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Coatings of titanium substrates with xCaO·(1 − x)SiO2 sol–gel materials: characterization, bioactivity and biocompatibility evaluation
Materials Science and Engineering C 58 (2016) 846–851
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Coatings of titanium substrates with xCaO·(1 − x)SiO2 sol–gel materials: characterization, bioactivity and biocompatibility evaluation M. Catauro ⁎, F. Papale, F. Bollino Department of Industrial and Information Engineering, Second University of Naples, Via Roma 29, 81031 Aversa, Italy
a r t i c l e
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Article history: Received 9 February 2015 Received in revised form 24 August 2015 Accepted 7 September 2015 Available online 12 September 2015 Keywords: Bioactivity Calcium silicate Glass Sol–gel method
a b s t r a c t The objective of this study has been to develop low temperature sol–gel coatings to modify the surface of commercially pure titanium grade 4 (a material generally used in dental application) and to evaluate their bioactivity and biocompatibility on the substrate. Glasses of composition expressed by the following general formula xCaO·(1 − x)SiO2 (0.0 b x b 0.60) have been prepared by means of the sol–gel route starting from tetraethyl orthosilicate and calcium nitrate tetrahydrate. Those materials, still in the sol phase, have been used to coat titanium substrates by means of the dip-coating technique. Fourier transform infrared spectroscopy (FTIR) and X-ray diffraction (XRD) allowed the materials to be characterized and a microstructural analysis of the coatings obtained was performed using scanning electron microscopy (SEM). The potential applications of the coatings in the biomedical field were evaluated by bioactivity and biocompatibility tests. The coated titanium was immersed in simulated body fluid (SBF) for 21 days and the hydroxyapatite deposition on its surface was subsequently evaluated via SEM–EDXS analysis, as an index of bone-bonding capability. To investigate cell-material interactions, mouse embryonic fibroblast cells (3 T3) were seeded onto the specimens and the cell viability was evaluated by a WST-8 assay. © 2015 Published by Elsevier B.V.
1. Introduction Continuous improvements in medicine and life quality are gradually leading to a senescence of the population in industrialized countries. This trend implies that a considerable portion of the population suffers from diseases caused by a progressive deterioration of the musculoskeletal system. The last decade has been characterized by many studies on biomaterials such as to allow the replacement the damaged tissues, to build long-life prostheses or to release drugs [1–4]. Metallic materials, thanks to their excellent mechanical properties, are used today in orthopedic and dental applications where the implants are subject to severe stresses. However, metallic materials have several significant disadvantages such as the tendency to release ions (toxic) when they come in contact with body fluids and the lack of integration in bone [5–9]. As a consequence, the formation of peri-implant fibrosis may occur, thus inducing prosthesis mobilization and then the reduction of its performance. Ceramic materials show a capability to bond spontaneously and to integrate with bones in living tissues, therefore stimulating osteogenesis [5,10–12]. The clinical use of such biomaterials is limited by their poor mechanical properties ⁎ Corresponding author. E-mail address: [email protected] (M. Catauro).
http://dx.doi.org/10.1016/j.msec.2015.09.033 0928-4931/© 2015 Published by Elsevier B.V.
