Chitosan–nanobioactive glass electrophoretic coatings with bone regenerative and drug delivering potential Kapil D. Patel, ab Ahmed El-Fiqi, ab Hye-Young Lee, ab Rajendra K. Singh, ab Dong-Ae Kim, abc Hae-Hyoung Lee ac and Hae-Won Kim * abc Received 14th June 2012, Accepted 3rd October 2012 DOI: 10.1039/c2jm33830k Nanocomposites with bone-bioactivity and drug eluting capacity are considered as potentially valuable coating materials for metallic bone implants. Here, we developed composite coatings of chitosan (CH)– bioactive glass nanoparticles (BGn) via cathodic electrophoretic deposition (EPD). BGn 50–100 nm in size with aminated surface were suspended with CH molecules at different ratios (5–20 wt% BGn) in aqueous medium, and EPD was performed. Uniform coatings with thicknesses of a few to tens of micrometers were produced, which was controllable by the EPD parameters (voltage, pH and time). Thermogravimetric analysis revealed the quantity of BGn within the coatings that well corresponded to that initially incorporated. Apatite forming ability of the coatings, performed in simulated body fluid, was significantly improved by the addition of BGn. Degradation of the coatings increased with increasing BGn addition. Of note, the degradation profile was almost linear with time; degradation of 5–13 wt% during 1 week became 30–40 wt% after 7 weeks at almost a constant rate. The CH–BGn coatings showed favorable cell adhesion and growth, and stimulated osteogenic differentiation. Drug loading and release capacity of the CH–BGn coatings were performed using the ampicillin antibiotic as a model drug. Ampicillin, initially incorporated within the CH–BGn suspension, was eluted from the coatings continuously over 10–11 weeks, confirming long-term drug delivering capacity. Antibacterial tests also confirmed the effects of released ampicillin using agar diffusion assay against Streptococcus mutants. The CH–BGn may be potentially useful as a coating composition for metallic implants due to the excellent bone bioactivity and cell responses, as well as the capacity for long-term drug delivery. 1. Introduction Commercial pure titanium (CPTi) and its alloys have been extensively used as implants in dental, cranial-maxillary facial reconstruction and orthopedic applications. 1 This is primarily due to their excellent corrosion resistance and biocompatibility, allowing bone-implant integration. 2,3 The biocompatibility of metallic implants can be improved by the surface modification, such as the control over roughness and topography, and the coating with bioactive compositions. While the coatings are the protective layer against corrosion of metals, they impart new compositions to the surface, allowing a large spectrum of possibilities of choosing compositions to trigger proper tissue reactions. A number of coating techniques have been developed, which include plasma spraying, anodic oxidation, sol–gel process, biomimetic coating, sputtering and electrochemical treatment. 4–11 Electrophoretic deposition (EPD) is one of the most useful and effective coating methods available, mainly due to its simplicity and low cost. Advantages also include the possibility of producing a coating layer with high uniformity and variable thickness (0.3–100 mm), the capacity to coat complex shapes, the ease of control over the coating composition and commercial availability. It is possible to apply either an anodic or cathodic treatment depending on the charge of the particles or molecules being deposited. 9 Using the EPD method, a range of composi- tions, including biopolymers, 9,12,13 bioactive ceramics 14,15 and composites 16–21 have been deposited for biomedical implants. Among the compositions, here we focus on biopolymer composites with bioactive inorganic nanoparticles. In fact, there has been significant attention to produce biopolymer composite coatings with inorganic particles by the EPD method. 17–23 Inor- ganic particles, including hydroxyapatite (HA), carbon nano- tube, silica, and their combinations, introduced into the polymeric solutions, were enabled to form co-deposits by the EPD process. Among the biopolymer sources, chitosan (CH) has been widely used, as it is biocompatible and degradable and is a Institute of Tissue Regeneration Engineering (ITREN), Dankook University, South Korea. E-mail: [email protected]; Fax: +82 41 550 3085; Tel: +82 41 550 3081 b Department of Nanobiomedical Science & WCU Research Center, Dankook University Graduate School, South Korea c Department of Biomaterials Science, College of Dentistry, Dankook University, South Korea This journal is ª The Royal Society of Chemistry 2012 J. Mater. Chem., 2012, 22, 24945–24956 | 24945 Dynamic Article Links C < Journal of Materials Chemistry Cite this: J. Mater. Chem., 2012, 22, 24945 www.rsc.org/materials PAPER Downloaded by Dankook University on 23 November 2012 Published on 03 October 2012 on http://pubs.rsc.org | doi:10.1039/C2JM33830K View Article Online / Journal Homepage / Table of Contents for this issue
