can preserve the mechanical integrity of polymeric scaffolds while
creating an ECM-like environment to the scaffolds. The surface chemistry of the implant
also plays an important role in deciding the cell characteristics. For example it was
reported that arginine-glycine-aspartic acid (RGD)-coated Ti disks greatly promoted
attachment and decreased apoptosis of MC3T3-E1 osteoprogenitor cells. Coating the
nanofibers with RGD or another positively charged molecule, such as calcium ion or
poly-lysine, may promote the attachment of cells. [19]
Currently, the most popular surface treatment for commercial artificial joints and dental
implants
is
plasma-spray coating
with
hydroxyapatite
(HA).
Plasma-sprayed
hydroxyapatite on titanium has been reported to show beneficial effects such as
osteoconductivity and direct-bone bonding ability [20]. However, the process has
disadvantages attributed to the high temperatures used during the process, such as the
possibility of fracture at the interface between the titanium and the HA due to the residual
stress at the interface, and changes in the composition, porosity, crystallinity, and
structure of the plasma-sprayed hydroxyapatite [21]. Therefore, new HA coating methods
25
have attracted great interests in recent years for replacing the high temperature techniques
like plasma spraying.
Besides, clinical trials were done by Wang et al., on canine trabecullar bone. He studied
the osseointegration of uncoated, Plasma- sprayed -HA-coated and electrodeposition –
HA -coated Ti–6Al–4V in a canine trabecular bone at 6 h, 7 days and 14 days postimplantation. The Plasma sprayed -HA was found to provide higher bone apposition ratio
than those exhibited by the bare alloy and electrodeposited-HA, owing to their earliest
mineralization (6 h—7 days) in the form of nano-ribbon cluster mineral deposits with a
Ca/P atomic ratio lower than that of hydroxyapatite [22]. In another study, pure titanium
was subjected to various surface modifications and examined in terms of morphology,
chemical characteristics and wettability. The results showed that etching in alkaline or
acid solutions resulted in significant changes in surface morphology; a characteristic
feature for the presence of sub-microporosity [23].
An earlier work done by Nicula et al., compared cp Ti, Ti–Al–V, Ti–Al–V–Cr and Ti–
Al–Mn–V–Cr prepared by high-energy ball-milling method, to achieve a microtextured
suface. Optimal cell adhesion was observed for the Ti–Mn–V–Cr–Al alloy, which might
be due to the surface morphology of this specimen (high-roughness, porosity in the
micron range). Thus the results showed that the surface properties are important for
implant materials, since the surface topography influences the mechanisms of cell
adhesion and growth [24].
26
The biomimetic scaffolds for tissue engineering can be manufactured by various
processes like electrospinning, phase separation, self–assembly and lithography.
Comparisons between the various techniques are shown in Table 2.2
Table 2.2: Various fabrication techniques along with their advantages and disadvantages
Fabrication
Advantages
Drawbacks
technique
Laser
Uniform distribution of pore size, simple Reduced resolution and
deposition
and fast.
Self assembly
Can generate fibrous networks capable Lack
poor surface finish
mechanical
of supporting cells in three dimensions. strength,
Limited
Cell-seeding problems associated with amphiphilic
using
prefabricated
scaffolds
eliminated
materials,
nanofibrous random and very short
owing
to nanofibers.
spontaneous assembly.
Lithography
Relatively good resolution.
Time consuming and
expensive.
Electrospinning The properties of electrospun nanofibers, Electrospinning yields a
such as fiber diameter, can be controlled flat mat that has limited
readily via manipulation of spinning three
dimensionality
parameters. Capable of mimicking the and suffers from cell
stem cell niche.
infiltration
problems
because of the small
pore size of the mats
Phase
A nanofibrous 3D scaffold can be Nanofiber distribution is
separation
constructed.
Has
controllable
high subject
to
the
the
27
porosity and surface-to- volume ratios.
processing.
2.2.2 Cell – substrate interaction
The implant's surface properties, surface chemistry, surface energy, topography and
roughness influence the initial cell response at the cell - material interface, ultimately
affecting the neo-tissue formation. Recent studies have shown higher osteoblasts
adhesion and enhanced alkaline phosphatase activity on rough Ti and Ti-6Al-4V [25, 26].
It is well known that cell response is affected by the physicochemical parameters of the
biomaterial surface, such as surface energy, surface charges or chemical composition.
Topography is one of the most crucial physical cues for stem cells and recently it has
been proven that nanotopography plays the main influencing factor, rather than
microtopography [27].
Though the surface modification techniques like grit-blasting, plasma treatment, sand
blasting, have been successful, the time required for osseointegration ranges from 3 to 6
months. Osteoblasts adhesion on nanostructured surfaces was first reported in 1999 by
Webster et al., [28]. He demonstrated that osteoblasts adhesion was improved when they
were cultured on nanostructured surfaces, compared to the conventional micro surfaces.
Specifically, alumina with grain sizes between 49 and 67 nm and titania with grain sizes
28
between 32 and 56 nm enhanced osteoblast adhesion compared to their respective micrograined materials.
It has been proved that the contact of cells to the surface of the biomaterials results in
changes to the cell shape and bioactivity depending on the topography of the surface [29].
For instance, cells cultured on pure Ti and Ti alloy exhibit differences in cell response
even though both are covered with TiO2 oxide layer. These differences may be attributed
to the surface morphology and chemistry differences between the two.
2.3 Tissue Engineering
2.3.1 Introduction
The current medical need is to address bone graft problems such as implant failure owing
to lack of tissue regeneration around the implant surface, resulting in poor bone
remodelling and loosening of the implants. In recent years, tissue engineering has
revolutionized the direction of research for orthopaedic applications because of the
success of nanotechnological advancements in creating new fabrication techniques for
nano-scale materials such as nanofibers and nanofibrous scaffolds. Previous studies
conducted by Ngiam et al., proved that n-HA on PLGA and PLGA/Collagen had a
positive modulation on early capture of osteoblasts compared to the non-functionalized
nanofibers. However no studies have been reported on the influence of hMSCs on the
functionalized nanofibers. The main advantage of using hMSCs for tissue engineering
applications is because of its direct clinical applications [30].
29
2.3.2 Nanofiber fabrication by electrospinning
Electrospinning is a simple and versatile technique that can produce non-woven ECMlike nanofiber scaffolds with nano-topographical cues to interact with the cells. Synthetic
polymeric nanofibers such as poly(ε-caprolactone) (PCL) [31], poly(L-lactic acid)
(PLLA) [32], poly(glycolic acid) (PGA) [33] and poly(lactic-co-glycolic acid) (PLGA)
[34], and natural-occurring polymeric nanofibers such as collagen [35] and gelatin [36]
have been widely explored for applications in the different areas of tissue engineering
such as skin, cartilage, bone, blood vessel, heart, and nerve [31 - 40]. The properties of
the commonly used polymers are discussed in Table 2.3.
Table 2.3 Commonly used polymers and their properties
Polymer
Properties
Degradation rate
Reference
PGA
Aliphatic polyester,
6-12 months
[41]
> 24 months
[42 – 44]
Crystalline, semi permeable
PLLA
Aliphatic polyester,
crystalline, porous; Roughlooking due to the open-pore
structure
PLGA
Semi pemeable
6 – 12 months
[40]
PCL
Semi permeable, amorphous
< 12 months
[45]
Collagen
Semi pemeable
1 – 9 months
[46]
30
Electrospinning process utilizes an electric field generated by an applied voltage that
subsequently introduces surface charges to the polymer solution. This results in the
formation of a Taylor cone polymeric droplet at the tip of the spinneret. Once the electric
potential that is created at the droplet surface exceeds a critical value, the electrostatic
forces will overcome the solution surface tension to initiate a polymer jet stream. The
charged jet is accelerated towards the grounded collector and undergoes bending
instability, elongation, and solvent evaporation or jet solidification which leads to rapid
thinning of the jet and deposition of dry fibers in a random manner onto the collector [33,
41, 42]. The experimental set up for electrospinning is shown in Figure 2.2. Several
factors can affect the electrospinning process and fiber morphology (Table 2.4).
Table 2.4. Factors that affect the electrospinning process and fiber morphology [47].
Process Parameter
Viscosity/concentration
Effect on fiber morphology
Low concentrations/viscosities yielded defects in the
form
of
beads
and
unction;
increasing
concentration/viscosity reduced the defects;
Fiber
diameters
increased
with
increasing
concentration/viscosity.
Conductivity
Increasing the conductivity aided in the production
of uniform bead-free fibers;
31
Higher conductivities yielded smaller fibers in
general (except PAA and polyamide-6).
Polymer
molecular
weight
Dipole
Increasing molecular weight reduced the number of
beads and droplets.
moment
and
dielectric constant
Flow rate
Successful spinning occurred in solvents with a high
dielectric constant.
Lower flow rates yielded fibers with smaller
diameters;
High flow rates produced fibers that were not dry
upon reaching the collector.
Field strength/voltage
At too high voltage, beading was observed;
Correlation between voltage and fiber diameter was
ambiguous.
Distance between tip
and collector
A minimum distance was required to obtain dried
fibers;
At distance either too close or too far, beading was
observed.
Fiber morphology
Smooth fibers resulted from metal collectors;
Aligned fibers were obtained using a conductive
32
frame, rotating drum, or a wheel-like bobbin
collector;
Ambient parameters
Yarns and braided fibers were also obtained.
Increased temperature caused a decrease in solution
viscosity, resulting in smaller fibers;
Increasing humidity resulted in the appearance of
circular pores on the fibers.
Figure 2.2. Schematic diagram of electrospinning set-up
In a work done by Ma et al., three different materials, silicon (Si), silicon oxide (SiO2),
and titanium oxide (TiO2), were used to construct nanofibers for surface coating of Ti
alloy Ti-6Al-4V. The results demonstrated that TiO2 nanofibers coated over the Ti alloy
facilitated a higher adhesion potential and higher cellular differentiation capacity than Ti
alloy and tissue culture–treated polystyrene surfaces (TCP). Thus, surface modification
33
using nanofibers of various materials was proved to alter the attachment, proliferation,
and differentiation of osteoprogenitor cells in vitro [48].
It was also reported that nanofibrous poly (L-lactide) (PLLA) scaffold fabricated by
phase separation and particle-leaching method showed biological function similar to
those of the collagen fibers of bone [49]. These results might implicate the possibility that
a nanofibrous surface can improve the osseointegration of implants. However, these
nanofibrous materials such as carbon and organic polymer are difficult to be immobilized
on titanium surface because of their low reactivity with titanium [50].
