Metal Oxide Nanostructures and their Applications
(ISBN 1-58883-170-1)
Edited by
Ahmad Umar, Najran University, Saudi Arabia
Yoon-Bong Hahn, Chonbuk National University, South Korea.
Published by
American Scientific Publishers, CA, USA, (http://www.aspbs.com/mona/)
2010
5-Volume Set, 4,000 pages, Hardcover
Volume 4 Applications (Part-2)
Chapter 12 Bioactive Bioceramic Coatings. Part I: Coatings on Non-metallic Biomaterials
(Vol.4, pp 451-478)
Proof-reading Corrections:
No. Original
1.
Reference [10]
2.
Reference [13]
3.
Reference [15]
4.
Reference [32]
Correction
10. L. L. Hench and J. Wilson, An Introduction to
Bioceramics, World Scientific Publishing, Singapore,
1993.
13. T. Nakamura, M. Neo, and T. Kokubo,
“Mineralization in Natural and Synthetic Biomaterials”
(P. J. Li, P. Calvert, T. Kokubo, R. Levy, and C. Scheid,
Eds.), Materials Research Society Symposium
Proceedings, Vol.559, p. 15, PA, 2000.
15. M. J. Read, S. L. Burkett, and A. M. Mayes,
“Mineralization in Natural and Synthetic Biomaterials”
(P. J. Li, P. Calvert, T. Kokubo, R. Levy, and C. Scheid,
Eds.), Materials Research Society Symposium
Proceedings, Vol.559, p. 337, PA, 2000.
32. M. Wang, L. J. Chen and J. Ni, Transactions of the
Society For Biomaterials 28th Annual Meeting, Tampa,
Florida, 2002, p. 259.
CHAPTER 12
Bioactive Bioceramic Coatings:
Part I. Coatings on Non-Metallic
Biomaterials
Jin-Ming Wu1 , Min Wang2
1
Department of Materials Science and Engineering, Zhejiang University,
Hangzhou 310027, P. R. China
2
Department of Mechanical Engineering, The University of Hong Kong,
Pokfulam Road, Hong Kong
CONTENTS
1.
2.
3.
4.
5.
Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1
Bioactive Bonding and In Vitro Bioactivity . . . . . . . . . . . . . . . 2
2.1. In Vivo Observations of Bioactive Bonding . . . . . . . . . . 2
2.2. In Vitro Evaluation of Bioactivity . . . . . . . . . . . . . . . . . 4
2.3. Correlation Between In Vitro Apatite Formation and
In Vivo Bone Bonding . . . . . . . . . . . . . . . . . . . . . . . . . . 8
Mechanisms of In Vitro Apatite Deposition . . . . . . . . . . . . . . 8
3.1. Silica Gel . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9
3.2. Titania Gel . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9
3.3. Calcium-Phosphate (Ca-P) Layer . . . . . . . . . . . . . . . . 11
3.4. Other Surface Functional Groups . . . . . . . . . . . . . . . . 12
3.5. Effects of SBF-Related Factors . . . . . . . . . . . . . . . . . . 12
Bioactive Coatings on Non-Metallic Biomaterial Substrates . 13
4.1. Bioceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13
4.2. Biomedical Polymers . . . . . . . . . . . . . . . . . . . . . . . . . 16
4.3. Biomedical Composites . . . . . . . . . . . . . . . . . . . . . . . 19
4.4. Scaffolds for Tissue Engineering . . . . . . . . . . . . . . . . 21
Concluding Remarks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 23
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 24
1. INTRODUCTION
The increasingly longer life expectancy of human beings requires more and more surgical operations for hard-tissue replacement [1]. A national survey (1991–2005) by the
ISBN: 1-58883-170-1
Copyright © 2010 by American Scientific Publishers
All rights of reproduction in any form reserved.
Metal Oxide Nanostructures and Their Applications
Edited by Ahmad Umar and Yoon-Bong Hahn
Volume 4: Pages 1–28
1
2
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
American Academy of Orthopedic Surgeons showed that 160,000 hip prostheses and
266,000 knee prostheses were implanted in the United States in 1998, and these figures
increased steadily to 235,000 and 534,000, respectively, in 2005. Transplantation, which
utilizes tissues from a human or other species donor, has been almost abandoned nowadays for hard tissue replacements. Man-made materials, including metals and alloys,
ceramics, polymers, and composites, have been widely used due to their availability,
reproducibility, and reliability to repair or replace the malfunctioning human tissues.
The tight and reliable fixation between the implants and surrounding tissues is always
favored in restorative surgery, especially in joint-replacement operations. For a total hip
replacement, poly(methyl methacrylate) (PMMA) bone cement is normally used to secure
hip prostheses. Unfortunately, there have been problems with this type of “mechanical fixation.” Severe abrasive wear, which results from the detachment of bone cement,
is usually reported clinically. In addition, it is not easy to retrieve the failed implants
when a revision surgery is required [2]. To overcome such problems, the so-called “biological fixation,” which relies on the in-growth of the bone into the porous surface or
textured surface of implants for stabilization, has been developed and clinically used.
The porous surface structure can be achieved by plasma spraying metal powders, sintering wire meshes, or sintering metal beads on the implants, and the pore size range can
be optimized in the range of 100∼350 m for the best bone in-growth in clinical practices [3, 4]. Although the strategy of using surface porous implants without bone cement
(uncemented implants) have proven to be successful, bone in-growth is seldom observed
[5]. Therefore, “bioactive fixation” has been advocated in the past several decades since
the milestone work by Hench et al. in 1971 [6].
Natural bone is a composite of collagen and mineral nanoparticles made of carbonated
hydroxyapatite (CHA). The basic building block of bone material is collagen fibril of
ca. 100 nm in diameter, which consists of an assembly of 300 nm long and 1.5 nm thick
collagen molecules, deposited by the osteoblasts (bone-forming cells) into the extracellular space and then self-assembled into fibrils. Collagen fibrils are filled and coated with
tiny mineral CHA crystals with sizes of ca. 1.5–4.5 nm [7]. Therefore, it is not surprising
that the bioactive fixation is defined as the chemical, direct bonding at the molecular level
between implants and their surrounding bone tissues through an apatite layer consisting
of nano-sized particles [8]. After implantation in the human body for 3∼6 months, the
interfacial strength achieved by bioactive fixation is comparable to or higher than the
strength of the surrounding bone tissue [2]. The key for implants to achieve bioactive
fixation when implanted in the human body is their high bioactivity, i.e., the ability to
attach to the surrounding bone tissue through chemical bonding in the human body
environment [9]. Therefore, over the past 20 or so years, apart from the development
of bioactive materials, such as bioactive glasses, ceramics, and glass-ceramics, attention
has also been paid to providing bioinert biomaterials with bioactivity through various
coating techniques.
The two chapters of this series focus on reviewing bioactive bioceramic coatings on
various biomaterial substrates. The in vivo and in vitro evaluation of the bioactivity, mechanisms for apatite formation, and various coating techniques are reviewed. The first
chapter deals with several common issues and also coatings on non-metallic biomaterials. Because of the significance and the relatively larger amount of published papers
concerning bioactive coatings on metallic biomaterial substrates, the second chapter concentrates on the work of coating metallic biomaterials, with special attention paid to
nanostructured titania coatings.
2. BIOACTIVE BONDING AND IN VITRO BIOACTIVITY
2.1. In Vivo Observations of Bioactive Bonding
Host tissues respond naturally to any material that invades the body. There are three
types of implant-tissue interactions for biomaterials, as listed in Table 1. In most cases,
the human body tends to protect itself spontaneously by forming a fibrous tissue to isolate the invading material. For a material with a controlled level of chemical reactivity,
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
3
Table 1. Implant-tissue interactions for biomaterials.
Implant-tissue reaction
Bioinert
Bioactive
Biodegradation
Consequence
Illustrational biomaterials
Tissue forms a non-adherent
fibrous capsule surrounding
the implant
Tissue forms an interfacial
bond with the implant
Tissue replaces the implant
Alumina, zirconia, metals, most polymers
Hydroxylapatite (HA), bioactive glasses,
bioactive glass-ceramics
Tricalcium phosphate, bioactive glasses
Source: Reprinted with permission from [10], L. L. Hench and J. Wilson, World Scientific Publishing, Singapore, 1, 1993.
© 1993.
through certain surface reactions a bioactive interface can be developed that prevents
motion and mimics the type of interface that is formed when natural tissues repair
themselves. A highly reactive material dissolves or resorbs gradually and is replaced
eventually by the newly formed tissues [10]. Biodegradable materials are the most ideal
materials to be used as hard tissue substitutes as nothing is left in the human body once
the diseased or fractured tissue has been repaired. However, the degradation of the materials surely causes a loss of the implant’s mechanical properties. It is not easy to reach a
compromise between the degradation (or disappearance) rate and the rate of decrease in
mechanical properties during the tissue-regeneration process. Therefore, stable bioactive
materials are still attracting considerable attention at the present time.
A cross-sectional scanning electron microscopy (SEM) micrograph published by
Fujibayashi et al. [11] clearly shows the typical bonding between bioactive materials and
bone (Fig. 1). After implantation for a certain duration, a silicon-rich layer formed on
the bioactive glass surface through certain reactions, on which apatite deposition was
triggered. The surrounding bone tissue attached to the implant through this apatite layer
[11, 12]. The formation of a thin apatite layer, which was characterized by transmission
electron microscopy (TEM) to be collagen-free fine nanocrystals having no preferred orientation [13], can be well reproduced in vitro using simulated body fluids (SBF) and
hence be well characterized (see the Mechanisms of In Vitro Apatite Deposition section
for details). However, the bone apposition process on the thin apatite layer is not yet
well clarified due to the complicated bone biology, bone cell mechanotransduction, and
cell-surface interactions involved in the new bone development process [14]. Six events
have been suggested to be involved in sequence: adsorption of biological molecules in
the apatite layer, action of macrophages, attachment of stem cells, differentiation of stem
Figure 1. In vivo observation of a bioactive bone bonding between a bioglass implant and its surrounding
bone tissue. Reprinted with permission from Ref. [11], S. Fujibayashi et al., Biomaterials 24, 1349 (2003). © 2003,
Elsevier Science Ltd.
4
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
cells, generation of a matrix, and mineralization of the matrix [12]. Many reports have
appeared giving observations on proteins adsorption [15, 16] and cells attachment and
differentiation [13, 17, 18] on apatite films. However, the latter stage is relatively less
well-defined, and the factors controlling the rate still remain unknown [12].