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Chitosan–nanobioactive glass electrophoretic coatings with bone regenerativeand drug delivering potential
Kapil D. Patel,ab Ahmed El-Fiqi,ab Hye-Young Lee,ab Rajendra K. Singh,ab Dong-Ae Kim,abc Hae-Hyoung Leeac
and Hae-Won Kim*abc
Received 14th June 2012, Accepted 3rd October 2012
DOI: 10.1039/c2jm33830k
Nanocomposites with bone-bioactivity and drug eluting capacity are considered as potentially valuable
coating materials for metallic bone implants. Here, we developed composite coatings of chitosan (CH)–
bioactive glass nanoparticles (BGn) via cathodic electrophoretic deposition (EPD). BGn 50–100 nm in
size with aminated surface were suspended with CH molecules at different ratios (5–20 wt% BGn) in
aqueous medium, and EPD was performed. Uniform coatings with thicknesses of a few to tens of
micrometers were produced, which was controllable by the EPD parameters (voltage, pH and time).
Thermogravimetric analysis revealed the quantity of BGn within the coatings that well corresponded to
that initially incorporated. Apatite forming ability of the coatings, performed in simulated body fluid,
was significantly improved by the addition of BGn. Degradation of the coatings increased with
increasing BGn addition. Of note, the degradation profile was almost linear with time; degradation of
5–13 wt% during 1 week became 30–40 wt% after 7 weeks at almost a constant rate. The CH–BGn
coatings showed favorable cell adhesion and growth, and stimulated osteogenic differentiation. Drug
loading and release capacity of the CH–BGn coatings were performed using the ampicillin antibiotic as
a model drug. Ampicillin, initially incorporated within the CH–BGn suspension, was eluted from the
coatings continuously over 10–11 weeks, confirming long-term drug delivering capacity. Antibacterial
tests also confirmed the effects of released ampicillin using agar diffusion assay against Streptococcus
mutants. The CH–BGn may be potentially useful as a coating composition for metallic implants due to
the excellent bone bioactivity and cell responses, as well as the capacity for long-term drug delivery.
1. Introduction
Commercial pure titanium (CPTi) and its alloys have been
extensively used as implants in dental, cranial-maxillary facial
reconstruction and orthopedic applications.1 This is primarily
due to their excellent corrosion resistance and biocompatibility,
allowing bone-implant integration.2,3 The biocompatibility of
metallic implants can be improved by the surface modification,
such as the control over roughness and topography, and the
coating with bioactive compositions. While the coatings are the
protective layer against corrosion of metals, they impart new
compositions to the surface, allowing a large spectrum of
possibilities of choosing compositions to trigger proper tissue
reactions. A number of coating techniques have been developed,
which include plasma spraying, anodic oxidation, sol–gel
aInstitute of Tissue Regeneration Engineering (ITREN), DankookUniversity, South Korea. E-mail: [email protected]; Fax: +82 41 5503085; Tel: +82 41 550 3081bDepartment of Nanobiomedical Science & WCU Research Center,Dankook University Graduate School, South KoreacDepartment of Biomaterials Science, College of Dentistry, DankookUniversity, South Korea
This journal is ª The Royal Society of Chemistry 2012
process, biomimetic coating, sputtering and electrochemical
treatment.4–11
Electrophoretic deposition (EPD) is one of the most useful and
effective coating methods available, mainly due to its simplicity
and low cost. Advantages also include the possibility of
producing a coating layer with high uniformity and variable
thickness (0.3–100 mm), the capacity to coat complex shapes, the
ease of control over the coating composition and commercial
availability. It is possible to apply either an anodic or cathodic
treatment depending on the charge of the particles or molecules
being deposited.9 Using the EPD method, a range of composi-
tions, including biopolymers,9,12,13 bioactive ceramics14,15 and
composites16–21 have been deposited for biomedical implants.