which hamper monolithic glasses from being used in load-bearing applications, where metallic alloys are still the materials of choice. The coating of metal implants with ceramic materials has enabled corrosion [13] and ion release to be reduced, and bioactivity and biocompatibility [14] to be improved, maintaining the metal's mechanical properties [15]. A common technique used to obtain ceramic or glassy coatings on substrates, even with complex shapes, is the sol–gel dip coating process. Sol–gel is the method used to synthesize ceramic and/or glass materials at a relatively low temperature. Usually the chemical reagents used in the sol preparation are metal organic compounds such as metal alkoxides [M(OR)n], where M represents a network-forming element (such as Al, B, Si, Ti, Zr, etc), and R is an alkyl group. The conversion of the system from sol to gel is the result of hydrolysis and polycondensation reactions of the metal alkoxide precursors. The hydrolysis generates hydroxyl groups which tend to establish new bonds (by means of oxygen bridges) through polycondensation reactions. This process has several advantages when compared to traditional synthesis, such as the capability to obtain homogeneous materials with fine control of the chemical composition; it also allows the densification temperature of the layer to be decreased and it requires less equipment, and is thus cheaper. Materials prepared by the sol–gel process have shown to be more bioactive than those with the same composition but prepared with different methods. In fact, surface analysis of sol–gel materials
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and coatings has shown the presence of hydroxyl groups that can promote the nucleation of calcium and phosphate and, therefore, the osteointegration when those materials are implanted. The sol–gel technique leads easily to various coating techniques such as dip-coating, spin-coating and spray [16–18] . In the literature the use of the sol–gel dip coating technique in dental and orthopedic fields is reported for the surface modifications of metals (e.g. titanium-alloys) in order to reduce corrosion and ion release, and to improve their biological properties [6,19–21]. Bioactive and thin coatings (about 70 to 1000 nm depending on the number of coating layers) can promote an early attachment of the bone to the implant. This increased bond strength and accelerated osteoconductivity are very important features in clinical applications, as the success of an implant is also a function of the initial bone attachment. The dip coating method requires the substrate to be dipped into the ‘sol’ and withdrawn at a constant speed to allow the sol drainage and its instantaneous gelation. Many parameters influence the thickness and the morphology of the obtained coating, such as the withdrawal speed, sol viscosity and drying treatment [22]. The sol–gel dip coating process is an advantageous technique for coating preparation due to the higher purity and homogeneity, lower processing temperatures, reduced thickness, simple and cheap method of preparation and ability to form a physically and chemically uniform coating over complex shapes. The aim of this study is to prepare materials via the sol–gel method with the composition of the adapted glasses expressed by the following general formula: xCaO·(1 − x)SiO2 with x = 0.00; 0.30; 0.40; 0.50; 0.60 film on Titanium grade 4 (Ti-4), to improve their biocompatibility and bioactivity. Ti-4 is a metallic material widely used in implants for orthopedic, dental and orthodontic wires. The coating of these materials allow the main drawback to be overcome which limits their use as implants in many applications [23]. Although titanium and its alloys are able to form spontaneously a stable and protective TiO2 surface layer [24], corrosion can take place when this film is removed due to local mechanical abrasion [23]. Different strategies can be followed to avoid the early failure of the implants due to wear and corrosion and to extend the prosthese lifetime. Many studies describe the use of nitric acid passivation protocols in order to increase the thickness of the oxide films [23,25]. Other studies report that a protective coating can be applied on titanium substrates for this purpose. In particular, the use of sol–gel coatings for corrosion protection which also leads to an improvement of the substrate's bioactivity and biocompatibility was proposed [13,26]. The choice of the material's constituents used to make the coating has been targeted to obtain layers with high bioactivity and biocompatibility. It is known [27] that calcium silicate amorphous glasses, as bulks, are able to form strong bonds to living bones and tissues. They stimulate osteogenesis via both their dissolution and degradation products. Indeed, soluble silica and calcium ions are proposed to both activate and stimulate osteoprogenitor cells at the implant site, promoting bone tissue growth. Different formulations of such materials, which differ in Ca-content, were prepared in order to study the influence of the calcium amount on the biological properties of the coating. In the literature [28,29], indeed, it has been proven that the presence of cations in a material can stimulate hydroxyapatite nucleation when it is soaked in simulated body fluid (SBF). In the present work, the bioactivity and biocompatibility of the prepared coatings have been evaluated by bioactivity tests and WST-8 assays on murine fibroblasts (3 T3) seeded on the layers. In order to assess that the good biological properties detected for calcium silicate have been retained when these materials have been used as coatings and have been transferred to the Ti-4 substrates, a comparison between the biological response of the uncoated and coated substrates has been carried out. Moreover, particular attention has been focused on the