The formation of a thin bone-like apatite layer has been observed without exception
on all bioactive materials. For example, after eight weeks of implantation in rats, a thin
collagen-free apatite layer was observed on the surfaces of all bioactive glass-ceramics of
Ceravital® (Na2 O-K2 O-MgO-CaO-SiO2 -P3 O5 glass containing Ca10 (PO4 )O6 crystals), glassceramic A-W (Na2 O-K2 O-MgO-CaO-SiO2 -P3 O5 glass containing hydroxyapatite and wollastonite crystals), dense hydroxyapatite (HA), and Bioglass® [13].
2.2. In Vitro Evaluation of Bioactivity
2.2.1. Simulated Body Fluid and Thermodynamics of
In Vitro Apatite Formation
In vivo observations have confirmed that a common feature of all bioactive implants is
the formation of a carbonate incorporated hydroxylapatite (CHA) layer, which is equivalent in composition and structure to the mineral phase of bone, on their surfaces when
implanted [10]. In vivo experiments require sacrificing animals as well as conducting
surgical operations and waiting for long durations for revealing effects. To evaluate
the bioactivity of biomaterials simply and effectively, Kokubo et al. developed the now
widely used SBF, which is a metastable calcium phosphate (Ca-P) solution similar in inorganic ion concentrations to those of human blood plasma, as shown in Table 2 [19–21].
Occasionally, Hank’s balanced salt solution (HBSS) is also used to evaluate in vitro
bioactivity.
SBF is a highly supersaturated solution with respect to apatite. It is stable because
of the high nucleation energy. Once nucleated, apatite precipitates follow the reaction
below:
10Ca2+ + 6PO43− + 2OH− = Ca10 PO4 6 OH2
(1)
The driving force for the above reaction can be expressed as:
G = −RT lnS
(2)
where G is the Gibbs free energy of formation, R is the gas constant, T is the absolute
temperature, and S is supersaturation. S is calculated by
S = IP /Ksp 1/9
2+ 5
IP = Ca (3)
PO43− 3 OH− (4)
Ksp is the solubility of the apatite. At 37 C, Ksp = 235 × 10−59 [23].
The supersaturation corresponding to apatite in SBF is as high as 167 [19]. At the normal body temperature of 36.5 C, the free energy calculated using Eq. (2) is −13.2 KJ/mol.
Therefore, once nucleated, apatite grows spontaneously. The control step of apatite deposition in SBF is the nucleation process.
The nucleation rate of apatite is determined by
J = K exp−G∗ /kT (5)
Table 2. Ion concentrations (mM) of simulated body fluid (SBF) [19], Hank’s balanced salt solution (HBSS)
[21, 22] and human blood plasma (HBP).
SBF
HBSS
HBP
Na+
K+
Ca2+
Mg2+
Cl−
HCO−3
HPO2−
4
SO2−
4
142. 0
142. 0
142. 0
5.0
5.81
5.0
2.5
1.26
2.5
1.5
0.898
1.5
148.8
146.0
103.0
4.2
4.17
27.0
1.0
0.779
1.0
0.5
0.406
0.5
5
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
where k is the Boltzmann constant, K is a kinetic factor, G∗ is nucleation activation
energy. The value of G∗ for a spherical nuclei can be expressed by [23]
G∗ =
16 2 3 f 3k2 T 2 ln S2
(6)
where v is the molecular volume of apatite (26324 × 10−28 m3 ), is the interfacial tension
between apatite and the solution (10.4 mJ/m2 ), and f is a function of the contact
angle between the nucleus and substrate. In addition to the solution temperature, the
nucleation rate of apatite is determined mainly by the solution supersaturation (S) and
the contact angle (). A higher supersaturation of the solution and a better wetability
of the substrate contribute to a higher nucleation rate of apatite. In addition, defects on
the substrate surface effectively improve the nucleation rate by reducing the activation
energy.
There are several Ca-P compounds, shown in Table 3, that are of considerable interest
in the biomaterials field. A theoretic analysis of Ca-P precipitation in SBF indicates that
apatite precipitation exhibits a higher thermodynamic driving force than does OCP and
DCPD in SBF, but OCP precipitation is kinetically favorable in SBF. The nucleation rate
of apatite is significantly affected by the pH value. A high pH environment favors apatite
nucleation. Precipitation of CHA is more kinetically favorable than that of stoichiometric
HA and has the same level of thermodynamic driving force. The precipitation of calciumdeficient HA is also more kinetically favorable, but its thermodynamic driving force is
lower than that of stoichiometric HA [23].
Observation with a dynamic light-scattering photometer equipped with a high-power
Ar-ion laser revealed Ca-P clusters with an initial hydrodynamic diameter of about 1 nm
in the SBF [25]. The cluster remained almost unchanged for a prolonged time up to 7 d at
36.5 C. A modified SBF, which has totally equal concentrations of ions to those in blood
plasma, has also been recommended for in vitro bioactivity evaluation [26].
Juhasz et al. argued recently that human blood serum (HBS) served as well to mimic
the in vivo crystallization process on brushite, tricalcium phosphate (TCP) and HA.
By using SBF with the addition of HBS, the influence of proteins can be assessed,
which hence provides a more accurate in vitro model for the bone-bonding ability of
biomaterials [27].
2.2.2. Characterization of Apatite Formed In Vitro
The apatite precipitated from SBF is characteristic of bone mineral; i.e., small crystal size,
the presence of carbonate ions, ionic substitution by Mg2+ and Cl− , and calcium-deficient
[28]. A typical apatite coating formed in SBF on a chemically modified commercial pure
titanium (CPTi) substrate is shown in Figure 2. Each globule of the coating is actually
aggregates of numerous tiny flakes united together. Figure 3 exhibits the corresponding thin-film X-ray diffraction (TF-XRD) pattern of the coating. It gives an abnormally
Table 3. Calcium phosphate compounds.
Compound
Monocalcium phosphate monohydrate
Monocalcium phosphate anhydrous
Dicalcium phosphate dihydrate (brushite)
Dicalcium phosphate anhydrous
Octacalcium phosphate
-Tricalcium phosphate
-Tricalcium phosphate
Tetracalcium phosphate
Calcium hydroxyapatite
Fluoroapatite
Abbreviation
Formula
MCPM
MCPA
DCPD
DCPA
OCP
-TCP
-TCP
TTCP
HA
FAP
Ca(H2 PO4 )2 · H2 O
Ca(H2 PO4 )2
CaHPO4 · 2H2 O
CaHPO4
Ca8 H2 (PO4 )6 · 5H2 O
-Ca3 (PO4 2
-Ca3 (PO4 2
Ca4 (PO4 2 O
Ca10 (PO4 6 (OH)2
Ca5 (PO4 3 F
Source: Reprinted with permission from [24], L. C. Chow, “Mineralization in Natural and Synthetic Biomaterials” (P. J. Li,
P. Calvert, T. Kokubo, R. Levy, and C. Scheid, Eds.), p. 27. Materials Research Society, PA, 2000. © 2000.
6
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
(a)
(b)
1 µm
10 µm
Figure 2. A typical apatite coating formed in SBF after 1 d soaking, on the surface of CPTi subjected to treating
with 30 mass% H2 O2 containing 3 mM TaCl5 at 80 C for 3 d. (a) low magnification and (b) high magnification
showing the tiny flakes united together to form large globules.
strong (002) peak, suggesting a preferential growth of the apatite crystal. The XRD peaks
corresponding to (211), (112), and (300) planes overlap to give a broad diffraction located
around 2 = 32 . The apatite is poorly crystallized. Figure 4 displays a typical Fourier
transform infrared reflection (FT-IR) spectrum of the apatite. The band at 875 cm−1 and
the bands between 1418 and 1456 cm−1 correspond to the carbonate group incorporated
in the apatite. The two bands at 597 and 555 cm−1 are from the 4 mode of O–P–O bending
in apatite, whereas the peaks at 1110 and 1000 cm−1 indicate the 3 band of P–O stretching
mode. The peak at 472 cm−1 is ascribed to the 2 mode of O–P–O bending in apatite. The
shoulder at 961 cm−1 reflects the 1 band of P–O stretching mode in apatite [28, 29].
The electrochemical impedance spectroscopy (EIS) technique was found to serve well
to investigate apatite growth on the surface of alkali-treated titanium and its alloys as
well as oxide films on the untreated titanium surface. EIS is sensitive to detecting an
increase in electrical resistance on the outmost surface when apatite nucleates [30, 31]. At
this nucleation stage, apatite cannot be clearly seen even under SEM. Typical EIS spectra
of alkali-treated Ti before and after immersion in SBF are shown in Figure 5. For these
spectra, by considering the nucleation and growth of apatite on the substrate surface, an
equivalent circuit was proposed and curve-fitting for obtained spectra was conducted. It
was observed that the electrical resistance of apatite formed in vitro (Rap ) increased with
the immersion time, indicating the growth (in thickness) of the apatite layer [30, 31].
Nano-indentation was used to determine the mechanical properties of apatite formed
on a bioactive composite in SBF [32]. Cross-sections of composite specimens that had been
immersed in SBF for various periods of time were ground and polished. A Berkovich
diamond indenter was used for the nano-indentation tests, and a series of indentations
were made across the specimen surface. Figure 6 exhibits a nano-indentation and a nanoindentation curve of the apatite. The elastic modulus across the apatite layer had little
variation, with an average value being around 5.22 GPa. This modulus value did not
Ap (002)
Ti
Ap
Ti
Ti
Ap
Ap
R
20
25
30
35
2 theta/degree
Figure 3. TF-XRD pattern of the apatite coating shown in Figure 2.
40
7
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
Phosphate
∆ Carbonate
∆
Reflectance /arb. unit
∆
2000
1600
1200
800
400
Wavenumber/cm–1
Figure 4. FT-IR spectrum of the apatite coating shown in Figure 2.
Bode Z & Phase (deg)
1.0E+07
90
Impedance
Angle
80
60
Z (ohms)
1.0E+05
50
40
30
1.0E+03
Phase of Z (deg)
70
20
10
1.0E+01
0.001
0.1
10
1000
0
100000
Frequency (Hz)
Bode Z & Phase (deg)
1.0E+07
Impendance
Angle
90
80
Z (ohms)
60
50
40
1.0E+03
Phase of Z (deg)
70
1.0E+05
30
20
1.0E+01
0.001
0.1
10
1000
10
100000
Frequency (Hz)
Figure 5. Bode plot for pretreated Ti before and after immersion in SBF for eight weeks. Reprinted with permission from Ref. [31], C. X. Wang et al., Biomaterials 24, 3069 (2003). © 2003, Elsevier Science Ltd.