2.3.3 Modifications of the electrospun nanofibers
At present HA has been widely used as bioceramics in orthopaedics and dentistry due its
osteoconductive properties [29]. In the native bone tissue, HA nanocrystals grow in
intimate contact within collagen fibers, building up a nano-structured composite.
However, HA has a disadvantage that is attributed to low mechanical strength. Hence the
combination of a load bearing biomaterial like titanium with the osteoconductive
properties of HA is very attractive. HA related bone formation is believed to begin with
surface dissolution of the HA, which releases calcium and phosphate ions into the
vicinity around the implant. Reprecipitation of carbonated apatite then occurs on the
coating surface, thereby enhancing osteoblasts adhesion onto the surface.
34
Immobilization reaction of TiO2 nanofibers on the titanium plate was done by treating the
Titanium plates firstly in alkali and then in acid solutions. When immersed in NaOH, the
passive oxide layer of titanium dissolves to form amorphous titanate layer containing
Na+ ions. Immediately after immersion in simulated body fluid (SBF), Na+ ions from the
amorphous layer will be exchanged by H3O+ ions from the surrounding fluid resulting in
the formation of Ti–OH layer. And then hydroxyapatite was formed on titanium surface
by ionic bonding between Ti–O anions and Ca2+ cations in SBF [51].
Thus,
biomineralization originated from native process may provide some effective way for
osseointegration. In another study collagen fibrils/carbonate-hydroxyapatite coating has
been electrodeposited on Ti plates using Ca (NO3)2 and NH4H2PO4 solutions in a type I
collagen molecule suspension [52].
2.3.4 Potential application of mesenchymal stem cells for osseointegration
Stem cells are unspecialized cells that can self renew indefinitely and differentiate into
several somatic cells with proper environmental cues. In stem cell niche, the stem cell–
ECM interactions are very crucial for different cellular functions like adhesion,
proliferation and differentiation. Most recently, the importance of nanometric scale
surface topography, and roughness of biomaterials is, besides chemical surface
modifications, increasingly becoming recognized as a crucial factor as synthetic ECM for
cell survival and host tissue acceptance.
35
Recent work by Muschler [53, 54] demonstrated that it is possible to capture MSCs on
substrate such as allograft bone. He has developed a system where it is able to capture
MSCs on allograft bone with an enrichment-factor of 3-4x at best. A much higher
theoretical capture ratio is possible. The fabrication of a nanofibrous scaffold offers the
possibility to optimize cell capture as well as cell adhesion and proliferation.
Furthermore, MSCs derived from the bone marrow of neonatal rats, were used for
seeding on electrospun PCL scaffolds by Yoshimoto et al., [31]. MSCs not only attached
favourably and grew well on the surface of these scaffolds, but the MSCs were also able
to migrate inside the scaffold up to 114 µm within 1 week of culture.
Gelatin/PCL shows better biocompatibility than PCL nanofibrous material. The enhanced
adhesion and proliferation of MSCs on nanofibers matrix also showed up on PLA and
silk electrospun nanofibers [50, 51]. Hosseinkhani et al., investigated mesenchymal stem
cell (MSC) behavior on self-assembled peptide-amphiphile (PA) nanofiber scaffolds [55]
Significantly enhanced osteogenic differentiation of MSC occurred in the 3D PA scaffold
compared to 2D static tissue culture.
36
Chapter 3
Biomimetic surface modification of dental implant materials by
advanced electrospinning
3.1 Introduction
Implants can be divided into smooth (machined or turned) and rough on the basis of
surface roughness. The techniques used for preparing the surface roughness maybe either
additive or subtractive in nature. Additive techniques involve coating the implants with
titanium or HA using plasma spraying technique or sintering. Subtraction techniques
involve the use of sandblasting or acid/alkali etching treatments.
Titanium (Ti) has been widely used as implant materials in the dentistry and orthopaedics
owing to their excellent mechanical properties and biocompatibility [1]. Some of these
properties, in particular the biological response of titanium, are strongly determined by
the surface characteristics— its morphology, chemistry and physical properties. Ti and Ti
alloy facilitate new bone formation and provide long-lasting bone-implant stability. In
addition to being bio-inert and nontoxic, requirements for the next generation of
biomaterials include enhanced cell attachment and differentiation to accelerate
osseointegration of implants. Modified or coated Ti and its alloys have become
candidates for next-generation implants. Surface properties may be changed by applying
various surface modifications while the crucial bulk properties such as tensile strength
and fatigue resistance remain unchanged. However implant failures do occur owing to
37
loosening of the implants. One of the main strategies to enhance osteoconduction is the
use of a nanofiber-coated surface [56]. A nanofiber coating on Ti constructs a rough
surface, which may stimulate bone formation by triggering specific cell responses. Our
strategy is to design and fabricate biomimetic and bioactive implant surfaces that
resemble the native extracellular matrix (ECM) as closely as possible so as to create
conducive living milieu that will induce cells to function naturally. In this context, our
current endeavor is to use the natural polymer collagen along with PLGA as a matrix and
to deposit n-HA (nano – HA) by Calcium-phosphate (Ca-P) dipping method so as to
develop biomimetic n-HA containing nanocomposite nanofibers.
3.2 Materials and Methods
3.2.1 Mechanical Polishing/ etching:
Pure Titanium (15mm diameter) and Titanium alloy (Ti- 6Al- 4V) discs (25mm
diameter), purchased from Northwest Institute for Non – Ferrous Metal Research (Xian,
Shanxi, P.R. China) were mechanically polished using 320 grit and 400 grit SiC papers,
till a mirror finish was achieved. The discs were further polished using alumina (1M)
cloth for a smoother finish. The discs were then cleaned with ethanol using an
ultrasonicator for 15min. This ensures the removal of the impurities arisen due to the
mechanical polishing. The mechanical treatment was followed by chemical etching using
4% HNO3 in ethanol for 1min. The discs were then allowed to dry at room temperature.
38
3.2.2 Pretreatment of Ti
The polished/etched Ti plates were immersed in 10N (Normality) NaOH solution at 600C
for 24 hrs. The samples were then allowed to cool to the room temperature, followed by
treatment with 10N HCl solution for 1hr. The samples were then dried.
Titanium implants after the alkali treatment retained sodium and the sodium titanate layer
with limited formation of titania layers. To overcome these problems, in addition to water
treatment, a dilute hydrochloric acid (HCl) treatment was done, which almost completely
removes sodium, even from deep pores [12].
3.2.3 Electrospinning of PLGA and PLGA/Collagen nanofibers on the Ti discs
The materials used for electrospinning were Type I collagen (Koken Co. Tokyo, Japan),
PLGA (100,000 Da, Aldrich Chemical Company, Inc., St. Louis, U.S.) and 1,1,1,3,3hexafluoro-2-propanol (HFP, Aldrich Chemical Company, Inc., St. Louis, U.S.). PLGA
(75:25) pellets were dissolved in HFP at a w/v ratio of 15%. The electrospinning
parameters, the w/v ratio of PLGA in HFP and the fiber deposition time were optimized
till uniform nanofibers without bead formation was obtained, as shown in Table 3.1.
Electrospinning of blended PLGA/Collagen (50:50 w/w ratio) was also done following
the same procedure (Table 3.2).
The polymer solution was then loaded into a syringe (Becton Dickinson, BD, N.J, U.S.)
and a high voltage electric field (DC high voltage power supply from Gamma High
39
Voltage Research, Florida, U.S.) was applied to draw the fibers from the spinneret
(27G1/2 needle, Becton Dickinson, BD, N.J, U.S.) onto the collector plate, over which
the Ti disc was placed. The experimental setup was shown in figure 3.1. The spinneret
was first grounded to give a flat tip in order to produce continuous and uniform
nanofibers. A constant feed rate of 1 mL/h was applied using a syringe pump (KD
Scientific Inc., M.A., U.S.).
Polymer solution
spinneret
Ring electrode
Collector
Figure 3.1 Electrospinning set up
40
Needle tip
a
Ring
electrode
b
Z
Z
X
X
Figure 3.2 Electric field pattern a) 18kV at the needle tip and 10kV at the ring electrode
18kV at the needle tip and 14kV at the ring electrode
b)
In order to achieve focussed electrospinning, an additional ring electrode was used. A
high voltage electric field supply was connected to the ring electrode as shown in Figure
3.1. The voltages across the two power supplies were optimized by analysing the electric
field pattern developed from the two voltages, as shown in Figure 3.2. It was found that
when the voltage at the needle was 18kV and the voltage at the ring electrode was 14kV,
the electrospinning process was limited. This was because, at 14kV there were starting to
appear irregularities in the electric field, as shown in the Figure 3.2. However when the
voltage at the ring electrode was reduced to 10kV the electric field was more uniform
without any features that could disturb electrospinning process. As the voltage was
increased, the fiber deposition spot decreased.
41
The electrospun nanofibers were subsequently vacuum dried so that any residual solvent
present could be removed.
3.2.4 Biomineralization using Calcium-Phosphate dipping method
Biomineralization of nano-hydroxyapatite (n-HA) was achieved using Ca-P dip method.
The Ti plates coated with electrospun PLGA and PLGA/Collagen nanofibers were
initially immersed in 0.5M CaCl2 solution (Aldrich Chemical Company, Inc., St. Louis,
U.S.), for 10 min. The samples were then rinsed in deionised water for 1min. The
samples were then immersed in 0.3M of Na2HPO4 (Merck & Co. Inc., N.J, U.S.), for
10min and rinsed for 1min in DI water. This entire procedure was considered as 1 cycle.
The scaffolds were subjected to 3 cycles of the above treatment. The first cycle was for
10min and the subsequent cycles were for 5min in each solution. After that, the scaffolds
were removed and freeze dried overnight. The above process is schematically illustrated
in Figure 3.3.
42
Biomineralization Procedure:
Ti plate
xxxxxxxxxxx
x
Nanofibers
Freeze Dry (overnight)
Figure 3.3 Biomineralization procedure
3.2.5 Cell adhesion study
All the titanium samples were sterilized under UV light for 2 hrs. The discs were then
washed with phosphate buffered saline (PBS) thrice for 15min each. Human bone
marrow derived mesenchymal stem cells (hMSCs) (PT-2501, Lonza, USA) was cultured
in DMEM low glucose medium (DMEM, Invitrogen, CA, U.S.) with 10% FBS
(Invitrogen, CA, U.S.) and 1X Antibiotics (Sigma-Aldrich Chemical Company Inc., St.