Among the compositions, here we focus on biopolymer
composites with bioactive inorganic nanoparticles. In fact, there
has been significant attention to produce biopolymer composite
coatings with inorganic particles by the EPD method.17–23 Inor-
ganic particles, including hydroxyapatite (HA), carbon nano-
tube, silica, and their combinations, introduced into the
polymeric solutions, were enabled to form co-deposits by the
EPD process. Among the biopolymer sources, chitosan (CH) has
been widely used, as it is biocompatible and degradable and is
Fig. 10 (a) Na–ampicillin was incorporated within the coating layer during the EPD process and the release profile was observed for periods of up to
10–11 weeks. CH and 10BGn coatings were tested representatively. Na–ampicillin was added to the EPD solution in concert with CH or CH–10BG
nanoparticles; at two different concentrations (low 5 mg and high 10 mg; CH ¼ 100 mg and BG ¼ 10 mg). After the EPD process (40 kV, 5 min), the
coating layer was gently washed and dried and the sample was immersed in PBS at 37 �C for different time points, prior to an assay for the ampicillin
release amount using a UV-vis spectrophotometer. A continuous and highly sustained release for up to the period tested (10–11 weeks) was recorded.
Data well fitted according to the combined model of the zero-order model (initial stage) and Ritger–Peppas empirical equation (later stage), and
parameters are summarized in Table 1. (b) Antibacterial tests of the ampicillin-loaded 10BGn coating against Streptococcus mutants using an agar
diffusion plate. Antibacterial effective zone was formed around the ampicillin-loaded coating at 24 h and was maintained and even increased for up to 5
days (time point for the bacteria lifespan), which was however not observed in the coating without ampicillin loading, confirming the efficacy of the drug
delivery through the composite coating layer. Representative images of the agar diffusion test are shown for comparison purpose (1 and 5 days of
ampicillin-added 10BGn vs. 1 day of ampicillin-free 10BGn).
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and MN are the absolute amount of drug released at time t and
infinite time (N), respectively, and K0 and K are released rate
constants for each equation, incorporating structural and
geometric characteristics of the drug delivery device, and n is the
released exponent, indicative of the drug release mechanism. The
parameters determined from the curve fittings are summarized in
Table 1. Although the models are simplified forms without
considering the moving boundary problems as our coatings are
degradable and thus do not preserve constant volumes, they
should allow the interpretation of the drug release kinetics in a
much easier and simpler way, as have generally been carried out
elsewhere.58–61 The initial stage was shown to follow well the
linearity, with the R2 value lower than 0.99. The 2nd stage also
showed a power exponent of 0.44, 0.37 and 0.38, for CH (high),
10BGn (low) and 10BGn (high) coating, respectively, values
lower than 0.5 (an index of the diffusion-controlled process),
suggesting the stage is a sort of anomalous diffusion-controlled
(slight deviation from Fickian diffusion-controlled) release
phenomenon, which has been reported elsewhere, systems such
Table 1 Summary of release-model parameters (K0, K, and n), definingthe release mechanism of the drug from the coatings. Linear release withzero-order kinetics; Mt/MN ¼ K0t at the 1st stage, and then at the 2nd
stage with a power law relationship provided empirically by Ritger–Peppas; Mt/MN ¼ Ktn
Model Parameter
Coating sample with ampicillin
CH (low) 10BGn (low) 10BGn (high)
Zero-order model K0 2.82 3.38 4.16Ritger–Peppasempirical model
K 17.5 22.6 35.2n 0.44 0.37 0.38
24952 | J. Mater. Chem., 2012, 22, 24945–24956
as hydrogels, swelling polymers and semi-interpenetrating
networks.62–64
The initial drug release may be mainly ascribed to the degra-
dation (surface erosion) of the coating layer as the surface-
reaction (erosion) process has a linear dependence on time;