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study of the influence of the Ca-content on the material's biological properties. Therefore, all results have been reported as a function of the Ca content. 2. Experimental section 2.1. Sol–gel synthesis and coatings procedure Glasses were synthesized starting from tetraethyl orthosilicate (TEOS, Si(OC2H5)4, Sigma Aldrich) and calcium nitrate tetrahydrate (Ca(NO3)2·4H2O, Sigma Aldrich) as sources of SiO2 and CaO, respectively. To synthesize Ca-free gel, Si(OC2H5)4 was mixed, under stirring, in ethanol 99.8% (Sigma–Aldrich), nitric acid (HNO 3 ≥ 65%, Sigma–Aldrich) and water, according to the following molar ratios: TEOS:EtOH:H2O:HNO3 = 1:5:1:1. The other systems were obtained by adding a solution of calcium nitrate tetrahydrate in ethanol 99.8% (molar ratio Ca(NO3) 2·4H2O:EtOH = 1:22) . Different volumes of the two solutions were mixed to obtain glasses with various molar percentages in calcium, expressed by the following general formula: xCaO·(1 − x)SiO2 with x = 0–0.3–0.4–0.5–0.6 . In Table 1 the label and compositions of the synthesized materials are given. Mixtures were left stirring for 1 h and the resulting sols were uniform and transparent. Subsequently, samples on the dip coater arm (KSV LM, Stockholm, Sweden) were used as substrates for coating. Ti-4 disks were ultrasonically cleaned with acetone and passivated with HNO3 ≥ 65% (Sigma–Aldrich) for 60 min, and finally stored in ethanol 99.8%. Substrates were dried before use in an oven at 50 °C, immersed in the sols and left, protected from dust, for 24 h at room temperature. Titanium grade 4 (Ti-4) disks (Sweden & Martina, Padua, Italy) of 8 mm diameter and 2.20 mm thickness, equipped with a pin for attachment to the gel for 24 h. After various trials, the extraction speed (Ve) of disks from the sols was set at 15 mm/min to obtain a homogeneous layer. The coated substrates were heat-treated at 45 °C for 1 h to promote film densification. After coating, the residual sols were left to gel at room temperature. As a result of hydrolysis of precursor molecules and condensation of the oligomers obtained, after 10–30 days (depending on Ca-content) the gelation of all sols occurred and also the obtained gels were heattreated at 45 °C for 1 h. Glassy and transparent materials were obtained. In Fig. 1 is shown 0.4CaO·SiO2 as an example. The solid samples obtained were subsequently crushed to a powder in a mortar and the resulting powders were used for chemical characterization. 2.2. Chemical and structural characterization of the materials and coatings The pulverized sol–gel glasses were analyzed using Fourier transform infrared spectroscopy (FTIR) to characterize the chemical interactions between the material components. Disks containing 2 mg of sample powder and 198 mg of KBr were made. Transmittance FTIR spectra were recorded in the 400–4000 cm−1 region using a Prestige 21 spectrophotometer (Shimadzu, Tokyo, Japan) equipped with DTGS KBr (Deuterated Tryglycine Sulfate with potassium bromide windows) detector, with a resolution of 2 cm− 1 (45 scans). FTIR spectra were analyzed by Prestige software (IRsolution).
Table 1 Labels and composition of the synthesized gels. Label
Compositions
SiO2 0.3CaO·SiO2 0.4CaO·SiO2 0.5CaO·SiO2 0.6CaO·SiO2
0.0CaO·SiO2 0.3CaO·0.7SiO2 0.4CaO·0.6SiO2 0.5CaO·0.5SiO2 0.6CaO·0.4SiO2
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Fig. 1. 0.4CaO·SiO2 gel after drying.
Fig. 2. FTIR spectra of a) SiO2; b) 0.3CaO·SiO2; c) 0.4CaO·SiO2; d) 0.5CaO·SiO2; e) 0.6CaO·SiO2; f) Ca(NO3)2·4H2O.
The nature of the pulverized SiO2 and of the xCaO·(1 − x)SiO2 glasses was ascertained by X-ray diffraction (XRD) analysis using a Philips diffractometer. Powdered samples were scanned from 2Θ = 5° to 60° using CuKα radiation. Scanning electron microscopy (SEM, Quanta 200, FEI, The Netherlands) was used to investigate the films' morphologies and the influence of calcium. Coated Ti-4 disks were fixed on aluminium stubs with colloidal graphite; as Ti-4 disks are conductive, it was possible to avoid the coating of the samples with a conductive layer.
represent 100% viability. The cells were allowed to grow for 24 h on each sample. Afterwards they were washed 3 times with PBS (phosphate buffered saline) and again incubated with 10% v/v of WST-8 in a fresh medium for 2 h. WST-8 is a colorimetric assay and the number of viable cells is directly proportional to the absorbance value (450 nm). The absorbance value was measured with a UV-visible spectrophotometer (Biomate 3, Thermo Scientific). The statistical analysis of the results was performed using Student's t-test; significance was at the 0.05 level.