Louis, U.S) until 3 passages. The cells were then trypsinized and seeded onto all the
samples (untreated Ti, Ti + PLGA nanofibers, Ti + PLGA/Collagen nanofibers, Ti +
43
PLGA nanofibers + n-HA, Ti + PLGA/Collagen nanofibers + n-HA) at the concentration
of 10,000 cells per well. The well plates were incubated for different time points – 10, 20,
30 and 60 minutes at room temperature. After the incubation time, the media was
removed and the plates were washed thrice with PBS to remove the unbound cells. The
attachment efficiency of each sample was then evaluated using field emission scanning
electron microscopy (FESEM). The samples were fixed with 3% glutaraldehyde for 3
hours. The scaffolds were then rinsed with distilled water for 15min and then dehydrated
with a series of ethanol gradients starting from 30% to 50%, 75%, 90% and 100% (v/v).
Subsequently the samples were treated with HMDS (Hexamethyldisilazane) solution and
allowed to air – dry at room temperature in the fume hood. The samples were then gold
coated (JEOL JFC-1200 fine coater, Japan) and the cells were counted using FESEM.
Five locations were chosen on each scaffold – four corners and centre. The average
number of cells was then counted from the chosen locations. The distance of the location
was measured using Image J software. The total number of cells was then calculated for
the entire dimension of the scaffold.
3.2.6 Surface characterization analysis
The effect of the pretreatment on titanium was characterised using field-emission
scanning electron microscopy (FESEM) (Quanta 200F, FEI, Oregon, U.S.). Prior to
which the samples were gold coated (JEOL JFC-1200 fine coater, Japan).
The biomineralized and the non-biomineralized electrospun nanofibers were also
characterised using FESEM, to analyse the distribution of HA formed owing to the in
44
vitro biomineralization procedure. The diameter of the nanofibers was measured using
image analysis software (Image J, National Institutes of Health, Bethesda, U.S.).
3.2.7 Surface roughness analysis
The surface roughness of the Titanium discs were analysed before and after the
pretreatment procedure using Profilometer (Surftest SV-400, Mitutoyo, U.S.), using a
PC50 filter and at the rate of 0.5mm/s. Five discs were measured from each of the
different substrates and the average roughness value Ra was calculated.
The discs were further characterized using the Atomic Force Microscope (AFM) in
tapping mode (Dimension 3100 AFM, Veeco Instruments Inc., CA, U.S.). The roughness
data was then analysed using the Nanoscope software (Digital Systems, US). Similar to
the profilometer analysis, five discs were chosen in random for each of the procedures
and the average roughness value was analysed.
3.2.8 Fourier transform infrared spectroscopy (FTIR) and X-ray photoelectron
spectroscopy (XPS)
The TiO2 oxide layer on the titanium samples before and after the pretreatment was
analysed using ATR mode FTIR (Bio-Rad FTIR FTS 3500). A universal sampling
aperture at a grazing angle of 670 with respect to the surface was used. A spectral
resolution of 4 cm-1 in the 400 – 4000cm-1 range was employed to analyze the TiO2 oxide
45
layer. Besides the electrospun PLGA and PLGA/Collagen nanofibers were also analyzed
using FTIR to determine the functional groups present in them.
The chemical composition of the untreated and the pretreated Ti discs were analyzed
using XPS (Kratos AXIS HSi, X-ray Source: Mono Al K alpha, 15 kV, 10 mA (150 w)).
The following were the parameters used: Photoelectron accept angle: 90 degree. Base
pressure: 1.0X10 -9 Torr, working pressure: 1.0X10-8 Torr.
3.2.9 Water contact angle measurement
The contact angle of the Ti discs was measured before and after the pretreatment
procedure to study the influence of the treatment procedure on the wettability of the
substrate. The contact angle measurements were done using VCA Optima Surface
Analysis system (AST products, Billerica, MA). Distilled water was used for drop
formation.
3.2.10 Statistical analysis
Values (at least triplicate) were averaged and expressed as mean ± standard deviation
(SD). Each experiment was repeated twice for cell adhesion. Diameter of nanofibers was
calculated from 5 SEM images by randomly selecting 10 fibers from each SEM image.
Statistical differences were determined by Student two-sample t test. Differences were
considered statistically significant at p ≤ 0.05.
46
3.3 Results and Discussion
3.3.1 Surface characterization analysis
Electrospinning is a very simple and versatile process by which polymer nanofibers with
diameters ranging from a few nanometers to several micrometers can be produced using
an electrostatically driven jet of polymer solution. Figure 3.1 shows the conventional
electrospinning set-up. In a typical electrospinning process, an electrical potential is
applied between a droplet of a polymer solution, held at the end of a capillary tube and a
grounded target. When the applied electric field overcomes the surface tension of the
droplet, a charged jet of polymer solution is ejected. The route of the charged jet is
controlled by the electric field. This inherent property of the electrospinning process,
favors the control of deposition of polymer fibers onto any target substrate [57]. In the
work by Theron et al., an electrostatic field-assisted assembly technique was described,
which in combination with an electrospinning process was used to position and align
nanofibers on a tapered and grounded wheel-like bobbin [58]. Our experiment suggested
the targeted deposition of the nanofibers onto the cpTi (commercially pure Ti) and Ti
alloy (Ti6Al4V) discs. Results were also reported by Hohman, Shin, Rutlege and
Brenner, who studied electrospinning with regard to electrically forced jet and
instabilities, and proposed a stability theory for electrified fluid jets [59 - 62]. It was
demonstrated that at increasing field strengths, the electrical instabilities are enhanced
[59]. This was in correlation with the results obtained in our study using advanced
electrospinning technique; as shown in Figure 3.2. Wherein, on the left diagram it is
18kV at the needle and 10kV at the ring; on the right it is 18kV at the needle and 14kV at
the ring. As shown, for 14kV there are starting to appear irregularities in the electric field.
47
A bit higher voltage and this irregularity was bigger, also indicating a decrease in the
fiber deposition diameter. In fact it is repelling the electric field from the ring, but still too
small to stop the electrospinning action. On the contrary electric field for 10kV at the ring
is much more uniform without any features that could disturb electrospinning process.
Green lines indicated in the Figure 3.2 are only for representation of the decrease of the
fiber deposition spot. These are protons moving in given electric field and cannot be
referred to as nanofiber. Additionally in correlation with the earlier studies it is seen that
with higher voltage, fiber deposition spot should be smaller as proven in the experiment.
The deposition of n-HA was achieved by a feasible Ca-P dipping method as shown in
Figure 3.3.
The SEM images (Figure 3.4a) of the untreated Ti discs revealed no distinctive surface
topography. However after the pretreatment, the SEM images show topographical
distribution of α and β grains (Figure 3.4b). This shows that the acid treatment has led to
conversion of the initial microtextured surface to a nanotextured surface. The
morphology of the PLGA and the PLGA/Collagen nanofibers were analyzed using SEM
images (Figure 3.4c – 3.4f). It was not possible to measure the tensile strength of the
PLGA and PLGA/Collagen coated implant surface owing to the hard nature of the
implant material and the thin layer of nanofiber coating on the implant surface. The
tensile properties of PLGA and PLGA/Collagen nanofibers have been reported by Kun
Ma et al., where Young’s modulus (MPa) of PLGA and PLGA/Collagen nanofibers was
reported to be 190.42 ± 9.97 and 40.43 ± 3.53 respectively. Ultimate tensile stress (MPa)
was 4.82 ± 0.33 and 1.22 ± 0.12 for PLGA and PLGA/Collagen nanofibers respectively
[63]. The deposition time and the concentration parameters were varied till an optimum
48
fiber diameter was achieved as shown in Table 3.1 and Table 3.2. From the SEM
micrographs it was seen that as the fiber deposition time increased beyond 15 seconds,
the amount of fiber deposited also increased. In order to ensure osseointegration, it is
therefore desirable that the cells contact both the nanofibers and the nanotopography of
the Ti substrate. For the deposition time of 10 seconds and 5 seconds, the fiber deposition
was not uniform as can be seen from the SEM images. Hence 15 seconds was chosen as
the optimum deposition time for further studies. From the Tables 3.1 and 3.2, it was seen
that as the polymer concentration was increased, the fiber diameter also increased. For
polymer concentrations less than 10% and 15% in the case of PLGA/Collagen and PLGA
respectively, even though the fiber diameter was smaller, bead formation and nonuniformity of fibers was seen. Hence for further studies PLGA nanofibers were spun at a
polymer concentration of 15% and deposition time of 15 seconds and PLGA/Collagen at
a concentration of 10% and deposition time of 15 seconds.
The deposition of n-HA after the biomineralization treatment using Ca-P dipping method
was also shown using FESEM (Figure 3.4 g and 3.4 h). HA nanocrystals grow in intimate
contact within collagen fibers, building up a nanostructured composite. However, HA has
disadvantage attributed to low mechanical strength for implant applications. Hence the
combination of a load bearing biomaterial like titanium with the osteoconductive
properties of HA is very attractive. It was found that the n-HA deposition was uniform
and more predominant on the PLGA/Collagen nanofibers than on the PLGA nanofibers.
The attachment of nano-HA was more on PLGA/Collagen compared to PLGA nanofibers
because collagen is more hydrophilic and mimics the natural bone, thereby favouring
nano-HA deposition.