although a level of diffusion out of drug also occurred, the
degradation will be the rate determinant. The slightly faster
release of the drug in the 10BGn was also associated with the
more rapid coating degradation in the sample. However, after a
certain period (�14 days), when a depletion zone of drug was
created at the surface region, drugs below the zone could be
released mainly by a diffusion through the surface coating layer,
which would be evident as the curved parabolic-like pattern at a
longer period. Although the surface erosion is processed, and at
this step the drugs existing at the eroding surface should be
released, drugs in the deeper region could still be diffusing out
through the barrier of the coating layer. As the drug release
process results from a complex of the coating degradation and
the diffusion through coating layer, the outcome pattern with
respect to time will not be an abrupt transition, but rather might
be a smooth pattern. Coating degradability, interactions between
drug molecules and coating materials, and permeability or
diffusivity of drugs through the coating can significantly influ-
ence the drug release profile. These aspects need to be considered
carefully in the design of coatings to control the drug release
profile. Although this release pattern may not be applicable in
parallel to all other types of drugs, because of the difference in
the drug size and interactions with coating materials, particularly
for small hydrophilic (or possibly anionic) drugs this long-term
(over 2–3 months) sustained release can be envisioned. As the
ampicillin molecules are anionic-charged, a sort of weak charge–
charge interactions is possible with the BGn or CH molecules
This journal is ª The Royal Society of Chemistry 2012
characteristic wavelength, 230 nm. A series of standard Na–
ampicillin solutions in deionized water (10–100 mg ml�1) were
prepared to obtain a linear calibration curve (r2 ¼ 0.99) that
obeys Beer’s law A ¼ abc, where A is the absorbance, a is a
constant known as absorbtivity coefficient, c is the concentra-
tion, and b is the cell bath length, which is constant. To eliminate
any possible interference of the degraded products, blank solu-
tions for the UV-spec. assay was prepared, by collecting solu-
tions from the coatings free of ampicillin at the same incubation
time as the drug-eluting period.
Antibacterial effects of the ampicillin released from the coat-
ings were investigated by means of an agar diffusion test against
Streptococcus mutants (from ATCC, USA). Coated samples with
or without ampicillin were used. After spreading 100 ml aliquot
of Streptococcus mutants directly onto the agar plate and incu-
bated overnight at 37 �C, each sample was placed onto the agar
plate, and the inhibitory zone formed by the released ampicillin
from the coating layer was visualized during periods for up to 5
days with 24 h interval.
4.7. Statistics
Data are presented as the mean � standard deviation and the
differences between groups were compared using a Student’s
t-test. Statistical significance was considered at p < 0.05 and
p < 0.01.
Acknowledgements
This work was supported by the Priority Research Centers
Program (no. 2009-0093829) and WCU program (no. R31-
10069) through the NSF, funded by the MEST, Republic of
Korea. Authors also thank the assistance of Mrs Hwang KH in
the cellular assays.
References
1 D. M. Brunette, P. Tengvall, M. Textor and P. Thomsen, Titanium inMedicine, Springer Verlag, Berlin, 2001.
2 P. Lutjering, J. C. Williams and A. Gysler, Microstructure andMechanical Properties of Titanium Alloys, in, Microstructure andProperties of Materials, ed. Li J. C. M., World Scientific, Singapore,2000.
3 R. Br�anemark, P. I. Br�anemark, B. Rydevik and R. R. Myers,Osseointegration in skeletal reconstruction and rehabilitation. Areview:, J. Rehabil. Res. Dev., 2001, 38(2), 175–181.
4 M. Geetha, A. K. Singh, R. Asokamani and A. K. Gogia, Ti basedbiomaterials, the ultimate choice for orthopaedic implants – areview, Prog. Mater. Sci., 2009, 54, 397–425.