8
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
(a)
(b)
Figure 6. Nano-indentation of apatite formed on 20% HA/PHB (immersion time: 1 month): (a) nanoindentation and (b) nano-indentation curve.
vary significantly with regard to the immersion time of the composite in SBF. It was also
found that elastic modulus values of apatite formed in vitro on various biomaterials were
within a narrow range [32].
2.3. Correlation Between In Vitro Apatite Formation
and In Vivo Bone Bonding
It has been found, up until now, that nearly all materials that are able to bond directly to
the surrounding bone tissue without being separated by a fibrous capsule induce apatite
deposition in SBF at 36.5 C within short periods of time [11, 33, 34]. A comparative
study between in vivo bone formation and growth and in vitro apatite formation in SBF
was performed using five kinds of Na2 O-CaO-SiO2 glasses of different compositions [11].
These glasses had various apatite induction times in SBF of 12 h, 3 d, 7 d, 21 d, and
longer than 28 d. They were implanted in holes of the femoral condyles of rabbits. The
bioactive glasses with the apatite induction times of 12 h and 3 d directly bonded with the
surrounding bone tissue, while those with apatite induction times of 7 d, 21 d, and longer
than 28 d showed no bone-bonding ability even after 12 weeks of implantation. The
formation of an Si-rich layer and bone growth were significantly earlier on the bioactive
glass with the apatite induction time of 12 h than those on the glass with the apatite
induction time of 3 d. Therefore, the in vivo bioactivity of a material could be shown
through its apatite-forming ability in SBF. On the other hand, it was noted that the glasses
that induced apatite deposition in SBF with durations longer than 7 d failed to form
direct bonding with bone in the in vivo tests. In a separate study [35], it was shown that,
although an alkali-treated CPTi induced apatite deposition in SBF within 24 h, it also
failed to form a direct bonding with the surrounding bone tissue in vivo, because of the
reactive surface layer [36]. Therefore, it cannot be generalized that biomaterials being able
to induce in vitro apatite deposition in SBF would also exhibit in vivo bioactivity. Extreme
care should be taken when assessing the bioactivity of those biomaterials that induce
apatite deposition in SBF but require a significantly long time. An in vivo experiment
seems to be inevitable for confirming the bioactivity for such biomaterials.
3. MECHANISMS OF IN VITRO APATITE DEPOSITION
It is now clearly understood that all bioactive biomaterials are characterized of being
bonded with surrounding tissues by forming a layer of apatite on their surfaces. In vitro
apatite deposition in SBF could indicate in vivo bioactivity. Therefore, understanding the
in vitro apatite deposition mechanism and knowing the influencing factors for apatite
formation are important for the development of bioactive materials. Apatite deposition
on bulk bioactive materials or bioactive coatings can be ascribed to the functioning of
surface silica gel, titania gel, a Ca-P layer, other functional groups, or their combinations.
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
9
3.1. Silica Gel
Investigations into bioactive glasses have shown that five reaction stages occurred on the
material side of the interface during the formation of chemical bonding with surrounding
tissue [37]:
(1) A rapid exchange of Na+ or K+ with H+ or H3 O+ from solution, forming silanols
(Si-OH)
(2) Loss of soluble silica in the form of Si(OH) to the solution, resulting from the breaking of Si–O–Si bonds and the formation of silanols at the glass-solution interface
(3) Polycondensation of silanols to form a hydrated silica gel
(4) Formation of an amorphous calcium phosphate (ACP) layer through migration of
Ca2+ and PO3−
4 groups to the surface
−
(5) Crystallization of the ACP layer by the incorporation of OH− , CO2−
3 or F anions
from solution to form a bone-like apatite
The cation leaching, silica network formation, and phosphate reprecipitation occurred
simultaneously during the dissolution of Bioglass 45S5 in a tris-buffered solution at pH 8
[38]. The presence of silanol groups on the surface was confirmed to be necessary for
apatite deposition. The formation of the Si-rich layer as well as the apatite layer around
bioactive glass particles when implanted in bone has been confirmed in vivo [11].
Sol–gel derived silica induced apatite deposition but dense silica glasses or quartz did
not [39]. The silica film derived by a sol–gel dip coating followed by a heat treatment did
not induce apatite deposition in SBF either, due probably to the dense structure [40, 41].
These observations emphasize the importance of both the Si-OH groups and the defects
on the substrates for apatite deposition. However, larger amounts of Si-OH groups do
not necessarily lead to higher apatite nucleation rates [40, 42]. Pores acted as nucleation
sites for apatite nucleation, and the CHA nucleation rate increased with increasing pore
size and pore volume of a porous gel-silica substrate [42]. Pore sizes larger than 2 nm
were found to be necessary to achieve rapid apatite deposition [42]. However, no clear
relationship could be found among porosity, Si-OH contents, or synthesis conditions and
the in vitro bioactivity for sol–gel derived porous silica films with a wide variety of
microstructures [42]. It was shown that the mesoporous material MCM-41 could possess high specific surface (BET specific area of 1134 m2 /g), high porosity (pore volume
0.984 m3 /g; peak pore size 2.5 nm), and consist of siloxane and silanol groups, yet they
induced no apatite deposition even after two months soaking in SBF [43]. Therefore, it
seems that the presence of superficial silanol groups conferring negative charges to the
surface and a high porosity are not sufficient to ensure apatite formation. Incorporation
of 10% glass, which had a molar composition of 55SiO2 -41CaO-4P2 O5 , into the mesoporous material resulted in apatite deposition on the surface after 7 d soaking in SBF.
These results show the importance of ionic interchange between the substrate and SBF
for apatite deposition. The released Ca2+ cations left holes in the substrate that could
act as nucleation sites themselves, instead of the initial pores in the material. Ordered
mesopores were found to contribute to the apatite deposition ability of a bioactive glass
(58S) because of the increasing specific surface area and hence the increasing amounts
of silanol groups on the surface [44]. The mesoporous bioactive glass with a medium
concentration of Si/Ca ratio (80S) was found to possess the best in vitro bioactivity, in
contrast to conventional sol–gel derived bioactive glasses which increased in bioactivity
with increasing calcium contents [45]. This fact supports the idea that the mesoporous
structure do affect readily the in vitro bioactivity.
3.2. Titania Gel
CPTi implants attach to the surrounding bone without being separated by a fibrous tissue,
due to its natural oxide layer, which is hydrolyzed to form Ti-OH groups and negatively
charged when immersed in human physiological fluid. The Ca2+ ions in the fluid are
then attracted onto the surface, which in turn attract the PO2−
4 ions to initiate the deposition of apatite [46, 47]. This postulation of the apatite-forming process is supported by
10
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
the fact that the titania gel derived by sol–gel, which is negatively charged in SBF and
also full of Ti-OH groups, favors apatite deposition in SBF [19]. Apatite formation was
also found to be accelerated on an amorphous sodium titanate surface layer that was
produced by treating CPTi with an alkali solution [35, 48]. While exposed to SBF, the
alkali ions from the alkali titanate layer were released into SBF, which was accompanied
by ion exchange with H3 O+ ions and thus resulted in Ti-OH functional groups on the
surface as well as a pH increase in the surrounding fluid. The Ti-OH groups soon combined with the Ca2+ in the SBF to form a calcium titanate. After a certain period, the
calcium titanate took phosphate ions as well as calcium ions from SBF to form apatite
nuclei, which then grew spontaneously due to the high supersaturation of the fluid with
respect to apatite. The pH increase also gave rise to an increase in the driving force for
apatite deposition and hence favored apatite deposition [35, 48]. As proven by TEM and
X-ray photoelectron spectroscopy (XPS) examinations, the calcium and phosphate deposition on the alkali- and heat-treated Ti surface was in sequence. Calcium ions can be
adsorbed on the surface of the alkali- and heat-treated Ti while phosphate ions cannot
without the pre-adsorption of calcium ions. Calcium deposition is thus believed to be the
prerequisite for apatite deposition [26, 49]. Phosphate and calcium were detected by XPS
on the acid- and alkali-treated titanium surface after soaking in SBF for a time as short as
only 2.5 min. The Ca/P ratio for alkali-treated Ti was significantly higher than that of the
acid-treated samples [50]. These observations confirmed that the alkali-treated Ti surface
favors the adsorption of calcium. It was noticed that the subsequent heat-treatment of an
alkali-treated amorphous titania coating, which was derived by vacuum plasma spraying, inhibited the exchange between Na+ ions in coatings and H3 O+ ions in SBF; hence
reducing the Ti-OH surface groups, leading to a reduced in vitro bioactivity [51].
It should be noted that the apatite-forming process on the sodium titanate surface of
alkali- and heat-treated Ti cannot be simply applied to other titania gel surfaces. For
some other surface coatings containing mainly titania gel, it is likely that the adsorption
of both calcium and phosphate ions occurs simultaneously, which represents a different process for apatite growth on the titania surface. It was found that both calcium
and phosphate ions were adsorbed separately and quickly on different sites of a sol–gel
derived titania surface and that calcium and phosphate adsorption increased once the
corresponding counterion (phosphate or calcium) had been previously adsorbed [52]. The
possible reactions leading to the adsorption of calcium and phosphate ions on the titania
surface are [53]:
Ti − OH + Ca2+ ←→ Ti − OH − Ca+ + H+
(7)
Ti − OH + HPO42− ←→ Ti − O − HPO3− + OH−
(8)
The assumption that apatite deposition in SBF can be ascribed to a negatively charged
surface is supported directly by the fact that the negatively charged barium titanate
(BaTiO3 , BTO) surface, which is ferroelectric and piezoelectric after a poling treatment,
showed Ca-P crystal growth, while no Ca-P phase was observed on the positively
charged BTO surface [54]. Similarly, the negatively charged HA surface, which was
obtained by polarization, induced apatite deposition in 1.5 SBF (a solution with inorganic
ion concentrations 1.5 times the SBF), whereas the positively charged or the non-polarized
HA surface did not [55]. However, the assumption that the negatively charged surface
initiates apatite deposition has been challenged by the finding that the CPTi surface con2+
tained more HPO−
ions after soaking in SBF for a certain duration [53].