49
Table 3.1 Optimization of electrospinning parameters by varying the time and concentration for
PLGA nanofibers
Electrospinning PLGA nanofibers at various concentrations (15%, 18% and 20%
w/v) and time periods (5 seconds to 2 minutes)
Nanofiber diameter ± SD (micro meter)
Time
15%
18%
20%
2 min
0.68 ± 0.282
1.761 ± 0.371
1.800 ± 0.213
1.5 min
0.774 ± 0.227
1.212 ± 0.39
0.914 ± 0.151
1min
0.543 ± 0.153
1.114 ± 0.357
1.089 ± 0.267
30sec
0.768 ± 0.314
1.326 ± 0.479
1.067 ± 0.194
15 sec
0.957 ± 0.357
1.615 ± 0.472
1.731 ± 0.386
10sec
0.996 ± 0.344
1.721 ± 0.413
1.264 ± 0.269
5sec
0.759 ± 0.415
1.535 ± 0.594
1.381 ± 0.449
Table 3.2 Optimization of electrospinning parameters by varying the time and concentration for
PLGA/Collagen nanofibers
Electrospinning PLGA/Collagen nanofibers at various concentrations (10% and 15%
w/v) and time periods (5 seconds to 2 minutes)
Nanofiber diameter ± SD (micro meter)
Time
10%
15%
2 min
0.549 ± 0.213
0.827 ± 0.116
1.5 min
0.279 ± 0.085
0.898 ± 0.176
1min
0.368 ± 0.089
0.801 ± 0.147
30sec
0.251 ± 0.093
0.776 ± 0.136
15 sec
0.378 ± 0.068
0.817 ± 0.151
10sec
0.410 ± 0.093
0.828 ± 0.185
50
5sec
a)
c)
0.310 ± 0.089
0.783 ± 0.454
b)
d)
51
e)
f)
g)
h)
Figure 3.4 SEM images of a) untreated cpTi, b) cpTi after surface modification c) cpTi coated
with PLGA nanofibers at 1000X magnification d) cpTi coated with PLGA/Collagen nanofibers at
1000X magnification e) cpTi coated with PLGA nanofibers at 5000X magnification f) cpTi
coated with PLGA/Collagen nanofibers at 5000X magnification g) cpTi coated with
functionalized PLGA nanofibers h) cpTi coated with functionalized PLGA/Collagen nanofibers
3.3.2 Surface Roughness analysis
It is generally accepted that rough, textured and porous surfaces are able to stimulate cell
attachment, differentiation and formation of the extracellular matrix [64]. Moreover, an
52
appropriate surface roughness can produce beneficial mechanical interlocking at the
initial adhesion stage and aid in further cell adhesion [65].
Profilometer analysis revealed that the Ra value of the untreated cp Ti and Ti-6Al-4V
discs were 0.306 μm and 1.529 μm respectively. However after the pretreatment
procedure, the Ra values reduced to 0.022 μm and 0.042 μm respectively. This proves
that the microtextured samples, after the pretreatment have attained a nanotextured
surface. Thus the chemical pretreatment procedure using NaOH and HCl was a very
feasible procedure to achieve nanotopography. Nanotopography plays a very important
role in stem cell adhesion, proliferation and differentiation [63, 66].
Moreover, an
appropriate surface roughness can produce beneficial mechanical interlocking at the
initial adhesion stage and aid in further cell adhesion [64, 65]. Cell adhesion and
proliferation were reported to be more on a nanotextured surface than on a microsurface.
Hence it was essential that an implant surface has a nanotexture, in order to accelerate the
osseointegration process in vivo.
It was not possible to measure the surface roughness of the untreated Ti, using AFM, as it
was highly rough. But the Ra results from AFM (Figure 3.5) for the treated samples
complemented the profilometer results. The Ra values for the pretreated Ti and Ti alloy
were calculated using the Nanoscope software and were found to be 15.2nm and 18.9 nm
respectively.
53
Figure 3.5 AFM image of pretreated Ti showing the surface roughness
3.3.3 FTIR
The FTIR results (Figure 3.6) showed an increase in the TiO2 oxide layer thickness in the
pretreated samples. The TiO2 peak in the case of Ti was located at 667 cm-1. Though the
untreated samples also showed the TiO2 peak, it was more significant in the case of the
pretreated samples. The increase in the TiO2 oxide layer improves the biocompatibility of
the implant [67]. Besides the new peaks in the 400-800 cm-1 wavenumber range, in the
pretreated samples corresponds to the other oxides of Ti like TiO and Ti2O3.
54
In pure PLGA nanofibers the C=O stretch and the C–O stretch hovered around 1761 cm−1
and 1088 cm−1 respectively. Amide I and amide II of collagen were detected at 1658 cm−1
and 1544 cm−1 in PLGA/Collagen nanofibers.
a)
55
b)
c)
Figure 3.6 FTIR results for a) pure Ti treated and untreated, b) Ti-6Al-4V alloy treated and
untreated, c) PLGA and PLGA/Collagen nanofibers coated over the Ti surface.
56
3.3.4 Water contact angle measurement
The contact angle results showed that the pretreatment procedure has improved the
hydrophilicity of the scaffold. The contact angle of the control samples were 106.10 ±1.70
for cp Ti and 71.10 ± 5.40 for Ti alloy. The contact angle was reduced to 16.02 ± 0.80 and
11.04 ± 1.10 for cp Ti and Ti alloy respectively after the pretreatment as shown in Table
3.3. A decrease in the water contact angle indicates that the substrate has become more
hydrophilic. An increase in the wettability of the scaffold was said to improve cell
adhesion [66]. The functional groups present in collagen, i.e. carboxyl groups and
carbonyl groups [68] and [69], served as nucleation sites for apatite formation and
consequently, uniform distribution of n-HA was apparent on the outer and inner surfaces
of the PLGA/Collagen nanofibers compared to the PLGA nanofibers. Besides being a
favorable site for nucleation, the –COOH functional groups of collagen increased the
hydrophilicity of the nanofibers.
The contact angle measurements were taken after depositing the PLGA and
PLGA/Collagen nanofibers over the scaffold by electrospinning. The PLGA scaffolds
had a contact angle of 9.9 + 0.3° and the PLGA/Collagen scaffolds had a value of 0 as
shown in Table 3.4. This was because collagen is very hydrophilic in nature. The
hydroxyl groups present in the collagen forms hydrogen bonds with water molecules thus
imparting the relevant hydrophilicity. Hence after incorporating collagen to PLGA, the
water contact angle decreases.
57
Table 3.3 Water contact angle measurements for treated and untreated cp Ti and Ti6Al4V
SAMPLE
WATER
CONTACT
ANGLE(°)
Before pretreatment
Ti
106.10 ± 1.7
After pretreatment
Ti
16.02 ± 0.8
Before pretreatment
Ti alloy
71.10 ± 5.4
After pretreatment
Ti alloy
11.04 ± 1.1
Table 3.4 Water contact angle measurements for PLGA and PLGA/Collagen nanofibers
SAMPLE
WATER CONTACT ANGLE (°)
PLGA nanofibers
9.9 ± 0.3
PLGA/Collagen nanofibers
0
3.3.5 XPS
Figure 3.7 shows distinct peaks obtained in the range 450 – 470eV in the treated Ti
sample, corresponding to Ti 2p. We find that in the treated samples new peaks arise in
this range corresponding to the Ti oxides like Ti2O3, TiO2 and TiO. Besides the Oxygen
1s peak also increased indicating an increase in the oxide layer formed over Ti. The XPS
results proved that owing to the pretreatment of Ti, the oxide layers have increased,
which in turn improves the biocompatibility of the implant surface [67]. Improving the
oxide layer also favours enhanced initial osseointegration. The peaks for Na 1s and Cl 2p
58
increased in the pretreated samples owing to the NaOH/HCl pretreatment procedure
followed.
Figure 3.7 XPS results showing the Ti2p peaks in the treated samples
3.3.6 Cell culture analysis
As depicted in the Figures 3.8A and 3.8B, the biomineralized nanofibers show enhanced
cell adhesion when compared to the non- biomineralized nanofibers and the untreated
titanium samples. In the untreated samples no significant adhesion of cells occurs even
after 60min. This suggests that the untreated cp Ti and Ti alloy samples are not suitable
for early cell adhesion. However at 10min, no statistical difference (p ≤ 0.05) was found
between the Ti coated with PLGA/Collagen and the Ti coated with PLGA/HA. Also in
the case of Ti alloy at 10min there was no significant difference between Ti alloy coated
with PLGA and PLGA/HA. No statistical difference was also observed between Ti alloys
coated with PLGA/Collagen and PLGA/Collagen/HA. This maybe because 10min
59
duration was too short for HA to cause a significant cell adhesion. At 30min and 60min
the adhesion onto the Ti/PLGA/Collagen/HA and Ti alloy/PLGA/Collagen/HA substrate
was statistically significant from all the other Ti and Ti alloy samples respectively,
indicating that maximum adhesion of hMSCs occurred on the mineralized
PLGA/Collagen scaffolds compared to the other scaffolds. Figure 3.8A and 3.8B shows
one of the SEM images taken for the cp Ti and Ti alloy samples respectively at the time
intervals – 10, 20, 30 and 60 min.
Studies have reported lower cell adhesion and proliferation on less organized surfaces
(i.e. sandblasted ones) [70]. Hence for this reason, in the present study, regular and
uniform surface roughness on the surface of all samples was produced, resulting in
homogeneous surface texture on all the cpTi and Ti alloy disks. This study thus proves
that the nano-hydroxyapatite coated on the nanotextured titanium surface improves the
initial cell attachment, which is very crucial for enhanced osseointegration. It was found
that the cell adhesion was more on the biomineralized scaffolds compared to the nonbiomineralized scaffolds. This is because collagen along with n-HA synergistically
enhances early cell capture. Besides, calcium ions have also been suggested to be
advantageous to cell growth [71]. However the exact mechanisms by which calcium
phosphate ceramics improve bone bonding are not clearly understood, although it is
known that the bioactivity of ceramics is related to the dissolution rate and that the early
cellular response is of primary importance [72]. Similarly adhesion study has also been
reported earlier by Kun Ma et al., using bone marrow derived haematopoitic stem cells on
PLGA and PLGA/Collagen nanofibers coated with a surface adhesion molecule, E selectin. The study revealed that the haematopoitic stem cells capture efficiency on the
60
PLGA/Collagen nanofiber scaffold after coated with E-selectin, significantly increased
cell capture percentage from 23.40% to 67.41% within 30 min and from 29.44% to
70.19% within 60 min of incubation at room temperature [63]. Nevertheless our results
have indicated nearly 75% cell adhesion on to the bioceramic coated surface on both cpTi
and Ti alloy samples as shown in Figure 3.9. Table 3.5 indicates the statistics for the
number of cells adhered to the cpTi and Ti alloy implant surfaces at various
time intervals. The rationale for conducting a short-term cell adhesion study
on the nanofibrous scaffolds was to assess the viability of avoiding extended
culture periods of cell seeding on the substrates, thereby reducing the down-time from
material preparation to the material implantation in the patient, preferably in-situ during
surgery.