5 D. L. Cochran, R. K. Schenk, A. Lussi, F. I. Higginbottom andD. Buser, Bone response to unloaded and loaded titanium implantswith a sandblasted and acid-etched surface: a histometric study inthe canine mandible, J. Biomed. Mater. Res., 1998, 40, 1–11.
6 H. Kurzweg, R. B. Heimann and T. Troczynski, Development ofplasma sprayed bioceramic coatings with bond coats based ontitania and zirconia, Biomaterials, 1998, 19, 1507–1515.
7 S. H. Lee, H. W. Kim, E. J. Lee, L. H. Li and H. E. Kim,Hydroxyapatite–TiO2 coating on Ti implants, J. Biomater. Appl.,2006, 20, 195–208.
8 L. H. Li, Y. M. Kong, H. W. Kim, Y. W. Kim, H. E. Kim, S. J. Heoand J. Y. Koak, Improved biological performance of Ti implants dueto surface modification by micro-arc oxidation, Biomaterials, 2004,25, 2867–2875.
9 P. Sarkar and P. S. Nicholson, Electrophoretic deposition (EPD):mechanism, kinetics and application to ceramics, J. Am. Ceram.Soc., 1996, 79, 1987–2002.
This journal is ª The Royal Society of Chemistry 2012
10 D. S. Couto, N. M. Alves and J. F. Mano, Nanostructured multilayercoatings combining chitosan with bioactive glass nanoparticles, J.Nanosci. Nanotechnol., 2009, 9, 1741–1748.
11 S. H. Jun, E. J. Lee, S. W. Yook, H. E. Kim and H. W. Kim, Abioactive coating of a silica xerogel–chitosan hybrid on titanium bya room temperature sol–gel process,Acta Biomater., 2010, 6, 302–307.
12 A. Simichi, F. Pishbin and A. R. Boccaccini, Electrophoreticdeposition of chitosan, Mater. Lett., 2009, 63, 2253–2256.
13 L. Besra and M. Liu, A review on fundamentals and applications ofelectrophoretic deposition (EPD), Prog. Mater. Sci., 2007, 52, 1–61.
14 I. Zhitomirsky, Electrophoretic deposition of organic–inorganicnanocomposites, J. Mater. Sci., 2006, 41, 8186–8195.
15 Z. Zhang, T. Jiang, K. Ma, X. Cai, Y. Zhuo and Y. Wang, Lowtemperature electrophoretic deposition of porous chitosan–silkfibroin composite coating for titanium biofunctionalization, J.Mater. Chem., 2011, 21, 7705–7713.
16 F. Pishbin, A. Simchi, M. P. Ryan and A. R. Boccaccini,Electrophoretic deposition of chitosan–45S5 Bioglass� compositecoatings for orthopaedic applications, Surf. Coat. Technol., 2011,205, 5260–5268.
17 K. Rezwan, Q. Z. Chen, J. J. Blaker and A. R. Boccaccini,Biodegradable and bioactive porous polymer–inorganic compositescaffolds for bone tissue engineering, Biomaterials, 2006, 27, 3413–3431.
18 K. Grandfield, F. Sun, P. M. Fitz, M. Cheong and I. Zhitomirsky,Electrophoretic deposition of polymer–carbon nanotube–hydroxyapatite composites, Surf. Coat. Technol., 2009, 203, 1481–1487.
19 T. Casagrande, P. Imin, F. Cheng, G. A. Botton, I. Zhitomirsky andA. Adronov, Synthesis and electrophoretic deposition of single-walledcarbon nanotube complexes with a conjugated polyelectrolyte, Chem.Mater., 2010, 22, 2741–2749.
20 X. Pang and I. Zhitomirsky, Electrodeposition of compositehydroxyapatite–chitosan films, Mater. Chem. Phys., 2005, 94, 245–251.
21 K. Grandfield and I. Zhitomirsky, Electrophoretic deposition ofcomposite hydroxyapatite–silica–chitosan coatings, Mater. Charact.,2008, 59, 61–67.