2.3. Biological characterization
3. Results and discussion
The assessment of the in vitro bioactivity response was carried out by soaking the un-coated and coated disks for 21 days in Simulated Body Fluid (SBF) as proposed by Kokubo et al. [30]. SBF is an acellular aqueous solution with an inorganic ion composition almost equal to human plasma. It was prepared by dissolving NaCl, NaHCO3, KCl, MgCl2 · 6H2O, CaCl2, Na2HPO4, Na2SO4 (Sigma–Aldrich) in ultra-pure water and buffered at physiological pH 7.40 using 4-(2-hydroxyethyl)piperazine1-ethanesulfonic acid hemisodium salt (HEPES, Sigma–Aldrich) and NaOH. As the formation of the hydroxyapatite layer is influenced by the ratio of the surface sample to SBF volume, the ratio was kept equal to 10 mm2/mL as reported in the literature [30]. The disks and the SBF were placed in a polystyrene bottle. During the experiment the temperature is kept constant at 37 °C ± 0.5 °C by means of a water bath. The solution was exchanged every 2 days to avoid depletion of the ionic species. After 7 and 21 days the samples were removed from SBF, gently rinsed with distilled water and dried in a desiccator. The surface of the films was analyzed with SEM equipped with Energy Dispersive X-ray Spectroscopy (EDXS). The bioactive response of coatings was evaluated in terms of the hydroxyl-apatite layer formation on the surface of the samples. NIH 3 T3 murine fibroblasts cells (ATCC, VA, USA) were used to evaluate the biocompatibility of the films. Cells were seeded onto coated Ti-4 substrates and their viability was tested with the WST-8 Assay (Dojindo Molecular Technologies Inc., MD, USA) [31]. Cells were proliferated in DMEM medium (Gibco, CA, USA) with 10% (v/v) fetal bovine serum, 1% pen-strep, in a humidified incubator, at 37 °C and 5% CO2. The samples were sterilized with 70% ethanol and UV for 6 h and afterwards placed on the bottom of a polystyrene 24-well plate. For each system 3 disks were used and 5000 cells were plated onto each sample. As a negative control, uncoated titanium disks were used and the cells grown on the polystyrene wells have been considered to
3.1. Chemical characterization and coating morphology The infrared spectra of xCaO·(1 − x)SiO2 with x = 0;0.3;0.4;0.5;0.6; (a, b,c,d,e) and Ca(NO3)2·4H2O (curve f) powders are compared in Fig. 2. In all spectra, a broad band at 3440 cm−1 and a peak at 1640 cm−1 are visible which are due to the stretching and bending of water, respectively. The SiO2 spectrum (curve a) is characterized by a strong band at 1080 cm−1 associated with bands at 800 and 460 cm− 1 which are assigned to the symmetric and asymmetric stretching and bending
Fig. 3. XRD spectra of synthesized materials.
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modes of SiO4 tetrahedra, respectively [32]. Moreover, the band at 950 cm− 1 is attributed to surface silanols groups [33] and the sharp peak at 1382 cm−1 is ascribed to N\\O stretching modes [34,35]. The last peak shows that nitrate ions are present in the obtained materials, resulting from HNO3 used for the synthesis. The spectra of the xCaO·(1 − x)SiO2 samples (from curve b to e) show all the described peaks, but those related to SiO4 tetrahedra appear to be slightly shifted to lower wavenumbers (1075 cm−1 and 795 cm−1). That shift can be due to the interaction between Ca2+ ions and the negative charge of SiO4 units. Moreover, in all xCaO·(1 − x)SiO2 sample spectra some peaks typical of calcium nitrate (spectrum f) are visible, such as the strong bands at 1425 and 1357 cm−1 and a sharp peak at 823 cm− 1 due to asymmetric and symmetric stretching vibrations and the bending mode of nitrate ions [36,37]. The nitrate peaks have high intensity in all spectra, whereas the intensity of the main Si\\O\\Si band, at 1075 cm−1, decreases when the Ca2+ content increases. This observation suggests that a part of the calcium nitrate added during the
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synthesis process isn't incorporated into the silica network but coats it. The hypothesis is in agreement with Martin et al. [38]. The authors in their review reported that the entire removal of nitrate ions in calcium silicate materials obtained via sol–gel can only occur after heat-treatment of the gels at temperatures above 350 °C. The diffractograms of all xCaO·(1 − x)SiO2 powders are shown in Fig. 3. They show no Bragg diffraction peaks, indicating the absence of crystalline phases. Hence, the pristine sol–gel prepared glasses are amorphous in nature, having long-range structural disorder. The coated Ti-4 disks were observed by SEM after heat treatment and the thin film micrographs are shown in Fig. 4. In the SiO2 (x = 0) film shown in panel A cracks are evident in the coating. Panel B and C represent 0.3CaO·SiO2 and 0.4CaO·SiO2. The films are homogeneous and without cracks. 0.5CaO·SiO2 and 0.6CaO·SiO2 are shown in panels D and E. Here the surface shows large uncovered areas. This suggests that a high calcium percentage may interfere with the dip-coating process, changing the values of the
Fig. 4. SEM micrographs of A) SiO2; B) 0.3CaO·SiO2; C) 0.4CaO·SiO2; D) 0.5CaO·SiO2 E) 0.6CaO·SiO2 coatings.