10 min
20 min
30 min
60 min
a)Untreated cpTi
b)cpTi + PLGA
61
c)cpTi+PLGA/Col
d)cpTi+PLGA/HA
e)cpTi+PLGA/Col/HA
Figure 3.8A: Adhesion of hMSCs on the a) untreated cpTi implants, b) cpTi implant coated with
PLGA nanofibers, c) cpTi implant coated with PLGA/Collagen nanofibers, d) cpTi implant
coated with PLGA/HA, e) cpTi implant coated with PLGA/Collagen/HA nanofibers at 500x
10 min
20 min
30 min
60 min
a) untreated Ti6Al4V
62
b) Ti6Al4V+PLGA
c)Ti6Al4V + PLGA/Col
d)Ti6Al4V + PLGA/HA
e)Ti6Al4V+PLGA/Col/HA
Figure 3.8B: Adhesion of hMSCs on the a) untreated Ti6Al4V implants, b) Ti6Al4V implant
coated with PLGA nanofibers, c) Ti6Al4V implant coated with PLGA/Collagen nanofibers, d)
Ti6Al4V implant coated with PLGA/HA, e) Ti6Al4V implant coated with PLGA/Collagen/HA
nanofibers at 500x
63
Table 3.5: Average number of cells adhered to the Ti samples
Sample
10 min
20 min
30 min
60 min
cpTi
44
108
132
188
cpTi coated with PLGA nanofibers
683
1230
2188
3420
2482
3120
4601
biomineralized 746
1668
3842
5880
biomineralized 1080
3260
5460
7512
cpTi coated with PLGA/Collagen 1025
nanofibers
cpTi
coated
with
PLGA nanofibers
cpTi
coated
with
PLGA/Collagen nanofibers
Ti alloy
32
48
85
105
Ti alloy coated with PLGA nanofibers
592
1185
2018
3220
2380
3175
4500
1775
3724
5800
3120
5516
7475
Ti alloy coated with PLGA/Collagen 1105
nanofibers
Ti alloy coated with biomineralized 820
PLGA nanofibers
Ti alloy coated with biomineralized 1185
PLGA/Collagen nanofibers
64
Ti
P
Ti
P
er
s
er
s
/H
A
LG
A/
HA
lag
en
ol
Ti
P
ib
fib
no
f
na
no
na
LG
A/
C
lag
en
LG
A
nt
re
at
ed
100
ol
LG
A/
C
Ti
P
Ti
u
MSC capture percentage (%)
Ti
P
Ti
P
lag
en
ol
er
s
er
s
/H
A
LG
A/
HA
lag
en
Ti
P
ib
no
f
fib
nt
re
at
ed
na
no
na
LG
A
LG
A/
C
LG
A/
Co
l
Ti
P
Ti
u
MSC capture percentage (%)
100
Adhesion study on pure Ti
80
60
40
10 min
20
20 min
0
30 min
40
60 min
Adhesion study on Ti6Al4V
80
60
10 min
20
20 min
0
30 min
60 min
Figure 3.9: Percentage attachment efficiency of hMSCs on cpTi and Ti6Al4V alloy
65
3.4 Conclusion
Our work has proved the feasibility of creating a nanotextured surface on titanium by
simple acid/alkali treatment. The surface roughness can be tailored by modifying the
etching/ polishing procedures. Besides we have demonstrated that the cell adhesion can
be increased by coating the titanium surface with nanofibers. This is because the
nanofibers mimic the natural ECM and hence improve cell attachment. Through our
electrospinning set up we were able to achieve precise fiber deposition at a shorter
interval of time. We have increased the fiber deposition efficiency by our set up
compared to the conventional electrospinning. Moreover we have shown that the
adhesion efficiency of the hMSCs was the maximum on the cpTi and Ti alloy samples
coated with biomineralized PLGA/Collagen nanofibers compared to the other samples,
owing to the synergistic effect of collagen and n-HA.
66
Chapter 4
Mesenchymal stem cells proliferation and differentiation studies on the modified
implant surfaces
4.1 Introduction
When a biomaterial is implanted into the human body, it is unavoidable that blood will
contact the implant surface. Upon contact, the implant surface could be covered with a
layer of plasma proteins that mediate the next cellular responses. Therefore, the surface
characteristics of an implant should not only enhance the osteogenic cell–material
interactions but also optimize the initial blood–material interactions.
The success of a bone implant depends on how early the osseointegration is achieved
[73]. Hence the surface of the implants ought to be modified to improve early
osseointegration. Albrektsson et al., proposed six factors as especially important for
successful osseointegration. These include the implant material, implant design, surface
conditions, and status of the bone, the surgical technique and the implant loading
conditions
[74].
Nanofibers
have
demonstrated
excellent
cell
adhesion
and
differentiation. Ultimately, one of the main goals is to attract and induce the
osteoprogenitor cells to differentiate into osteoblasts. We hypothesize that coating
surfaces with nanofibers and the presence of biomolecules like n-HA would affect the
proliferation and differentiation of osteoprogenitors to osteoblasts. Therefore,
nanofibrous modification of dental and bone implants might enhance osseointegration.
There have been various techniques tried out in the past to improve the surface roughness
67
of the implant like plasma treatment, acid-etching and heat treatment. It is our hypothesis
that biomimetic bone like composite-coated metallic implants with loading capability
from the metal core and having a bioactive surface like nanofibers with nano-HA will
accelerate bone formation and implant fixation.
4.2 Material and methods
4.2.1 Mesenchymal stem cells culture
All the titanium samples were sterilized under UV light for 2 hrs. The discs were then
washed with phosphate buffered saline (PBS) thrice for 15min and were eventually
incubated with DMEM overnight before cell seeding. Human Mesenchymal stem cells
(PT-2501, Lonza, USA) was cultured in DMEM low glucose medium (DMEM,
Invitrogen, CA, U.S.) with 10% FBS (Invitrogen, CA, U.S.) and 1X Antibiotics (SigmaAldrich Chemical Company Inc., St. Louis, U.S) until 3 passages. The cells were then
trypsinized from the 75 cm2 cell culture flasks by adding 1 ml of 0.25% trypsin
containing 0.1% EDTA, purchased from GIBCO Invitrogen, USA. Detached cells were
centrifuged, counted by tryphan blue assay using a hemocytometer and seeded on the
scaffolds at a cell density of 1.0 × 104 cells/well for 24 well plates (pure Ti based
samples) and 2.0 × 104 cells/well for the 6 well plates (Ti alloy based samples) was added
and left in incubator for facilitating cell growth. The well plates were incubated for 60
minutes at room temperature to favor cell adhesion as described in chapter 3. After the
incubation time, the media was removed and the plates were washed thrice with PBS to
remove the unbound cells. Fresh media was then added to the wells and the plates were
then transferred to the incubator. The well plates were cultured for days 7, 14 and 21 to
68
carry out further cell culture analysis like proliferation, differentiation and mineralization.
The pure Ti samples – untreated Ti, treated Ti coated with PLGA nanofibers, treated Ti
coated with PLGA/Collagen nanofibers, treated Ti coated with PLGA nanofibers+nanoHA and treated Ti coated with PLGA/Collagen nanofibers+nano-HA, were cultured onto
the 24 well plates. The Ti-6Al-4V samples – untreated Ti-6Al-4V, treated Ti-6Al-4V
coated with PLGA nanofibers, treated Ti-6A-4V coated with PLGA/Collagen nanofibers,
treated Ti-6Al-4V coated with PLGA nanofibers+nano-HA and treated Ti-6Al-4V coated
with PLGA/Collagen nanofibers+nano-HA, were cultured onto 6 well plates. The cells
were cultured and analyzed for their proliferation and differentiation on days 7, 14 and
21. The optical images of hMSCs cultured on Tissue Culture Plate (TCP) at 24, 48, 72
and 96 hrs were shown in Appendix A
4.2.2 Cell Morphology Study
The cell morphology was analyzed using FESEM. After 6 days of seeding the hMSCs,
the media was removed from the wells and the samples were fixed with 3%
glutaraldehyde for 3 hours. The scaffolds were then rinsed with distilled water for 15min
and then dehydrated with a series of ethanol gradients starting from 30% to 50%, 75%,
90% and 100% (v/v). Subsequently the samples were treated with HMDS
(Hexamethyldisilazane) solution and allowed to air – dry at room temperature in the fume
hood. The samples were then gold coated and the cells morphology was analyzed using
FESEM. The mineral secreted by the cells was analyzed using FESEM equipped with
69
EDX. The same procedure was repeated for day 14 and day 21. The results are shown in
Appendix B.
4.2.3 Cell Proliferation Study
The cell proliferation on different scaffolds was analyzed using MTS assay (CellTiter 96
AQueous One solution reagent, purchased from Promega, Madison, WI). The principle
behind the MTS assay
involves the reduction of yellow tetrazolium salt [3-(4, 5-
dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2(4-sulfophenyl)-2H-tetrazolium] in
MTS to form purple formazan crystals by the dehydrogenase enzymes secreted by
mitochondria of metabolically active cells forms the basis of this assay. The formazan
dye shows absorbance at 492 nm and the amount of formazan crystals formed is directly
proportional to the number of cells. After 6 days of seeding, the media was removed from
the well plates and the scaffolds were washed in PBS. The scaffolds were then incubated
in a 1:5 ratio mixture of MTS assay and serum free DMEM medium for 3 – 5 hrs at 37°C
in a 5% CO2 incubator. After the incubation period, the samples were pipetted out into 96
well plates. The absorbance was then calibrated at 490 nm using a spectrophotometric
plate reader (Fluostar Optima, BMG Lab Technologies, Germany). The same procedure
was repeated for day 14 and day 21 samples.
70
4.2.4 Alkaline phosphatase activity
The osteogenic differentiation of hMSCs was analyzed by measuring the alkaline
phosphatase (ALP) activity. ALP activity was measured using Alkaline Phosphate
Yellow Liquid substrate system for ELISA (Sigma Life Sciences, USA). In this reaction,
ALP catalyzes the hydrolysis of colorless organic phosphate ester substrate, pnitrophenylphosphate (pNPP) to a yellow product, p-nitrophenol, and phosphate. After 6
days of seeding the scaffolds with the hMSCs were washed thrice with PBS. 400 μl of
pNPP liquid was added to the scaffolds and incubated for 30 min till the colour of
solution becomes yellow. The reaction was then arrested by the addition of 100 μl of 2 N
NaOH solution following which the yellow color product was aliquoted in 96-well plate
and read in spectrophotometric plate reader at 405 nm.