3 than Ca
CPTi subjected to close contact with other surface, such as polytetrafluoroethylene
(PTFE), enhanced its apatite-forming ability, due probably to the accumulated titanium
ions and the increased local supersaturation of SBF within the closed region [56]. Therefore, as a logical deduction, a porous surface full of cracks is expected to favor apatite
deposition. This view is supported by the finding that micro-holes with sizes of around
0.3 mm on the alkali-treated Ti-6Al-V favored apatite deposition both in vitro and in vivo
[57]. However, Peltola et al., on the contrary, found that only surface topographies giving
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
11
peak distances of 15–50 nm, observed by atomic force microscopy (AFM), favored apatite
deposition [58].
The crystalline structure of titania, which has three common polymorphs of anatase,
rutile, and brookite, appears to be more important to initiate apatite nucleation as compared to the increase of supersaturation with respect to apatite in SBF, due to the release
of certain ions from the surface layer. The amorphous sol–gel derived titania induced
no apatite deposition in SBF [56]. The incorporation of Na+ and Ca2+ ions in the amorphous titania did not result in apatite deposition either, in spite of the fact that the
release of Ca2+ ions as well as the exchange between Na+ and H3 O+ in SBF increased
the supersaturation with respect to apatite [60]. On the contrary, the anatase gel containing no such Na+ and Ca2+ ions induced apatite deposition effectively [60]. Similarly, the
amorphous titania gel containing CaO did not induce apatite deposition while the pure
anatase gel did [61]. Also, an intermediate hot water soaking to remove sodium from
the sodium titanate layer produced by the alkali-treatment of CPTi gave pure anatase
upon a subsequent heat treatment, and subsequently enhanced the in vitro bioactivity
[62, 63]. Therefore, it can be postulated that the perquisite for an enhanced ability of
titania layer for early apatite deposition is to achieve a crystalline structure of titania.
However, it was noticed that a film of the mixture of anatase and rutile fabricated by
micro-arc oxidation (MAO) of CPTi that contained Ca did not induce apatite deposition in 1.5 SBF [64]. Only after a subsequent hydrothermal treatment, which resulted in
numerous precipitates assumed to be Ca-P and did not alter the phase composition or
the porous surface morphology, could the film induce apatite deposition in 1.5 SBF. The
MAO-derived anatase coating containing up to 8 at.% P induced no apatite-formation
when subjected to a SBF-soaking for up to 28 d; however, after a subsequent hydrothermal treatment, apatite was induced on the anatase surface after only 9 h of immersion
[65]. Similarly, apatite formation was found only after the alkali-treatment (soaking in an
NaOH aqueous solution) of an MAO-derived anatase film [66]. These findings emphasize
the importance of the surface hydroxyl groups, which have resulted from the hydrothermal treatment, for apatite deposition. However, it is not clear why the crystalline titania
film derived by MAO failed to induce apatite deposition.
Apatite deposition was found to be more pronounced on anatase gel than on rutile gel,
both of which were prepared by the sol–gel process followed by heating at different temperatures [59]. Titania coating derived by soaking in an anatase slurry was also found to
induce more apatite deposition than that derived by soaking in a rutile slurry [67]. However, it was also shown that the low-temperature titania film prepared through simply
soaking CPTi at 80 C for 3 d in a 30 mass.% H2 O2 solution containing 3 mM TaCl5 , which
consisted predominantly of rutile, was equally effective for inducing apatite deposition
in SBF, as the lattice mismatch between rutile and apatite was as small as that between
anatase and apatite [29]. It was further noted that the pure rutile film derived by various
chemical and electrochemical methods exhibited a high ability to induce apatite deposition in SBF [68–72]. Therefore, it may be concluded that the rutile phase has a higher,
or at least the same, ability to induce apatite deposition in SBF, in situations where the
rutile film contains Ti-OH functional groups.
Titania nanotube arrays can be obtained by anodic oxidation of pure Ti in electrolytes
containing fluoride. It has been reported that the unique nanostructure of nanotubes
accelerated the apatite-forming procedure in SBF when compared to a compact anodic
titania layer. The in vitro bioactivity improved after annealing the amorphous titania
nanotube arrays, which crystallized to anatase or a mixture of anatase and rutile, depending on the annealing temperatures [73]. This further confirms the importance of a crystalline structure for titania to induce apatite formation in SBF.
3.3. Calcium-Phosphate (Ca-P) Layer
Apatite deposition on bioactive ceramics, such as HA, TCP, and calcium phosphate, containing bioactive glasses can be ascribed to the formation of a surface Ca-P layer. The
apatite deposition process on these bioactive bioceramics is relatively simple. The dissolution of Ca2+ ions into the surrounding body fluid increases the supersaturation of
12
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
the solution with respect to apatite and at the same time leaves newly formed defects
on the surface, which in turn favors apatite nucleation. The bone-like apatite deposited
on sintered HA surfaces in SBF comes into being through crystallization of the previously formed Ca-rich, amorphous (or more precisely, nanocrystalline) calcium phosphate
(ACP). The zeta potential characterized by electrophoresis indicated that during the exposure to SBF, the HA surface was negatively charged, which interacted with positive
calcium ions in the fluid to form ACP [74].
It was found that the ACP incorporated in HA coatings, especially the coatings prepared by plasma spraying, plays an important role in the in vitro bioactivity of these
coatings [75–78]. After soaking in SBF, the dissolution of the amorphous phase increased
the supersaturation of SBF with respect to apatite and at the same time also increased the
surface roughness, which provided nucleation sites because the existence of defects in
the structure lowers the nucleation energy [75]. A thermal treatment could increase the
crystallinity of plasma-sprayed HA coatings, leading to decreased solubility and hence
more stable mechanical properties, but it reduced bioactivity of the coatings [78]. The
totally crystallized plasma-sprayed HA coating achieved through heat treatment after
plasma spraying did not induce apatite deposition in SBF, although heat treatment did
not alter significantly the surface morphology [75, 78]. Therefore, if only the osteoconductive property of Ca-P coating is desired for the initial fixation of implants, having ACP
is advantageous. On the other hand, if coating longevity is desired, the crystalline HA
coating is preferred [79]. It was also noted that an increased osteoblast-like cell activity
appeared on a sintered pure HA surface as compared with the as-received and calcined
pure HA samples [18], which suggests that the crystal size, and probably the crystallinity
as well, of HA may play an important role in governing cellular response. However,
higher crystallinity of the Ca-P coating may not necessarily mean poorer bioactivity. No
differences were found in the bone apposition abilities of the implants coated with 62%
crystalline HA, 30% crystalline HA or fluorohydroxyapatite (Ca5 (PO4 3 OH1−x Fx , FHA),
with all coatings being made by plasma-spraying [80, 81]. In a separate study, HA coatings with different degrees of crystallinity were found to have no significant influence
on bone-bonding ability, bone-bonding strength, tensile failure mode, and degradation
of the coating [82].
3.4. Other Surface Functional Groups
Surface functional groups that are able to induce apatite deposition include Si-OH
[26, 37, 83], Ti-OH [26, 84], Ta-OH [26, 85–87], Nb-OH [26], Zr-OH [26, 88], -COOH
[83, 89, 90], -NH2 [83], and -H2 PO4 [89, 90]. When biomaterials having these surface functional groups are subjected to immersion in physiological fluid, the Ca2+ ions in the fluid
are attracted by the functional groups and bind to the surface, which in turn attracts
the phosphate ions and hence initiates the apatite nucleation process. Because the physiological fluid is a supersaturated solution with respect to the apatite, the subsequent
apatite growth is a spontaneous process. Toworfe et al. utilized a self-assembled monolayer (SAM) technique to create amine (-NH2 ), carboxyl (-COOH), and hydroxyl (-OH)
functionalized surfaces onto oxidized silicon wafers. The results show that the hydroxyl
groups exhibited the most remarkably apatite-inducing ability in SBF [83].
3.5. Effects of SBF-Related Factors
Very few papers can be found that clarify the possible effects of the SBF itself on apatite
deposition. After soaking alkali- and heat-treated CPTi in the R-SBF, which has a higher
concentration (27.0 mM, exactly equal to that in human plasma) of HCO−
3 than the
conventional SBF (4.2 mM), OCP, not apatite, nucleated directly from the previously
deposited amorphous calcium phosphate layer, due probably to the inhibiting effects
of HCO−
3 on apatite precipitation. The OCP crystals continuously grew on the treated
titanium surfaces rather than transforming into apatite [91]. However, Muller [92] and
Jalota et al. [21] achieved apatite on alkali-treated Ti-alloy substrates also using SBF
−
with a HCO−
3 concentration of 27.0 mM. Muller et al. found that the HCO3 content in
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
13
SBF influences the composition and structure of the obtained HA. B-type carbonated3−
in HA) precipitated on chemically-treated Ti substrates when
HA (CO−
3 replaces PO4
−
−
the SBF contained less than 20 mM HCO−
3 ; A-type carbonated-HA (CO3 replaces OH
−
in HA) formed with higher HCO3 concentration in SBF [92]. Jalota et al. found that
increasing the HCO−
3 content in SBF from 4.2 mM to 27.0 mM not only speeded up the
apatite-forming procedure on the alkali-treated Ti6Al4V strips, but also modified the morphology of the deposited apatite, which also yielded enhanced cell viability and protein
concentrations [21].
The influence of some ions in 5 SBF (a solution with ion concentrations five times
the SBF) on the nucleation of biomimetic Ca-P coatings on Ti-6Al-4V has been studied
[93, 94]. Sodium chloride increased the ionic strength of the solution and hence delayed
the homogeneous Ca-P precipitation in solution, allowing Ca-P to nucleate heterogeneously on the substrates [93]. HCO−
3 inhibited apatite formation and hence reduced the
apatite crystal size of the coating, allowing a better physical attachment of the coating to
Ti-6Al-4V substrates [93]. Mg2+ was found to have a stronger inhibitory effect on apatite
crystal growth than HCO−
3 [94]. An XPS depth profile showed significantly increased
contents of Mg2+ , as well as Ca2+ , at the titanium/coating interface. A high concentration
of Mg2+ favored heterogeneous nucleation of tiny Ca-P globules on the substrate and
hence enhanced the physical adhesion of coating at the early stage of coating formation
[94]. However, it is still not clear whether those ions function similarly in SBF.