22 H. Yi, L. Q. Wu, W. E. Bentley, R. Ghodssi, G. W. Rubloff,J. N. Culver and G. F. Payne, Biofabrication with chitosan,Biomacromolecules, 2005, 6, 2881–2894.
23 D. Zhitomirsky, J. A. Roether, A. R. Boccaccini and I. Zhitomirsky,Electrophoretic deposition of bioactive glass–polymer compositecoatings with and without HA nanoparticle inclusions forbiomedical applications, J. Mater. Process. Technol., 2009, 209,1853–1860.
24 M. Dash, F. Chiellini, R. M. Ottenbrite and E. Chiellini, Chitosan-Aversatile semi-synthetic polymer in biomedical applications, Prog.Polym. Sci., 2011, 36, 981–1014.
25 J. V. V. Pamela, W. T. M. Howard, P. D. Stephen, M. Lois, W. Binand H. W. Paul, Evaluation of the biocompatibility of a chitosanscaffold in mice, J. Biomed. Mater. Res., 2002, 59, 585–590.
26 K. P. H€ogg�ard, K. M. V�arum, M. Issa, S. Danielsen,B. E. Christensen, B. T. Stokke and P. Artursson, Improvedchitosan-mediated gene delivery based on easily dissociated chitosanpolyplexes of highly defined chitosan oligomers, Gene Ther., 2004,11, 1441–1452.
27 M. K. Lee, S. K. Chun, W. J. Choi, J. K. Kim, S. H. Choi, A. Kim,K. Oungbho, J. S. Park, W. S. Ahn and C. K. Kim, The use ofchitosan as a condensing agent to enhance emulsion-mediated genetransfer, Biomaterials, 2005, 26, 2147–2156.
28 R. A. A. Muzzarelli, G. Biagini, A. DeBenedittis, P. Mengucci,G. Majni and G. Tosi, Chitosan–oxychitin coatings for prostheticmaterials, Carbohydr. Polym., 2001, 45, 35–41.
29 L. L. Hench, R. J. Splinter, W. C. Allen and T. K. Greenlee, Bondingmechanisms at the interface of ceramic prosthetic materials, J.Biomed. Mater. Res. Symp., 1971, 2, 117–141.
30 L. L. Hench, Bioceramics: from concept to clinic, J. Am. Ceram. Soc.,1991, 74, 1487–1570.
31 L. L. Hench and €O. Andersson, Bioactive Glasses, in An Introductionto Bioceramics, ed. L. L. Hench and J. Wilson, World Scientific,Singapore, 1993, pp. 41–62.
32 M. M. Pereira, A. E. Clark and L. L. Hench, Calcium phosphateformation on sol–gel-derived bioactive glasses in vitro, J. Biomed.Mater. Res., 1994, 28, 693–698.
33 H. S. Yun, S. E. Kim and Y. T. Hyeon, Design and preparation ofbioactive glasses with hierarchical pore networks, Chem. Commun.,2007, 2139–2141.
34 H. W. Kim, H. E. Kim and J. C. Knowles, Potential of bioactive glassnanofiber as a next generation biomaterial, Adv. Funct. Mater., 2006,16, 1529–1536.
35 T. Waltimo, T. J. Brunner, M. Vollenweider, W. J. Stark andM. Zehnder, Antimicrobial effect of nanometric bioactive glass45S5, J. Dent. Res., 2007, 86, 754–757.
36 H. W. Kim, J. H. Song and H. E. Kim, Bioactive glass nanofiber –collagen nanocomposite as a novel bone regeneration matrix, J.Biomed. Mater. Res., Part A, 2006, 79, 698–705.
37 A. R. Boccaccini, M. Erol, W. J. Stark, D. Mohn, Z. Hong andJ. F. Mano, Polymer–bioactive glass nanocomposites forbiomedical applications: a review, Compos. Sci. Technol., 2010, 70,1764–1776.