4.2.5 Cell mineralization Study
The amount of minerals secreted by the hMSCs can be both qualitatively and
quantitatively analysed by using Alizarin red staining. Alizarin Red-S (ARS) is a dye that
binds selectively to the calcium salts and hence can be used for mineral staining. The Ti
and Ti alloy scaffolds with hMSC cells was washed thrice with PBS and fixed in ice-cold
70% ethanol for 1 h. These constructs were then washed twice with distilled water and
stained with ARS (40 mM) for 20 min at room temperature. After several washes with
distilled water, the scaffolds were observed under inverted optical microscope and images
were taken using image software (Leica FW4000, version v 1.0.2). The stain was eluted
by incubating the scaffold with 10% cetylpyridinium chloride for 1 h. The absorbance of
71
the collected dye was then read at 540 nm in spectrophotometer (Thermo Spectronics,
Waltham, MA, USA).
4.2.6 Statistical analysis
Values (at least triplicate) were averaged and expressed as means ± standard deviation
(SD). Statistical differences were determined by Student two-sample t test. Differences
were considered statistically significant at p ≤ 0.05.
4.3 Results and Discussion
4.3.1 Cell Morphology Study
The hMSC morphology was analyzed on the day 7, 14 and 21 using FESEM. The Figures
4.1 – 4.3 show the SEM micrographs of cell interaction with the nanofibers as well as the
Ti scaffold. Since the duration of electrospinning was short, only a thin layer of the
electrospun fibers have been deposited onto the Ti plates. Hence the cells begin to
migrate further beyond the fibers and interact with the Ti discs. Nanotopography favors
cell adhesion, proliferation and differentiation [1, 4, 63, 66, 86]. Since both cp Ti and the
Ti6Al4V alloy have a nanotextured surface it is believed to enhance the cell – scaffold
interactions. From the cell morphology as shown in Figure 4.1, it is seen that by day 7 the
hMSCs cultured on the treated Ti coated with functionalized PLGA/Collagen nanofibers
with nano-HA have extended their filopodia and contacted the adjoining cells and
proliferated. The morphology of the cells remains rounded in the case of the untreated
72
scaffolds, indicating that the surface is not suitable for cell culture. Since the duration of
electrospinning was short, only a thin layer of the electrospun fibers have been deposited
onto the Ti plates. Hence the cells begin to migrate further beyond the fibers and interact
with the Ti discs. The nanotopography of a scaffold surface favors cell adhesion,
proliferation and differentiation. The cell spreading, with spindle-like and polygonal like
cell shapes, was also observed on HA-based composites on days 14 and 21 (Figure 4.2
and 4.3) of culture and physical contact between cells were maintained via filopodia or
lamellipodia [75]. By virtue of these observations, n-HA or n-HA in combination with
collagen would result in greater cell motility due to better-developed filopodia and
lamellipodia, as reported earlier [76]. Vanessa et al. [77] concluded that the treatment of
titanium and titanium alloy implant surfaces with discrete crystalline deposits like HA
renders them bone bonding, and it is the increase in complexity of the resultant surface
which is the driving force for the bonding mechanism at the bone – implant interface.
Theoretically, the osteoconductive properties of HA would provide reproducible
attachment of implants to the skeleton by osseointegration and bone ingrowth. The
morphology of the cells cultured on cp Ti and Ti-6Al-4V alloy was similar as the
scaffolds were subjected to the same surface treatment procedures.
73
a)
c)
b)
d)
e)
Figure 4.1 SEM images of the hMSC morphology on day 7 on a) untreated Ti-6Al-4V, b) Treated
Ti-6Al-4V coated with PLGA nanofibers, c) Treated Ti-6Al-4V coated with PLGA/Collagen
nanofibers, d) Treated Ti-6Al-4V coated with functionalized PLGA nanofibers, e) Treated Ti6Al-4V coated with functionalized PLGA/Collagen nanofibers.
74
a)
b)
c)
d)
e)
Figure 4.2 SEM images of the hMSC morphology on day 14 on a) untreated Ti-6Al-4V, b)
Treated Ti-6Al-4V coated with PLGA nanofibers, c) Treated Ti-6Al-4V coated with
PLGA/Collagen nanofibers, d) Treated Ti-6Al-4V coated with functionalized PLGA nanofibers,
e) Treated Ti-6Al-4V coated with functionalized PLGA/Collagen nanofibers.
75
a)
b
a)
)
c)
d)
e)
Figure 4.3 SEM images of the hMSC morphology on day 21 on a) untreated Ti -6Al-4V, b)
Treated Ti-6Al-4V coated with PLGA nanofibers, c) Treated Ti-6Al-4V coated with
PLGA/Collagen nanofibers, d) Treated Ti-6Al-4V coated with functionalized PLGA nanofibers,
e) Treated Ti-6Al-4V coated with functionalized PLGA/Collagen nanofibers.
76
4.3.2 Cell Proliferation Study
The MTS assay was used to study the cell proliferation of hMSCs on cpTi and Ti alloy
surfaces as shown in Figures 4.4a and 4.4b respectively. It was seen that even though the
cell proliferation rate initially was high in the TCP (Tissue Culture Plate), by day 21 the
cell proliferation was maximum in the functionalized nanofiber coated Ti. By day 21 the
cells seeded onto the biomineralized PLGA/Collagen scaffolds had proliferated by 257%
compared to day 7. In the case of cpTi samples on day 7, significant difference was
observed between the mineralized PLGA/Collagen nanofibers and the untreated Ti
samples. The samples coated with PLGA nanofibers were statistically different from the
sample coated with PLGA/Collagen nanofibers and the mineralized scaffolds; indicating
that PLGA alone was not sufficient to significantly enhance the cell proliferation rate.
Additionally, for the HA coated PLGA/Collagen nanofibers, the cell proliferation was
higher compared to the HA coated PLGA nanofibers. This was because of the presence
of collagen, which is a principle component of the ECM. The rate of cell proliferation
was rather slow in the untreated Ti samples. This was because the samples have a low
cell capture ratio as seen in the cell adhesion study in chapter 3. The proliferation in the
case of Ti coated with the PLGA and PLGA/Collagen nanofibers was also high as the
nanofibers mimic the ECM and thereby enhance the cell proliferation rate. However
owing to the presence of collagen the proliferation rate was higher in the PLGA/Collagen
nanofibers compared to the PLGA nanofibers. Thus it was seen that functionalization of
the nanofibers enhance the cell proliferation compared to the non-functionalized
scaffolds. As shown in the Figures 4.4a and 4.4b, the results for both the cp Ti and
Ti6Al4V alloy were similar. This maybe because the surface treatment and coating on
77
both the substrates were similar, indicating that both cpTi and Ti6Al4V alloy are
imparting similar mechanical and biological cues.
a)
*
*
b)
*
Figure 4.4 MTS assay for hMSC cells proliferation on a) cpTi based scaffolds – untreated, coated
with PLGA nanofibers, coated with PLGA/Collagen nanofibers, coated with PLGA/HA and
coated with PLGA/Collagen/HA nanofibers b) Ti-6Al-4V based scaffolds - untreated, coated with
PLGA nanofibers, coated with PLGA/Collagen nanofibers, coated with PLGA/HA and coated
with PLGA/Collagen/HA nanofibers for day 7, 14 and 21. * represents p≤ 0.05 statistical
difference. Control refers to the Tissue Culture Plate (TCP); TiK refers to Ti6Al4V alloy.
78
4.3.3 Alkaline phosphatase activity
Alkaline phosphatase is a membrane bound enzyme and its activity is used as an
osteoblastic differentiation marker [78], as it is produced only by cells showing
mineralized ECM [79]. The ALP activity indicates the osteogenic differentiation capacity
of the cells. The ALP activity for the cpTi and Ti alloy samples are depicted in Figures
4.5a and 4.5b. It was seen that even though initially at day 7 the ALP activity was similar
on all the scaffolds, the activity started to increase from day 14. There was not much
increase in the ALP activity for the untreated scaffolds, indicating that the cells cultured
on the untreated surfaces have not undergone osteogenic differentiation. This maybe
because the untreated scaffold have no biological and mechanical cues capable of
inducing the osteogenic differentiation of hMSCs. However TCP showed more ALP
activity compared to the untreated Ti surfaces, indicating that the untreated surfaces were
not suitable for cell differentiation compared to TCP. Comparison of the nonfunctionalized scaffolds show the ALP activity was higher for Ti coated with
PLGA/Collagen compared to the scaffolds coated with PLGA nanofibers (p ≤ 0.05).
There was also statistically significant difference between the PLGA/HA scaffolds and
the PLGA/Collagen/HA scaffolds indicating an enhanced ALP activity in the presence of
both collagen and HA. By day 21, significant difference was observed between the
mineralized scaffolds and the untreated and the non – mineralized scaffolds. As explained
earlier this increase in the ALP activity upon the incorporation of collagen was due to its
presence. Type I Collagen being the principle component of the organic part of the bone
matrix, induces the hMSCs to differentiate into bone cells. Thus the functionalized
nanofibers show the ability to induce osteogenic differentiation of mesenchymal stem
79
cells. However the hMSC undergoing osteogenic differentiation was more in the
biomineralized PLGA/Collagen fibers rather than the biomineralized PLGA nanofibers.
This difference is due to the presence of collagen in addition to HA which synergistically
enhances the osteogenic differentiation capacity of the hMSCs. However in the case of
cpTi, no significant difference was observed between the mineralized scaffolds and the
non – mineralized scaffolds on the days 7 and 14. This maybe because cpTi surface,
unlike the Ti alloy, is significantly slow in inducing the hMSCs differentiation. By day 21
however significant difference was observed between the mineralized and non –
mineralized scaffolds, indicating that collagen and HA being the native constituent of the
bone, synergistically induces hMSC differentiation. Ohgushi et al. [80] suggested that
mesenchymal cells could be influenced to differentiate into osteoblasts in the presence of
bioceramics. Our study has shown that the presence of bioceramics like HA has triggered
the differentiation of hMSC into osteoblasts as proved by the enhanced ALP activity in
the mineralized implants. As shown in the Figure 4.5a and 4.5b, the results for both cpTi
and Ti-6Al-4V alloy were similar. This similarity is due to the surface treatment
employed and the nanofibrous coating. Similar results have been found by others [81 84], who concluded that both rough Ti and Ti-6Al-4V surfaces enhance alkaline
phosphatase activity and mineralization. But none of the above studies involved coating
the rough Ti and Ti alloy discs with a mineralized nanofibrous coating. It has been shown
in vivo, that there are many cells which are capable of differentiating into osteoblastic
cells and hence contributing to the production of extracellular matrix, and that
differentiation is encouraged most by the HA coating on the titanium [85].