As reported in numerous publications, the in vitro bioactivity of biomaterials is generally assessed using SBF in the static condition. It has been suggested that a dynamic
process, with continuous exchange of SBF solution with the material, better mimics the
actual environment in the human body and hence serves better to evaluate bioactivity
[50, 95, 96]. Both static and dynamic processes were used to evaluate the bioactivity of
a sol–gel derived glass having the composition of 55%SiO2 -41%CaO-4%P2 O5 [95]. In the
static condition, changes in the Ca2+ ion concentration and pH value were significant,
while these two parameters remained almost unchanged in the dynamic condition during
in vitro tests. Apatite deposition was observed in both conditions. However, the dynamic
condition led to an accelerated apatite-forming rate. The morphology of the apatite layers was different: in the static condition, the glass was covered by an incoherent layer
of spherical particles of about 1 m in size; while in the dynamic condition, the glass
was fully covered by a layer in which the spherical particles were constituted by hundreds of needle-like crystal aggregates. Similar phenomena were observed on a sol–gel
silica-polymer composite [96].
It appears that many factors can influence apatite deposition on biomaterials. These
factors include the material itself, pretreatment of the material, material surface morphology (including porosity), and the material’s surface composition and crystallinity.
Although great efforts have been made to study the mechanisms of apatite formation on
various materials, conflicting results have been reported, and the exact apatite-forming
mechanisms still remain ambiguous for a large percentage of biomaterials investigated.
4. BIOACTIVE COATINGS ON NON-METALLIC
BIOMATERIAL SUBSTRATES
4.1. Bioceramics
Structural ceramics, such as high purity alumina with fine grains [97, 98], and toughened
zirconia ceramics [99] possess high strength, acceptable fracture toughness, high chemical
stability, and excellent wear resistance and have already been used for femoral heads of
total hip prostheses. Apart from their applications as coatings on metallic or polymer
substrates, bioceramics are also used in the forms of particles and bulk shapes in various
clinical situations. For bulk ceramics, such as alumina, bioactivity is achieved through
either having bioactive coatings or incorporating secondary bioactive phases. The apatiteforming processes on some bioactive glasses, ceramics, and glass-ceramics have been
investigated extensively.
14
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
4.1.1. Alumina
Thin HA films have been deposited on alumina substrates through radio-frequency sputtering [97, 100] or ion-beam assisted deposition [101], followed by a thermal treatment.
Sol–gel dip coating from a solution of 1.67Ca(NO3 -PO(OH)x (OEt)3−x [102–104] or spin
coating with a sol consisting mainly of calcium nitrate and phosphoric acid in methanol
[105], with a subsequent thermal treatment, also deposited HA layers on alumina. The
addition of an appropriate citric acid into the dipping solution resulted in not only a
rough or porous morphology but also an improved adhesion strength of the HA coating
[102, 104].
Porous -TCP films with a thickness of ∼50 m were formed on alumina ceramics
using a spray-pyrolysis technique [106]. Improved cellular responses with respect to cell
proliferation and differentiation were found on the coated specimen as compared with
the uncoated alumina.
Full-density medical-grade -alumina was coated with a bioactive FHA glass-ceramic
layer (50 m thick) through a suspension deposition and sintering process, with an intermediate layer of a SiO2 -CaO glass (400 m) to aid the adhesion and also to impede
the coating/alumina interfacial reaction [107, 108]. Bone-like apatite precipitated on the
coating, but didn’t cover the whole coating, after 1 month of soaking in SBF [108]. An
in vitro study using human osteoblast-like cells showed that the coatings exhibited cell
attachment and proliferation similar to bulk FHA glass-ceramic and obviously better than
alumina [107]. Enhanced protein adsorption that favored osteoblast-like cell adhesion
and spreading and reduced the macrophage complement-mediated local inflammatory
reaction as well, were observed on the FHA glass-ceramic coatings as compared with the
alumina [108].
4.1.2. Hydroxyapatite
Soaking HA in SBF leads to the formation of bone-like apatite on its surface through the
following surface structural changes in sequence: formation of a Ca-rich amorphous or
nano-crystalline Ca-P, formation of a Ca-poor crystalline calcium phosphate, crystallization into bone-like apatite. The exposure time for the formation of the Ca-P was delayed
on HA sintered at 1200 C, as compared to the one sintered at 800 C, because a higher
sintering temperature resulted in reduced surface hydroxyl groups and surface negativity [74]. Stoichiometric HA is generally considered to possess a limited in vitro reactivity
[55, 109]. A composite of HA and 5% sol–gel glass exhibited higher in vitro bioactivity
than pure HA [109]. The small amount of the bioactive glass was enough to induce a
positive response in just a few hours in SBF, which led to the formation of a new apatitelike layer that reached a thickness of 3 m after about 3 d. The incorporation of silicon
into the HA structure provides another way to improve the bioactivity [110]. An XPS
4−
investigation revealed the substitution of PO3−
4 groups with SiO4 groups. An increase of
the silicon content up to 1.6 wt% led to the polymerization of the silicate species on the
surface. The formation of apatite on the surface of silicon-substituted HA was strongly
enhanced as compared to pure HA. The samples containing monomeric silicate species
showed higher in vitro bioactivity than that of silicon-rich sample containing polymeric
silicate species [110].
Under the influence of an electric field of DC 1 kV/cm at 300 C, HA can be polarized due to proton movement [55]. After soaking in 1.5 SBF for 1 d, apatite deposition
took place on the negatively charged surface (N-surface), whereas no crystal growth
was observed on the positively charged surface (P-surface) even after a 3 d immersion
in 1.5 SBF. The apatite growth rate on the N-surface was also greater than that on the
non-polarized HA surface. Similarly, the negatively charged surface of barium titanate
(BaTiO3 , BTO), which is ferroelectric and piezoelectric after poling treatment, showed
Ca-P crystal growth, while no Ca-P phase was detected on the positively charged BTO
surfaces [54].
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
15
4.1.3. Glass-Ceramics
Bioactive glass-ceramics were prepared through tape casting 45S5 Bioglass followed by
sintering at 800, 900, or 1000 C for 3 h, which produced a crystalline material with a
major phase of Na2 Ca2 Si3 O9 [111]. Samples sintered at 1000 C induced apatite deposition
in SBF after 20 to 24 h, while the apatite induction time for samples sintered at 900 C
was only 2 h. The difference in bioactive response was attributed to the specific surface
area of sintered 45S5 Bioglass® [111].
The bioactive glass-ceramic A-W, which was prepared through heat treatment of a
MgO-CaO-SiO2 -P2 O5 glass to precipitate crystalline apatite and wollastonite, showed
high in vitro bioactivity. However, the addition of a small amount of Al2 O3 to this glassceramic to improve its bending strength resulted in the elimination of its apatite deposition ability. After being treated with 0.1 M HCl, apatite deposition occurred due to
the formation of hydrated silica on the surface of the alumina-containing glass-ceramic
A-W [112].
The glass-ceramic 60CaO · 30P2 O5 · 7Na2 O3 · TiO2 formed apatite after 20 d soaking in
SBF [113]. After a hydrothermal treatment in water at 140 C for 1 h, this glass-ceramic
was completely covered with bone-like apatite after 10 d soaking in SBF, due mainly to
the hydrated titania groups resulting from the hydrothermal treatment.
4.1.4. Other Bioceramics
A wide variety of bioinert surfaces, such as Teflon, silicon wafers, gold, and quartz, can
be converted to activated surfaces for biomolecule adsorption by coating them with a thin
sol–gel titania film, which can bind strongly with proteins, peptides, and polysaccharides.
Before applying the sol–gel titania coating, the bioinert surfaces should be oxygenated or
contain free hydroxyl groups [114].
It has been shown that both the 80SiO2 -20CaO and the 80SiO2 -20CaO-3P2 O5 (in mol%)
gel glasses induced apatite deposition in SBF [115]. The addition of P2 O5 in the gel glass
led to a quicker growth of apatite crystals [115]. Substitution of M2 O3 (M = Ga, Al, In,
La, Y) for CaO in the binary CaO-SiO2 glass composition progressively reduced the ability
to form a Ca-P layer on the surfaces exposed to SBF [116].
The trace amounts of Fe2 O3 in the -wollastonite (CaSiO3 ) matrix of ferromagnetic
glass-ceramics, which were designed as thermoseeds for hyperthermia treatment of
bone tumors and consisted of magnetite (Fe3 O4 ) precipitation in the -wollastonite glass
matrix, were found to inhibit the apatite deposition in SBF [117]. The addition of 3 wt%
Na2 O and the combination of B2 O3 and P2 O5 into the glass-ceramics improved in vitro
bioactivity by inducing apatite deposition in SBF within 10 to 30 d, because for Na2 O,
the exchange of Na+ with the H3 O+ in the SBF increased the pH value of the surrounding fluid, and the combination of B2 O3 and P2 O5 may have accelerated apatite formation
through the formation of hydrated silica on the surfaces.
A study revealed that the akermanite ceramic (Ca2 MgSi2 O7 ) possessed an apatiteformation ability comparable to the wollastonite ceramic after soaking in SBF for 20 d. In
addition, the Ca, Si, and Mg ions dissolved from akemanite at certain ranges of concentration stimulated significantly osteoblast and cell proliferation, which suggests additionally
that the ceramic is a suitable candidate as bone substitutes [118].
DCPD plays an important role in the biological mineralization process as well as the
setting of a variety of calcium phosphate cements for orthopedic and dental uses. It
was shown that a layer of apatite could be coated on the surface of gel-grown DCPD
single crystals through a hydrothermal-electrochemical deposition method, which used
heated SBF in an autoclave as an electrolyte [119]. The apatite formed was confirmed
through SEM, XRD, and FT-IR analyses. Under physiological conditions, DCPD gradually
transforms into apatite.
Phospholipids have been found to initiate calcium phosphate deposition in cartilage, bone, healing fracture callus, and certain calcifying bacteria. To study the possibility of using Ca-P complexed phospholipid (Ca-PL-PO4 ) coatings on solid surfaces
for enhancing bone-implant interactions, Ca-PL-PO4 complexes were prepared using
16
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
CaCl2 , (NH4 2 HPO4 , and various phospholipids, such as phosphatidylcholine (PC), phosphatidylserine (PS), and phosphatidylinositol (PI) [120]. After coating the complexes on
glass substrates, samples with coatings were subjected to soaking in SBF, and cell culture was also carried out. Results showed that surfaces coated with Ca-PS-PO4 exhibited
the highest amount of calcium deposition and enhanced osteoblast differentiation [120].