38 M. Peter, N. S. Binulal, S. Soumya, S. V. Nair, T. Furuike, H. Tamuraand R. Jayakumar, Nanocomposite scaffolds of bioactive glassceramic nanoparticles disseminated chitosan matrix for tissueengineering applications, Carbohydr. Polym., 2010, 79, 284–289.
39 F. Pishbin, A. Simchi, M. P. Ryan and A. R. Boccaccini, A study ofthe electrophoretic deposition of Bioglass� suspensions using theTaguchi experimental design approach, J. Eur. Ceram. Soc., 2010,30, 2963–2970.
40 R. F. S. Lenza and W. L. Vasconcelos, Structural evolution of silicasols modification with formamide, Mater. Res., 2001, 4, 175–179.
41 I. Manjubala, S. Scheler, J. Bossert and K. D. Jandt, Mineralisationof chitosan scaffolds with nano-apatite formation by doublediffusion technique, Acta Biomater., 2006, 2, 75–84.
42 H. Shen, X. Hu, F. Yang, J. Bei and S. Wang, Combining oxygenplasma treatment with anchorage of cationized gelatin forenhancing cell affinity of poly(lactide-co-glycolide), Biomaterials,2007, 28, 4219–4230.
43 S. E. Bae, J. Choi, Y. K. Joung, K. Park and D. K. Han, Controlledrelease of bone morphogenetic protein (BMP)-2 from nanocomplexincorporated on hydroxyapatite-formed titanium surface, J.Controlled Release, 2012, 160, 676–684.
44 J. K. Leach, D. Kaigler, Z. Wang, H. K. Paul and D. J. Mooney,Coating of VEGF-releasing scaffolds with bioactive glass forangiogenesis and bone regeneration, Biomaterials, 2006, 27, 3249–3255.
45 H. W. Kim, H. H. Lee and G. S. Chun, Bioactivity and osteoblastresponses of novel biomedical nanocomposites of bioactive glassnanofiber filled poly(lactic acid), J. Biomed. Mater. Res., Part A,2008, 85, 651–663.
46 R. Zhang and P. X. Ma, Biomimetic polymer–apatite compositescaffolds for mineralized tissue engineering, Macromol. Biosci.,2004, 4, 100–111.
47 I. Yamaguchi, K. Tokuchi, H. Fukuzaki, Y. Koyama, K. Takakuda,H.Monma and J. Tanaka, Preparation andmicrostructure analysis ofchitosan–hydroxyapatite nanocomposites, J. Biomed. Mater. Res.,Part A, 2001, 55, 20–27.
48 R. K. Singh and A. Srinivasan, Apatite-forming ability and magneticproperties of glass-ceramics containing zinc ferrite and calciumsodium phosphate phases, Mater. Sci. Eng., C, 2010, 30, 1100–1106.
49 I. Manjubala, S. Scheler, J. Bossert and K. D. Jandt, Mineralisationof chitosan scaffolds with nano-apatite formation by doublediffusion technique, Acta Biomater., 2006, 2, 75–84.
50 B. Dorj, J. H. Park and H. W. Kim, Robocasting chitosan–nanobioactive glass dual-pore structured scaffolds for boneengineering, Mater. Lett., 2012, 73, 119–122.
51 S. Maeno, Y. Niki, H. Matsumoto, H. Morioka, T. Yatabe,A. Funayama, Y. Toyama, T. Taguchi and J. Tanaka, The effect ofcalcium ion concentration on osteoblast viability, proliferation and
24956 | J. Mater. Chem., 2012, 22, 24945–24956
differentiation in monolayer and 3D culture, Biomaterials, 2005, 26,4847–4855.
52 M. Y. Shie, S. J. Ding and H. C. Chang, The role of silicon inosteoblast-like cell proliferation and apoptosis, Acta Biomater.,2011, 7, 2604–2614.
53 S. A. Oh, S. H. Kim, J. E. Won, J. J. Kim, U. S. Shin and H. W. Kim,Effects on growth and osteogenic differentiation of mesenchymalstem cells by the zinc-added sol–gel bioactive glass granules, J.Tissue Eng., 2010, 2010, 475260.