80
a)
*
*
b)
*
*
Figure 4.5 ALP activity for hMSC cells on a) cpTi based scaffolds - untreated, coated with PLGA
nanofibers, coated with PLGA/Collagen nanofibers, coated with PLGA/HA and coated with
PLGA/Collagen/HA nanofibers b) Ti-6Al-4V based scaffolds - untreated, coated with PLGA
nanofibers, coated with PLGA/Collagen nanofibers, coated with PLGA/HA and coated with
PLGA/Collagen/HA nanofibers for day 7, 14 and 21. * represents p≤ 0.05 statistical difference.
Control refers to the Tissue Culture Plate (TCP); TiK refers to Ti6Al4V alloy.
81
4.3.4 Cell mineralization study
Mineralization refers to the cell-mediated deposition of extracellular calcium and
phosphorus salts where anionic matrix molecules take up the calcium ions and the
phosphate ions and serve as nucleation and growth sites leading to calcification [86]. The
cell mineralization on the cpTi and Ti alloy samples have been represented quantitatively
in Figures 4.6a and 4.6b respectively, and qualitatively in Figures 4.7(A-C) and 4.8(A-C)
respectively. From Figures 4.6a and 4.6b, it was seen that the cells cultured on the
functionalized nanofiber scaffolds start secreting their minerals by day 7. In the case of Ti
alloy, even as early as day 7 significant difference (p ≤ 0.05) was observed between the
mineralized scaffolds and the non – mineralized scaffolds. This suggests that the HA
deposited on the fibers stimulates the mineralization of the cells. HA serves as biological
cues for stimulating the osteogenic differentiation and mineralization of hMSCs. The
untreated cpTi and Ti alloy samples (Figure 4.7A (a) and 4.8A (a)) showed similar
limited mineralization profile. Since it has not taken up the Alizarin red stain, it appears
green under the optical microscope. They did not favor cell mineralization owing to the
absence of any biological cues. However, the Ti coated with PLGA/Collagen favored
more cell mineralization than Ti coated with PLGA nanofibers. The presence of collagen
stimulates the secretion of minerals by the hMSCs. The cell mineralization was the
maximum for the functionalized PLGA/Collagen nanofibers. The presence of collagen
along with HA induces a synergic effect for the mineral secretion. In accordance to the
cell proliferation and the differentiation studies, the results were similar for both cp Ti
and Ti6Al4V alloy. Harris et al., demonstrated an increase in the extracellular matrix
production by osteoblastic cells when cultured on HA coatings [87]. Rough surface may
82
allow the osteoblastic cells to obtain more points of adhesion, as described by Niederauer
et al., [88], and to produce more extracellular matrix [89]. Although similar results have
not been reported with regard to hMSC, Tenenbaum et al., [90] have shown that, given
the right environment, many cells are capable of behaving in an osteoblast-like way and it
may be that the correct microenvironment is provided by the HA and collagen present in
the nanofiber surface coating. From the qualitative representation of the ARS staining as
shown in Figures 4.7(A-C) and 4.8(A-C), it was seen that the ARS staining was more
preponderant on the functionalized PLGA/Collagen nanofibers coated scaffolds
compared to the other scaffolds. This was because the cell mineralization was more on
the functionalized nanofibers and hence they take up more alizarin red stain giving a
bright red appearance. The untreated Ti scaffolds had not taken up any stain due to the
absence of the cell mineralization. Hence it gives a green appearance on the optical
microscope. Moreover, the ARS staining was increased by day 21 compared to the day 7
and day 14 on cpTi and Ti alloy samples coated with nanofibers indicating that the cell
mineralization has increased. More alizarin red staining uptake can be noticed on the
samples coated with mineralized PLGA/Collagen nanofibers on the cpTi and Ti alloy
samples as indicated in Figure 4.7C and 4.8C compared to the day 7 (Figure 4.7A and
4.8A) and day 14 (Figure 4.7B and 4.8B).
The mineralization was further analyzed using FESEM EDX (Supplementary data). The
Ca/P ratio of mineral was 1.3 on day 7 on the functionalized PLGA/Collagen nanofibers
with nano-HA. This ratio increased further as the cell mineralization increases. It was
seen that the Ca/P ratio on day 14 and day 21 was 1.5 and 2.03 respectively. The EDX
data were attached in the Appendix B.
83
a)
*
*
b)
*
*
Figure 4.6 Quantitative data for Alizarin red staining on hMSC cells on a) cp Ti scaffolds b) Ti6Al-4V scaffolds for days 7, 14 and 21. * represents p≤ 0.05 statistical difference
84
a)
b)
c)
d)
e)
Figure 4.7A: Optical image of the ARS stained hMSCs on the cpTi scaffolds on day 7 a)
untreated cpTi, b) Treated Ti coated with PLGA nanofibers, c) Treated cpTi coated with
PLGA/Collagen nanofibers, d) Treated cpTi coated with functionalized PLGA nanofibers, e)
Treated cpTi coated with functionalized PLGA/Collagen nanofibers.
85
a)
b)
c)
d)
e)
Figure 4.7B Optical image of the ARS stained hMSCs on the cpTi scaffolds on day 14 a)
untreated cpTi, b) Treated cpTi coated with PLGA nanofibers, c) Treated cpTi coated with
PLGA/Collagen nanofibers, d) Treated cpTi coated with functionalized PLGA nanofibers, e)
Treated cpTi coated with functionalized PLGA/Collagen nanofibers.
86
a)
c)
b)
d)
e)
Figure 4.7C Optical image of the ARS stained hMSCs on the cpTi scaffolds on day 21 a)
untreated cpTi, b) Treated cpTi coated with PLGA nanofibers, c) Treated cpTi coated with
PLGA/Collagen nanofibers, d) Treated cpTi coated with functionalized PLGA nanofibers, e)
Treated cpTi coated with functionalized PLGA/Collagen nanofibers.
87
a)
b)
c)
d)
e)
Figure 4.8A Optical image of the ARS stained hMSCs on the Ti-6Al-4V scaffolds on day 7 a)
untreated Ti-6Al-4V, b) Treated Ti-6Al-4V coated with PLGA nanofibers, c) Treated Ti-6Al-4V
coated with PLGA/Collagen nanofibers, d) Treated Ti-6Al-4V coated with functionalized PLGA
nanofibers, e) Treated Ti-6Al-4V coated with functionalized PLGA/Collagen nanofibers.
88
a)
c)
b)
d)
e)
Figure 4.8B Optical image of the ARS stained hMSCs on the Ti-6Al-4V scaffolds on day 14 a)
untreated Ti-6Al-4V, b) Treated Ti-6Al-4V coated with PLGA nanofibers, c) Treated Ti-6Al-4V
coated with PLGA/Collagen nanofibers, d) Treated Ti-6Al-4V coated with functionalized PLGA
nanofibers, e) Treated Ti-6Al-4V coated with functionalized PLGA/Collagen nanofibers.
89
a)
b)
c)
d)
e)
Figure 4.8C Optical image of the ARS stained hMSCs on the Ti-6Al-4V scaffolds on day 21 a)
untreated Ti-6Al-4V, b) Treated Ti-6Al-4V coated with PLGA nanofibers, c) Treated Ti-6Al-4V
coated with PLGA/Collagen nanofibers, d) Treated Ti-6Al-4V coated with functionalized PLGA
nanofibers, e) Treated Ti-6Al-4V coated with functionalized PLGA/Collagen nanofibers.
90
4.4 Conclusion
The studies reported so far [64, 91, 92] have used various cell culture models
representing osteoblasts and various stages in their lineage. The present study used
human mesenchymal stem cells, which are capable of differentiating into osteoprogenitor
cells and osteoblasts. However the increase in proliferation rate in the biomineralized
samples, which also showed an enhanced mineralization, is in contrary to the theory that
proliferation is down-regulated when extracellular matrix maturation is induced and
mineralization occurs [93]. This absence in down-regulation of the hMSCs cell number
on the biomineralized scaffolds which also showed enhanced mineralization maybe
because of the presence of surface cues like HA present on the implants which triggers
proliferation as well as mineralization simultaneously, without inhibiting one of them.
Similar results were obtained by Deepika et al., [66] who cultured osteoblasts on the HA
sprayed and HA blended PCL/Gelatin nanofibers; both proliferation and mineralization
increased continuously and was maximum on the HA sprayed nanofibrous scaffold. Cell
behaviour such as adhesion, spreading and proliferation represent the initial phase of
cell–scaffold interaction that subsequently effect differentiation and mineralization [86].
Animal studies done in the past have suggested that HA stimulates bone to bridge gaps,
induces fibrous connective tissue (FCT) metaplasia to bone, and increases bone
mineralization when compared to uncoated implants of equal size, material, and structure
[94]. This is in correlation to the increased mineralization observed on the HA coated Ti
samples compared to the untreated Ti samples, in our study. Although in the earlier
studies the deposition of HA was by plasma spray technique, the process has
disadvantages attributed to the high temperatures used during the process, such as the
91
possibility of fracture at the interface between the titanium and the HA due to the residual
stress at the interface, and changes in the composition, porosity, crystallinity, and
structure of the plasma sprayed HA [87]. Our study demonstrated the deposition of HA
on to the nanofibers by a feasible Ca–P dipping method, thereby overcoming the
disadvantages of the plasma-spraying technique, in addition to increasing the bonebonding ability. The difference in the bone reaction between HA coated and uncoated
implants, as reported by Alzubaydi et al., [95] not only suggests a high osteoconductive
potential of the coated HA material but also its osteoinductivity, which is very much
essential for early osseointegration. However most techniques used to deposit inorganic
Ca-P coatings involve either extremely high temperatures or other non-physiological
conditions that impede the incorporation of biomolecules such as collagen [96- 99].