The Ca-PS-PO4 -coated CPTi surfaces exhibited enhanced protein synthesis and alkaline phosphatase (ALP) activity as compared to other phospholipids-coated or uncoated
surfaces [121].
HA was also deposited on carbon fabric through a sonoelectrodeposition approach. An
increased current density achieved a much denser and uniform structure with smaller
HA crystal sizes [122].
4.2. Biomedical Polymers
Polymers have a long history for their applications in medicine. Synthetic polymers
such as poly(methyl methacrylate) (PMMA) and ultra high molecular weight polyethylene (UHMWPE) and poly(glycolic acid) (PLA) are successfully used in orthopedics and
surgery. In recent years, natural polymers such as collagen and chitosan have gained
increasing attention. Major efforts have been made to use polymers for hard tissue
replacement/regeneration. For achieving bone-bonding, investigations have been conducted to make polymer surfaces bioactive (osteoconductive). On the other hand, bone at
the microstructural level is a natural composite consisting of bone apatite about 50 nm ×
25 nm × 3 nm in size and collagen [123]. Efforts have thus been made to produce bone
substitutes with an analogous structure to that of natural bone through either industrialized plastics processing technologies or biomimetic processes. These materials should
have the ability to induce in vitro apatite deposition.
Several approaches to provide biomedical polymers with the ability to induce apatiteformation ability in SBF will be discussed in the next several paragraphs. However,
it has also been reported that certain polymers possess an inherent ability to induce
apatite deposition without additional treatments. For example, when soaking poly(2hydroxyethylmethacrylate) (PHEMA)-based hydrogels in SBF at 37 C for several weeks,
apatite deposition was significant on all the hydrogels. The nucleation of apatite was
ascribed to the strong tendency of Ca2+ ions to chelate to oxygen atoms in the unit structure of PHEMA, which was supported by the fact that the degree of calcification for the
hydrogels lowered with decreasing oxygen atoms as a result of certain copolymerizations
[124]. A type of silk protein, sericin, also possessed the ability to induce apatite nucleation in a 1.5 SBF when having a -sheet structure that forms polypeptides with ordered
and concentrated carboxyl groups [125].
4.2.1. Surface Seeding with Ca-P Phases
Many biological materials, including bone that is discussed above, are composites with
particles reinforcing a polymer matrix. In the biological environment, the particles are
grown in situ within the polymer matrix and under the control of the matrix. Calvert and
Mann reviewed some synthetic and biological composites formed by in situ precipitation
[126]. According to the matrix-mediated processes, they classified biological composites
into four categories: Type I (matrix-inert), Type II (nucleation), Type III (amorphous)
and Type IV (oriented) biocomposites. They also discussed the principles governing the
growth of particles when constrained by a polymer matrix. It appears possible to produce
various types of precipitates within a polymer matrix. However, through an extensive
literature survey, it seems that the majority of in situ precipitation investigations have
ended with products that have precipitates either on polymer surfaces or into a limited
depth of the bulk polymer if the polymers are soaked in the (normally ion-saturated)
solutions. There are good examples of Ca-P particles precipitating only in areas near the
surface of collagen membranes.
A biomimetic process was used to coat biomedical polymers with apatite [79, 127–131].
This biomimetic process included the following two steps: (1) seeding the organic substrates with apatite nuclei by bringing them in contact with particles of CaO-SiO2 -based
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
17
bioactive glasses (termed “G glass”) soaked in SBF, and (2) coating the organic substrates
with apatite by a subsequent soaking in 1.5 SBF, in which apatite nuclei grew on the
substrate in situ to cover the whole surface. This biomimetic method was also used to
coat apatite on biomedical hydrogels [132]. When using this method to coat apatite on
titanium substrate, the G glass could be replaced by a sodium silicate solution at the
seeding stage [133].
Apatite deposition on polystyrene could be accelerated by soaking the polymer substrate firstly at 37 C in 5 SBF, followed by soaking in a second 5 SBF solution which
was devoid of crystal growth inhibitors of Mg2+ and HCO−
3 [134]. The pH value of the
first soaking solution affected the final apatite structure. A high pH value led to larger
single crystal plates while a low pH value resulted in minute, polycrystalline plate-like
structures. The surface of silica gel beads or plates could also be coated with an apatite
layer through alternative soaking in a CaCl2 /tris-HCl aqueous solution and a Na2 HPO4
aqueous solution [135].
Calcium alginate fibers were produced by extruding an aqueous sodium alginate solution through nozzles into an aqueous calcium chloride solution first and then through a
calcium chloride methanol solution [136]. After soaking the fibers in an aqueous saturated
calcium hydroxide solution for 5 d and then in SBF for 7 d, apatite was deposited on
the surface of the fibers with certain alginate subunits, due to the release of the calcium
ions from the fibers, which left free carbonxyl groups and increased the supersaturation with respect to apatite in SBF as well. Du et al. produced various types of calcium
phosphate/collagen composites through the mineralization of the type I collagen matrix,
2+
which was realized by soaking collagen in a PO3−
4 containing solution and then in a Ca
containing solution [137, 138]. The mineral contributed up to 60–70% of the weight of
the composites, and the strength and Young’s modulus of the composites exceeded the
lower bound of bone. The interconnecting porous structure of the composites provided
a large surface area for cell attachment and sufficient space for nutrient transportation.
New bone matrix was synthesized at the bone-implant interface when the composite was
in contact with bone fragments [137].
4.2.2. Surface Functional Groups of Si-OH, COOH, Ti-OH, and so on
Silanol groups were produced on silicones through a sol–gel process [139]. Although the
silanol-modified silicones did not induce apatite formation in SBF within 21 d, a bonelike apatite layer formed on their surfaces in 1.5 SBF within 7 d. Silanol groups were also
introduced to the polyethylene (PE) surface through the vapor-phase photografting of
vinyltrimethoxysilane (VTMS) followed by hydrolysis [140]. The photografting process
formed methoxysilyl groups on the PE surface, which were transformed into silanol
groups subsequently by hydrolysis in a 1.0 M HCl solution. The surface-modified PE
was covered with a dense and homogeneous bone-like apatite layer after soaking in
1.5 SBF for 7 d. Modification of alginate with silanol groups, through reacting alginate
with 3-aminopropyltriethoxysilane (APES) that gave silanol groups after hydrolysis, and
calcium ions, through soaking in 1 mole/L CaCl2 solution at 36.5 C for 24 h, could induce
apatite deposition in SBF or 1.5 SBF both on its surface and within its structures [141].
Polymers previously subjected to a glow discharge treatment in oxygen gas to produce
polar groups on the surfaces were soaked in sodium silicate solutions to activate the
surfaces [142]. After soaking in 1.5 SBF for a certain time, bone-like apatite thoroughly
covered the polymer surface. Silicate oligomers with structures such as dimer, linear
trimer and cyclic tetramer contributed most to the apatite nucleation.
EVOH was treated with a silane coupling agent and calcium silicate solutions.
A smooth and uniform bone-like apatite layer was formed on both the EVOH plate
and the EVOH-knitted fibers in SBF within 2 d [143]. When the silane coupling agenttreated EVOH was soaked in a titania sol, Ti-OH groups were introduced on its surface
[144]. A subsequent treatment with an HCl solution crystallized the amorphous titania
to anatase. Apatite was then formed on the anatase-coated copolymer substrates after
soaking in SBF for 2 d.
18
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
A Ca-P layer was coated on the surface of a soft tissue (tendon) through alternative
soaking in a CaCl2 solution and a Na2 HPO4 solution [145]. The Ca2+ ions were suggested
to interact with the tendon surface, most probably with the carboxyl (COOH) functional
groups of collagen, and subsequently form nucleation centers for the Ca-P crystals. Similarly, alternately dipping poly(L-lactic acid) (PLLA) substrates, which were previously
treated with oxygen plasma to produce oxygen-containing functional groups on the surface, in aqueous CaCl2 solution and a K2 HPO4 solution deposited apatite precursors
on the surface. A dense and uniform surface apatite layer formed after immersion for
24 h in SBF [146]. Apatite deposition was observed on both raw silk cloth and sericin
film in 1.5 SBF after 7 d immersion, attributable to the catalytic effect of sericin, which
contained approximately 20 mol% of acidic amino acids that resulted in a surface structure abundant in carboxyl groups [147]. Pre-soaking in a 1 M CaCl2 solution accelerated
apatite deposition on silk fiber. After hydrolysis in water at 37 C for 7 d to produce carboxyl groups on the surface, PLLA fibers also induced apatite deposition in 1.5 SBF at
37 C and pH 7.3 [148]. Carboxyl groups were also introduced on the surface of a chitin
non-woven fabric [149] and gels of chitin and gellan gum [150]. After immersion in a
saturated Ca(OH)2 aqueous solution under nitrogen gas, apatite began to precipitate on
the polymer surface when soaked in SBF.
The biomimetic deposition of apatite on polymers has also been extended to making
organic–inorganic hybrids as bone substitutes. Low temperature formation of HA/poly
(alkyloxybenzoate)phosphazene composites made by the formation of HA through reacting Ca(PO4 2 O with CaHPO4 · 2H2 O in the presence of poly(ethyloxybenzoate) phosphazene (PN-EOB) and poly(propyloxabenzoate)phosphazene (PN-POB) was reported
[151]. The formation of HA through the following acid–base reaction caused the pH of
the solution to increase:
H2 O
2CaHPO4 · 2H2 O + 2Ca4 PO4 2 O −−→ Ca10 PO4 6 OH2 + 4H2 O
(9)
In the presence of PN-EOB or PN-POB, the generation of alkaline conditions promoted partial hydrolysis of the alkyl ester side-groups, resulting in the formation of free
carboxylate groups on the polymer. Calcium ions in the solution cross-linked these carboxylate groups, leading to the formation of a calcium cross-linked polyphosphazene
surface layer, which in turn initiated the deposition of apatite on the surface. A similar in situ hybridization procedure was used to prepare an HA-reinforced chitosan
nanocomposite [152].
Apatite could be deposited on silicone in SBF within 7 d after anatase particles had
been introduced on its surface [153]. Soaking silicone in tetraisopropyltitanate (TIPT) at
30 C for various periods and then, in sequence, in water at 25 C, in a hydrochloride
solution at 25 C, in an ammonia solution at 25 C or hot water at 80 C for 24 d, anatase
precipitated on the silicone surface. The amount of the apatite deposited on silicone
increased with an increasing amount of the precipitated anatase.