54 X. Huang and C. S. Brazel, On the importance and mechanisms ofburst release in matrix-controlled drug delivery systems, J.Controlled Release, 2001, 73, 121–136.
55 M. C. Berg, L. Zhai, R. E. Cohen andM. F. Rubner, Controlled drugrelease from porous polyelectrolyte multilayers, Biomacromolecules,2006, 7, 357–364.
56 L. Peng, A. D. Mendelsohn, T. J. LaTempa, S. Yoriya, C. A. Grimesand T. A. Desai, Long-term small molecule and protein elution fromTiO2 nanotubes, Nano Lett., 2009, 9, 1932–1936.
57 P. L. Ritger and N. A. Peppas, A simple equation for description ofsolute release I. Fickian and non-Fickian release from non-swellabledevices in the form of slabs, spheres, cylinders or discs, J.Controlled Release, 1987, 5, 23–36.
58 C. Strobel, N. Bormann, A. Kadow-Romacker, G. Schmidmaier andB. Wildemann, Sequential release kinetics of two (gentamicin andBMP-2) or three (gentamicin, IGF-I and BMP-2) substances from aone-component polymeric coating on implants, J. ControlledRelease, 2011, 156, 37–45.
59 A. M. Young and S. M. Ho, Drug release from injectablebiodegradable polymeric adhesives for bone repair, J. ControlledRelease, 2008, 127, 162–172.
60 B. Jeong, Y. H. Bae, S. H. Lee and S. W. Kim, Biodegradable blockcopolymers as injectable drug-delivery systems, Nature, 1997, 388,860–862.
61 P. R. Chen, M. H. Chen, F. H. Lin and W. Y. Su, Releasecharacteristics and bioactivity of gelatin–tricalcium phosphatemembranes covalently immobilized with nerve growth factors,Biomaterials, 2005, 26, 6579–6587.
62 L. Serra, J. Domenechc and N. A. Peppas, Drug transportmechanisms and release kinetics from molecularly designedpoly(acrylic acid-g-ethylene glycol) hydrogels, Biomaterials, 2006,27, 5440–5451.
63 P. L. Ritger and N. A. Peppas, A simple equation for description ofsolute release II. Fickian and anomalous release from swellabledevices, J. Controlled Release, 1987, 5, 37–42.
64 Y. Fu and J. W. Kao, Drug release kinetics and transport mechanismsfrom semi-interpenetrating networks of gelatin and poly(ethyleneglycol) diacrylate, Pharm Res., 2009, 26, 2115–2124.
65 A. L. Oliveira, P. B. Malafaya and R. L. Reis, Sodium silicate gel as aprecursor for the in vitro nucleation and growth of a bone-like apatitecoating in compact and porous polymeric structures, Biomaterials,2003, 24, 2575–2584.
66 M. Lebourg, J. S. Anton and J. L. G. Ribelles, Characterization ofcalcium phosphate layers grown on polycaprolactone for tissueengineering purposes, Compos. Sci. Technol., 2010, 70, 5182–5190.
67 H. W. Kim, J. C. Knowles and H. E. Kim, Hydroxyapatite–poly(3-caprolactone) composite coatings on hydroxyapatite porous bonescaffold for drug delivery, Biomaterials, 2004, 25, 1279–1287.
68 E. L. Hedberg, C. K. Shin, J. J. Lemoine, M. D. Timmer,M. A. Lieschner, J. A. Jansen and A. G. Mikos, In vitrodegradation of porous poly(propylene fumarate)–poly(DL-lactic-co-glycolic acid) composite scaffolds, Biomaterials, 2005, 26, 3215–3225.
69 Y. Hu, K. Cai, Z. Luo, R. Zhang, L. Yang, L. Deng and K. D. Jandt,Mineral-coated polymer microspheres for controlled protein bindingand release, Adv. Mater., 2009, 21, 1960–1963.
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