Hence our method by which collagen is electrospun along with PLGA followed by the
dipping method to deposit Ca-P is advantageous. This in-situ method of producing n-HA
on polymeric nanofiberous scaffolds coated on nanotextured implant surface, may be a
probable option for future implant materials.
92
Chapter 5
Conclusions and Recommendations
5.1 Conclusions
This work has proved the feasibility of creating a nanotextured surface on titanium by
simple acid/alkali treatment. The surface roughness can be tailored by modifying the
etching/ polishing procedures. Another strategy to enhance osteoconduction is the use of
a nanofiber-coated surface. Besides we have demonstrated that the cell adhesion can be
increased by coating the titanium surface with nanofibers. This is because the nanofibers
mimic the natural ECM and hence improve cell attachment. The cell attachment can be
increased further by depositing n-HA onto the nanofibers. The n-HA biomolecule
improves the adhesion efficiency of the hMSCs compared to the non-biomineralized
nanofibers. Through our electrospinning set up we were able to achieve fiber deposition
at a shorter interval of time. We have increased the fiber deposition efficiency by our
experimental set up compared to the conventional electrospinning.
A combination of structural, mechanical and biological properties of an implant material
play a critical role in cell seeding, proliferation and new tissue formation in orthopedic
research. Nano-biomaterials should promote cell adhesion and be optimized for ECM
production, mineralization and subsequent tissue regeneration. PLGA/Collagen/HA
nanofibers coated implant surfaces fabricated by a modified advanced electrospinning
technique, and hMSCs grown on them showed higher cell proliferation, and increased
93
ALP activity and mineralization, compared to the PLGA, PLGA/HA and PLGA/Collagen
nanofiber coated implant surfaces. Hence, electrospun biomimetic PLGA/Collagen/HA
nanofibers coated Ti surfaces hold great potential for adhesion, proliferation,
differentiation and mineralization of hMSCs.Our results suggested that the nanotextured
oxidised titanium surfaces, both cpTi and Ti alloy, coated with biomineralized
PLGA/Collagen nanofibers enhanced the initial cell capture ratio. Incorporating
biomolecular cue like collagen and n-HA have enhanced the cell proliferation, osteogenic
differentiation and cell mineralization. Thus the healing time can be reduced, leading to
enhanced initial osseointegration. To our knowledge, functionalized nanofiber treated
dental implant is a novel idea for enhanced osseointegration using bone regeneration
concept. The complete bone integration between dental implant and host bone will be
enhanced by following three biomimetic aspects:
1. Natural ECM like nanofiber coated on the dental implant;
2. The biomineralization treatment;
3. Controllable MSCs incorporation with dental implant.
5.2 Recommendations
From the above results it can be seen that osseointegration of the implant can be
improved by coating the implant surface with biomineralized PLGA/Collagen nanofibers.
This is very useful to reduce the duration required for osseointegration as the
biomineralized scaffolds have shown tremendous promise for early cell capture. This
enhances the bone bonding ability of the implant in vivo, resulting in early
94
osseointegration. For sure, animal study in a rabbit model will prove this concept
ultimately.
On the other hand, growth factors like bone morphogenic proteins (BMPs), especially
BMP-7 can be incorporated into the nanofibers and the increase in the early
osseointegration will be possible.
95
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109
Appendix A
Figure: MSC proliferation on TCP at 24, 48, 72 and 96 hrs.
110
Appendix B
FESEM-EDX results on Ti alloy coated with PLGA/Collagen/HA on day 7, 14 and 21.
111
1/1
View000
--------------------------Title
: IMG1
--------------------------Instrument
: 6700F
Volt
: 15.00 kV
Mag
: x 3,000
Date
: 2009/06/16
Pixel
: 512 x 384
---------------------------
10 µm
001
Acquisition Parameter
Instrument
: 6700F
Acc. Voltage : 15.0 kV
Probe Current: 1.00000 nA
PHA mode
: T3
Real Time
: 60.87 sec
Live Time
: 50.00 sec
Dead Time
: 18 %
Counting Rate: 3595 cps
Energy Range : 0 - 20 keV
001
4400
4000
3600
3200
2400
CaKa
PKa
Counts
2800
2000
CaKb
1600
1200
800
400
0
0.00
1.00
2.00
3.00
4.00
5.00
6.00
7.00
8.00
9.00
10.00
keV
ZAF Method Standardless Quantitative Analysis
Fitting Coefficient : 0.7151
Element
(keV)
mass% Error%
At% Compound
P K
2.013
32.94
1.62
38.86
Ca K
3.690
67.06
3.41
61.14
Total
100.00
100.00
AnalysisStation
mass%
Cation
K
33.0606
69.2746
1/1
View001
--------------------------Title
: IMG1
--------------------------Instrument
: 6700F
Volt
: 15.00 kV
Mag
: x 1,600
Date
: 2009/06/16
Pixel
: 512 x 384
---------------------------
20µm
001
Acquisition Parameter
Instrument
: 6700F
Acc. Voltage : 15.0 kV
Probe Current: 1.00000 nA
PHA mode
: T3
Real Time
: 60.51 sec
Live Time
: 50.00 sec
Dead Time
: 17 %
Counting Rate: 3487 cps
Energy Range : 0 - 20 keV
0.00
1.00
2.00
CaKa
CaKb
PKa
001
3.00
4.00
5.00
6.00
7.00
8.00
9.00
10.00
keV
ZAF Method Standardless Quantitative Analysis
Fitting Coefficient : 0.9557
Element
(keV)
mass% Error%
At% Compound
P K*
2.013
40.19
12.62
46.51
Ca K*
3.690
59.81
27.19
53.49
Total
100.00
100.00
AnalysisStation
mass%
Cation
K
41.1509
61.4710
1/1
View000
--------------------------Title
: IMG1
--------------------------Instrument
: 6700F
Volt
: 15.00 kV
Mag
: x 3,000
Date
: 2009/06/16
Pixel
: 512 x 384
---------------------------
10 µm
001
Acquisition Parameter
Instrument
: 6700F
Acc. Voltage : 15.0 kV
Probe Current: 1.00000 nA
PHA mode
: T3
Real Time
: 57.00 sec
Live Time
: 50.00 sec
Dead Time
: 12 %
Counting Rate: 2651 cps
Energy Range : 0 - 20 keV
001
6000
5500
5000
4500
Counts
4000
3500
3000
2500
1500
CaKa
CaKb
PKa
2000
1000
500
0
0.00
1.00
2.00
3.00
4.00
5.00
6.00
7.00
8.00
9.00
10.00
keV
ZAF Method Standardless Quantitative Analysis
Fitting Coefficient : 0.9655
Element
(keV)
mass% Error%
At% Compound
P K*
2.013
43.16
16.72
49.56
Ca K*
3.690
56.84
36.39
50.44
Total
100.00
100.00
AnalysisStation
mass%
Cation
K
44.5089
58.2319
[...]... knowledge, dental implant using functionalized nanofibers as a surface modification is a novel idea to enhance osseointegration using the bone regeneration concept 15 Chapter 1 Introduction 1.1 Background In the past 20 years, the number of dental implant procedures has increased steadily worldwide, reaching about one million dental implantations per year [1] Dental implants are useful for restoration of oral... period of 12 weeks [5] and this was reduced to 6 to 8 weeks with the introduction of the SLA (sand blasted, acid etched) surface [6] The differences in the contact angle and the surface roughness of the implant surface owing to the various surface modification techniques were shown in Table 2.1 Table 2.1 Different types of implant surface modifications and their surface roughness and contact angle Type of. .. short- and long-term success of the implants These parameters are associated with delicate surgical techniques, a prerequisite for a successful early clinical outcome High success rates for dental implants are reported in healthy patients with good bone quality In the future, with an aging population, more patients may be considered for dental implants; osseointegration of dental implants under less than... called the ―dip‖ Many implant failures occur during this period, and this period seems to be critical to the successful integration of the implant [2] 17 Figure 1: A typical Ti dental implant 18 1.3 Hypothesis and Objectives: Hypothesis This project is to develop a surface modification system for dental implant using electrospun nanofiber and biomineralization to fabricate a biomimetic substrate We... be used for the surface modification or activation of an implant surface Among these methods, chemical modifications seem to be relatively simple and inexpensive Hence it is widely used There have been various techniques tried out in the past to improve the surface roughness of the implant like plasma treatment, acid-etching and heat treatment For example, the TPS (titanium 21 plasma sprayed) surfaces... average 13 SUMMARY The introduction of dental implants has changed the way dentists approach the replacement of missing teeth The clinical success of dental implants is related to their osseointegration, which is a property virtually unique to titanium and has enhanced the science of joint replacement techniques Generally, the time between implant placement and implant loading ranged from 3 months... nanostructured surfaces of nanometallic and nanoceramic materials have several advantages compared to the conventional surfaces These include, (i) they possess greater surface roughness resulting from both decreased grain size and possibly decreased diameter of surface pores, (ii) enhanced surface wettability due to greater surface roughness and (iii) greater numbers of grain boundaries There are a number of physical... cases, enhanced bone formation around the implant would be an important criterion 16 This may be achieved by implant coatings that are able to interact actively with the surrounding tissues 1.2 Clinical problems associated with osseointegration: There are two types of responses exhibited by the body after implantation The first type involves the formation of a soft fibrous tissue around the implant. .. differentiation, which are crucial for enhanced osseointegration Objectives: Modify the implant surface to produce nanotextured topography Develop a nanofibrous coating from biodegradable synthetic polymers and/or natural polymers to mimic extracellular matrix Functionalization of the nanofiber by biomineralization Evaluate adhesion, proliferation of MSCs on the modified implant surface Investigate osteogenic... chemistry of the surfaces In all the cases TiO2 being the principle chemical component [16] The Ti metal spontaneously forms a protective TiO2 layer in the atmosphere When the Ti implant is inserted into the human body, the surrounding tissues directly contact the TiO2 layer on the implant surface The surface characteristics of the TiO2 layer determine the biocompatibility of Ti implant Therefore, it ... strength of the PLGA and PLGA/Collagen coated implant surface owing to the hard nature of the implant material and the thin layer of nanofiber coating on the implant surface The tensile properties of. .. integration of the implant [2] 17 Figure 1: A typical Ti dental implant 18 1.3 Hypothesis and Objectives: Hypothesis This project is to develop a surface modification system for dental implant using... roughness of the implant surface owing to the various surface modification techniques were shown in Table 2.1 Table 2.1 Different types of implant surface modifications and their surface roughness