Osteopontin is a prominent, highly acidic phosphoglycoprotein of mineralized bone
and dentin. It is believed to bind to extracellular matrix such as type I collagen to
mediate adhesion of cells to mineralize tissues. Osteopontin, which was purified from
bovine milk, was cross-linked or adsorbed to agarose beads [154]. It was then found that
the cross-linked osteopontin induced apatite deposition while the adsorbed osteopontin
failed to induce apatite formation.
Highly porous PTFE membranes, which are used in facial reconstructive surgery, can be
modified to possess bioactivity through the introduction of ionic groups onto its surface.
After introducing a monolayer of monoacryloxyethyl phosphate on the surface, which
did not alter the surface morphology but improved the hydrophilicity of the surface,
apatite-like coatings were obtained on PTFE after soaking in SBF [155].
4.2.3. Electroless Plating
An electroless plating technique has been used to coat an HA layer on high molecular weight polyethylene (HMWPE), starch/ethylene vinyl alcohol (SEVA) blends and
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
19
starch/cellulose acetate (SCA) blends [156]. The bath solution consisted mainly of calcium chloride, sodium hypophosphate, and other additions. PdCl2 was used as a catalyst
and worked as a channel for ionic exchange between the substrate and the solution.
After immersing the polymers in the acidic bath at 60 C or in the basic bath at 80 C
for 1 h, bone-like apatite with in vitro bioactivity was deposited on the polymers
auto-catalytically.
4.3. Biomedical Composites
4.3.1. Ceramic/Ceramic Composites
Zirconia-toughened alumina (ZTA) composites have been developed to replace alumina
ceramics in orthopedic applications requiring high fracture resistance. In ZTA composites,
unstabilized or stabilized zirconia particles were embedded homogeneously throughout the alumina matrix, which enhances the flexural strength, fracture toughness, and
fatigue resistance due to the stress-induced phase transformation that tetragonal zirconia
undergoes into the more stable monoclinic phase. However, these composites are bioinert. A Y-PSZ/Al2 O3 composite, in which the reinforcement is partially stabilized zirconia
containing 2∼4% Y2 O3 , was first subjected to incubation in a sodium silicate solution at
37 C for 7 d to deposit surface functional group of Si-OH on the composite, which initiated apatite nucleation and growth in the subsequent immersing in the 1.5 SBF at 37 C
for another 6 d [157]. The use of sodium silicate solution in place of generally utilized
bioactive glass as a nucleant was suggested to be effective for the formation of a carbonated hydroxyapatite layer on surface of the ZTA composites [157]. Applying a similar
biomimetic method, by soaking Mg-PSZ/Al2 O3 composite in SBF containing either a bed
of wollastonite (CaSiO3 ) ceramics or bioactive glass for 7 d and then in 1.4 SBF for 14 d,
led to the formation of a 15 to 30 m thick apatite. It is found that using wollastonite as
nucleants, the deposited apatite layer was thicker than that using bioactive glass [158].
A dense and homogeneous bone-like apatite layer 20 m in thickness was also achieved
simply by soaking the Mg-PSZ/Al2 O3 composite in SBF for 7 d followed by soaking in
1.4 SBF for 14 d, which provided bioactivity for the ZTA composite [159].
Ceramic/ceramic composites consisting of constituents with distinct bone cell
responses are of interest as a new kind of bone replacement material. Zinc shows a
simulating effect on osteoblastic cell prolifieration and bone formation; therefore, ZnO
has been added to a bovine bone-derived HA. The results indicate that the addition of
5 wt% ZnO increases the microhardness of the sintered composite but negligibly affects
the compressive strength [160]. CaCO3 /HA composites were also prepared by infiltrating
porous HA scaffolds with calcium carbonate slip, followed by sintering. The resultant
composite possessed enhanced elastic and shear moduli compared to pure HA, falling
into the range of human bone [161]. Slosarczyk et al. fabricated a carbon fiber-reinforced
HA composite by hot pressing the two components. Pre-coating of the carbon fiber with
a sol–gel calcium phosphate layer improved the bonding of the reinforcement and the
matrix; thus achieving a homogeneous composite with ideal mechanical properties [162].
The addition of 10 mol% strontium in the HA ceramic was reported to improve not
only the in vitro bioactivity in SBF, but also the cellular attachment, proliferation, and
differentiation [163].
TCP coating has been achieved on a C/C composite through a sonoelectrodeposition
approach. A subsequent soaking of the deposited TCP coating in a mixed solution of NaF,
K2 HPO4 and KH2 PO4 changed the composition to a mixture of HA and F-rich apatite
and also increased the crystallization degree of the coating [164].
4.3.2. Ceramic/Metal Composites
Another composite developed for load-bearing implants is HA/Ti, which combines the
bioactivity of HA and the load-bearing ability of Ti. The composite was produced by hot
pressing a mixture of 50 vol% HA, 40 vol% Ti and 10 vol% bioactive glass [165, 166].
20
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
The bioactive glass was added to aid sintering and to enhance the bioactivity of the
composite. During hot pressing, the following reaction occurred:
Ti + Ca10 PO4 6 OH2 → Ti2 O + CaTiO3 + CaO + Tix Py
(10)
The composite thus had additional CaO, Ti2 O, CaTiO3 , and TiP-like phases. In vitro
bioactivity tests showed that the HA/Ti composite induced apatite deposition in SBF
within 2 h. An apatite layer covered the entire surface after 24 h soaking in SBF. It was
suggested that apatite deposition was initiated through the interaction of Ti2 O with SBF
to form Ti-OH groups and a negatively charged surface [165]. The dissolution of the
CaO phase provided favorable conditions for apatite formation by forming open pores
on the surface and by increasing the degree of supersaturation of SBF with respect to
apatite [166].
4.3.3. Ceramic/Polymer Composites
Ceramic/polymer composites, also called organic–inorganic hybrids, combine the high
strength and wear resistance of ceramics and high toughness and flexibility of polymers
to overcome their individual disadvantages [167–174]. Apart from biological properties, if
mechanical properties similar to those of surrounding tissue are obtained, such biomaterials can minimize the stress-shielding problem among the materials and the surrounding
tissue [167, 175–177]. Therefore, ceramic/polymer composites have been extensively studied and their bioactivity evaluated. Wang provided a comprehensive review of bioactive
composite materials for tissue replacements [167].
Table 4 lists some organic–inorganic hybrids and their in vitro bioactivity. The apatiteforming ability of the hybrids has been ascribed mainly to the bioactive reinforcements,
such as HA, TCP, and Bioglass, and also silica, titania and CaO. The apatite induction
time decreases with increasing contents of the reinforcements [178, 179, 191, 193–195, 197].
The addition of TiO2 in the hybrid of SiO2 -CaO-P2 O5 /PDMS increased readily its apatiteforming ability, due probably also to the increasing specific surface area of the hybrid as
a result of the addition of TiO2 [198]. Enhanced in vitro bioactivity can also be obtained
by modifying the matrix. For example, after treating the HA/chitin composite with an
alkali solution, more D-glucopyranose was introduced in the chitin, which dissolved into
SBF, leading to the increase in hydroxyl concentration in the vicinity of the composite
surface and hence favored apatite deposition [183]. It was also noted that SBF-soaking
can significantly affect the mechanical properties of the hybrids. The storage modulus of
the HA/PHB composite increased initially with immersion time in SBF, due to apatite
Table 4. In vitro bioactivity of some organic–inorganic hybrids.
Matrixa
Reinforcements
Apatite inducer
Shortest Apatite
induction time in SBF
Refs.
PEEK
PHB
Chitin
PHB-PHV
PHB, PHV
PE
PE
PTMO
PTMO
Aerogel
PTMO
PDMS
PCL
ORMOSIL
HA
HA
HA
HA, TCP
HA, TCP
Bioglass®
HA
CaO, TiO2
CaO, SiO2
CaSiO3
TiO2
CaO, SiO2 , TiO2
SiO2
Silica
HA
HA
HA, D-glucopyranose
HA, TCP
HA, TCP
Bioglass®
HA
CaO, TiO2
CaO, SiO2
CaSiO3
TiO2
CaO
SiO2
Ca
3d
1d
2d
7d
1d
3d
7d
1d
12 h
25 d
1d
12 h
3h
3d
[178]
[179, 181, 182]
[183]
[184]
[185, 186]
[187]
[188]
[189, 190]
[191]
[192]
[193]
[194]
[96, 195, 196]
[197]
a
PEEK: Polyetheretherketone; PHB: Polyhydroxybutyrate; PHV: Polyhydroxyvalerate; PE: polyethylene; PTMO:
Poly(tetramethylene oxide); PDMS: Polydimethylsiloxane; PCL: Poly(-caprolactone); ORMOSIL: Organically modified silicate.
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
21
formation on the surface, and decreased after prolonged immersion in SBF, because of
the degradation of the PHB in SBF [179].
A nanocomposite consisting of a bioactive glass nanofiber (58SiO2 -38CaO-4P2 O5 ) and
type I collagen matrix has been reported [180]. The nanofiber was prepared by electrospinning the bioactive glass sol, which combined with self-assembled collagen sol in the
aqueous solution, and then cross-linked to produce the nanocomposite. Such nanocomposites induced the rapid formation of bone-like apatite on the surface when incubated in
SBF. Similarly, the type I collagen matrix dispersed with bioactive glass (CaO-P2 O5 -SiO2 )
and silica micro-particles possessed in vitro apatite-forming ability. Interestingly, the silica
particles and collagen hydrogel alone induced no apatite deposition, which suggests a
possible synergetic effect between collagen and silica on the apatite-forming ability [199].
Eglin et al. [96] used three types of tests to assess in vitro apatite forming ability of
silica/PCL composites: soaking in a static SBF at 37 C, soaking in a dynamic (circulating)
SBF with a flow rate of 0.2 ml/min at 37 C, and alternate soaking in 0.2 M CaCl2 at
37 C for 60 s and 0.12 M Na2 HPO4 at 37 C for 60 s. The results obtained indicated
that even though the dynamic condition best mimics the in vivo condition, both the
static apatite-forming ability test and alternative soaking process are useful in providing
simple methods for gaining quantitative information on the apatite-forming ability of
biomaterials.
4.4. Scaffolds for Tissue Engineering
Tissue engineering requires suitable biodegradable scaffolds in order to produce biological substitutes for reconstructive surgery. In the case of bone tissue engineering, scaffolds
are used either to induce the formation of bone from the surrounding tissue or to act as
a carrier or template for bone cells and/or other agents such as growth factors. Materials
for tissue engineering scaffolds must be biocompatible and bioactive and have a macroporous structure. Bioceramics such as HA [200–203] and -TCP [204, 205], polymers
[206, 207], and organic–inorganic hybrids [181, 200, 208, 211] having porous structures
have been made for bone tissue engineering. Scaffolds made of biodegradable polymers
are very weak in strength. Porous ceramic materials have a high risk of breaking from
normal handling during operations. Composite scaffolds based on polymers and containing bioactive materials such as HA or Bioglass appear to be more suitable for bone
tissue engineering [212].
4.4.1. Ceramics
Two deposition methods, suspension and thermal decomposition, were used to deposit
HA on porous alumina [187]. Reticulated alumina, identified as a strong ceramic with
interconnecting pores, was used as the substrate. The coatings were found to be formed
uniformly on inner pore surfaces of the alumina. In vitro bioactivity was assessed by
immersing the HA-coated alumina in SBF for up to 21 d. It was shown that the in vitro
bioactivity of coated samples was affected by both the crystallinity and specific surface
area of the HA coating. Well-crystallized HA coatings achieved by heat treatment at high
temperatures led to reduced bioactivity. Their stability in SBF is thought to be associated
with the high driving force required for the formation of apatite on their surfaces.
Powder mixtures of HA and phosphate-based glass (CaO · Na2 O · P2 O5 ) were dipcoated onto a porous ZrO2 structure and heat-treated at above 800 C for 2 h in air, which
resulted in a composite coating consisting of HA, TCP, DCP, and residual glass [200].
The adhesion strength of the composite coating increased with increasing glass addition
and decreasing CaO content in glass. The differentiation of cells was improved after such
a coating was applied [200]. In another study, different types of calcium phosphates,
i.e., HA, FA, TCP, and their composites of HA+FA, HA+TCP, were deposited onto the
porous ZrO2 using a powder slurry method [214]. The intermediate layer of FA effectively
suppressed the reaction between the Ca-P layer and the ZrO2 substrate. All coatings
indicated favorable and comparable cell viability. The ALP activity of the cells on pure
HA and HA composite coatings was expressed at higher levels than those on pure FA
22
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
and pure TCP coatings. The ALP levels shown on all coatings were higher than that on
the pure ZrO2 scaffold [214].
Porous CaSiO3 scaffolds were obtained by sintering a ceramic slip-coated polymer
foam at 1350 C. The obtained scaffolds, which were degradable in Ringer’s solution, had
a well-interconnected porous structure with pore sizes ranging from several micrometers
to more than 100 m and porosities of 885 ± 28%. After soaking in SBF for 7 d, HA
was formed on the surface of the scaffolds. Osteoblast-like cells were found to be able to
seed into the CaSiO3 scaffolds, and the proliferation rate and ALP activity of the cells in
the scaffolds improved as compared to the controls [215].
4.4.2. Polymers
The biomimetic apatite deposition process has been used to coat polymer scaffolds with
a layer of bone-like apatite.
Polyactive® is a series of segmented block copolymers consisting of alternating
hydrophilic polyethylene oxdide (PEO) and hydrophobic polybutylene terephthalate
(PBT) segments. Polyactive® 1000/70/30 is a member of the copolymer family with a PBT
content of 30% in weight and a molecular weight of PEO of 1000 g/mol. It could be used
as a bone defect filler for bone tissue engineering due to its bone-bonding ability and
an intermediate degradation rate. To enhance the bioactivity, a continuous biomimetic
ACP or bone-like apatite coating was produced on Polyactive® 1000/70/30, in the form
of either plates or scaffolds, through 1 d soaking in a 5 SBF solution, which was enabled
by bubbling CO2 gas [206].
To make the scaffolds osteoconductive, an accelerated bone-like apatite deposition was
achieved on the surfaces of biodegradable scaffolds of poly(L-lactic acid) (PLLA) and
poly(glycolic acid) (PGA) by soaking them in 5 SBF for only 1 d [212]. An accelerated apatite formation could be very useful for biodegradable polymers, which are used
for bone tissue engineering, as some of these polymers degrade in vitro too fast to be
coated with an apatite layer in a normal biomimetic process. The apatite formed in 5 SBF
was similar in morphology and composition to that formed in the classical biomimetic
process employing SBF. Current work concentrates on the accelerated deposition of an
apatite/collagen composite coating on PLLA scaffolds and a dynamic system for thoroughly coating surfaces of inner pores of the scaffolds.
Carbonxylate groups were introduced on the surface of poly(-caprolactone) (PCL)
plates and scaffolds by treatment with an aqueous NaOH solution [216]. When subsequently soaked in a Ca-P solution, the carbonxylate groups formed apatite nuclei, which
induced a dense and uniform bone-like apatite layer on the PCL surface after incubation
for 24 h in SBF.
Slurry-dipping and electrophoretic deposition techniques were developed to achieve
Bioglass® coatings on a poly(DL-lactide) foam [209]. Stable and homogeneous Bioglass®
coatings as well as infiltration of Bioglass® particles throughout the porous network were
obtained after the process was optimized. The coated foam exhibited in vitro bioactivity
as it induced apatite deposition in SBF after 7 d.
4.4.3. Organic–Inorganic Hybrids
Many organic scaffolds for tissue engineering applications encounter problems of releasing acidic degradation products which lead to inflammatory responses and poor celladhesion due to their hydrophobicity. The incorporation of certain bioactive inorganic
components is effective in controlling these degradation behaviors because of the buffering effect of such alkaline inorganic phases. Composite scaffolds containing bioactive
particles can be produced using a variety of techniques. These bioactive particles are
incorporated in the scaffold and can induce apatite deposition in vitro and in vivo.
A poly(lactic acid)/vaterite composite sponge was prepared through a conventional
particle-leaching technique using sucrose as the sacrificial phase [208, 217, 218]. After
soaking the composite sponge in SBF for 3 d, apatite particles were found covering
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
23
the sponge. In a separate study, a large number of apatite particles were also found
on the pore surface of a composite scaffold consisting of poorly crystallized HA and
PLA, after soaking in SBF for 7 d [181]. The carboxyl (COOH) groups that resulted from
hydrolyzation of PLA are believed to initiate apatite nucleation on the surfaces of the
above-mentioned hybrids.
Spherical Ca-P particles were prepared from highly crystalline feedstock HA particles
by plasma spraying the HA particles onto the surface of ice blocks or into water [219].
The Ca-P particles obtained were highly bioactive due to the presence of ACP in these
particles. Bioactive and biodegradable composite scaffolds produced by incorporating
plasma-sprayed Ca-P particles into a degradable PLA polymer were subjected to in vitro
bioactivity tests. The scaffold exhibited high bioactivity through the formation of apatite
in the SBF within a short period of 24 h [219].
The incorporation of calcium carbonate in vaterite crystalline form in macroporous
poly(DL-lactide) foam achieved a organic–inorganic hybrid scaffold with the ability to
induce apatite deposition after soaking in SBF for 7 d. The hybrid was prepared by dipcoating the organic foam in a slurry of vaterite in methanol or ethanol. The release of
Ca2+ ions as a result of the decomposition of vaterite in the hybrid was suggested to be
responsible for the apatite deposition on the surface [220].
Hirata et al. developed a new biodegradable scaffold. Carbonate apatite (CHA) was
mixed with neutralized collagen gel, which was then lyophilized into sponges in a porous
HA frame ring. After complexing with a bone morphogenic protein (rh-BMP2), the scaffold was implanted beneath the periosteum cranii of rats, and significant new bone was
observed after 4 weeks [221]. Also, in vitro cell culture experiments revealed that the
addition of wollastonite into a PHB-PHV matrix stimulated osteoblasts to proliferate and
differentiate. The hybrids containing Si and Ca as ionic products of wollastonite were
hence more suitable for bone implant and tissue engineering scaffolds [222].
5. CONCLUDING REMARKS
The prerequisite for biomaterials to form a chemical bond to living bone tissue is their
bioactivity; that is, the ability to form a bone-like apatite layer under physiological fluid
when implanted in vivo. The Kokubo’s SBF is generally used to evaluate the bioactivity in vitro. The in vitro apatite deposition of biomaterials is initiated by the surface
functional groups of Si-OH, Ti-OH, COOH, and so on. Other factors contributing to
the in vitro bioactivity include the negatively charged surface, surface micropores, large
specific surface area, certain surface defects, and certain crystal structures. Incorporation of elements such as Ca, P, K and Na in the surface layer also benefits the apatite
deposition through either directly dissolving the Ca and P ions into the surrounding
solution or indirectly exchanging K and Na ions with the H+ ions in the surrounding
solution to increase the supersaturation with respect to apatite. For apatite deposition
on titania gel, a crystalline titania, either in an anatase or rutile structure, is necessary
for apatite-forming within a short time. However, too many factors influence the apatiteforming ability of biomaterials, and some contradictory results have been reported. There
is still much work to do to clarify the exact apatite deposition procedure both in vitro
and in vivo.
Although nearly all the biomaterials capable of forming bone bonding in vivo induce
apatite deposition in SBF within a short time, not all the biomaterials capable of inducing
apatite deposition in SBF are able to develop a direct bone bonding when implanted
in vivo. Thus, although the SBF is simple and helpful to select a bioactive material or
coating, an in vivo animal evaluation seems to be necessary to ensure the bioactivity
before bringing the biomaterials to a clinical application.
Various coating techniques have been applied to provide biomaterials of ceramics,
polymers, and composites with bioactivity. Some coating techniques on polymers have
been utilized successfully to prepare the organic–inorganic hybrids analogous to natural
bone.
24
Bioactive Bioceramic Coatings: Part I. Coatings on Non-Metallic Biomaterials
ACKNOWLEDGMENTS
The authors would like to thank their colleagues and students for useful discussions.
J. M. Wu is grateful to Professor Akiyoshi Osaka for his advice in the research on the
surface treatment of Ti and subsequent apatite formation. M. Wang has benefited greatly
from the interactions with and friendship of Professors Larry Hench, Tadashi Kokubo,
Klaas de Groot, Paul Ducheyne, Akiyoshi Osaka, and Chikara Ohtsuki, who over the
years have provided him with their insights into various bioceramic coatings. M. Wang
also gratefully acknowledges funding from various sources for his research on bioactive
coatings. This chapter was written with the support of research grants from the University of Hong Kong and National Natural Science Foundation of China (NSFC) under
project No. 50502029.
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