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Fig. 5. SEM micrographs of un-coated Ti-4 (panel A) and coated (panel B) samples after 7 days of soaking to SBF. Panel C shows coated samples after 21 days. Panel D: EDXS analysis of globular crystals.
viscosity, surface tension and density. These modifications make sols unsuitable for the formation of thin films. 3.2. Films' biological properties Un-coated (Fig. 5A) and coated (Fig. 5B) disks were observed by SEM after 7 days. Un-coated Ti-4 disks show few globular grains in comparison to all coated samples at same time. The amount of this precipitate increases with the exposure time of the samples to SBF solution. Coated disks, after 21 days in SBF, show the surface completely covered by globular grains (Fig. 5C). As all samples show a similar distribution of globules, only a 0.4CaO·SiO2 sample is shown in Fig. 5 as an example. This globular morphology is typical of the bone-like apatite which precipitates on bioactive materials when they are soaked in SBF solution, as widely reported in the literature [30,39–41]. In order to confirm that the chemical composition of the observed grains was in agreement with the hydroxyl-apatite formula [Ca10(PO4)6(OH)2], EDXS analysis of the precipitate has been carried out (Fig. 5D). An atomic content
ratio Ca/P b 1.67 (in the range 1.56–1.62, percentage report in Table 2) has been recorded for all samples. This result suggests that the precipitate doesn't consist of stoichiometric hydroxyapatite but of Ca-deficient type apatite. The apatite formed in SBF is, indeed, similar to bone apatite because it is generally a Ca-deficient type apatite with lower Ca/P atomic ratio than stoichiometric one (ISO 23317:2014). According to Kokubo et al. [30] materials able to induce the hydroxyapatite nucleation when soaked in SBF, will be able to bond the living bone when implanted in vivo by the formation of a bone-like apatite layer on their surface. Therefore, this in vitro test is widely used in the
Table 2 EDXS percentages of hydroxyapatite after 21 days. Atomic content (%)
SiO2 0.3CaO·SiO2 0.4CaO·SiO2 0.5CaO·SiO2 0.6CaO· SiO2
Ca
P
Ca/P
18.16 19.10 18.65 18.38 19.69
11.21 12.14 11.73 11.78 12.23
1.62 1.57 1.59 1.56 1.61
Fig. 6. WST-8 assay value: cells grown on polystyrene were considered 100% viable, uncoated titanium (Ti-4) show the lowest viability value. * denotes significance.
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biomaterials field in order to predict the osseointegration ability of new materials in vivo. A part of the scientific community doesn't share the use of this test because it can lead to false positive and false negative results [42]. However, the data obtained on the bioactivity of the synthesized materials are in accordance with the results of other researcher groups [43] which proved that calcium silicate-based materials can be osseointegrated when implanted in vivo. An in vitro study [44] demonstrated that these materials can release calcium and silicate ions which induces the proliferation of osteoblasts by means of gene activation. The results of a cell viability test are shown in Fig. 6. The value of cells grown on polystyrene was considered as 100%. The cells grown on uncoated disks (Ti-4) show the lowest viability while the results improved using all coated disks. The data obtained from 0.3CaO·SiO2 are significantly better than that of 0.5CaO·SiO2 and 0.6CaO·SiO2. No statistically significant differences were found among the SiO2 group and 0.5CaO·SiO2 − 0.6CaO·SiO2 groups. This suggests that, in the latter systems, some cells grow in direct contact with the less biocompatible titanium substrate because of the non-homogeneous films (as observed by SEM), while a uniform film allows the best results of cell viability to be obtained. The results are in agreement with data reported in other studies where a decrease of cell viability is recorded on no homogeneous coatings [3,45]. 4. Conclusions Glass materials, composed of silica and calcium in different molar ratios, were synthesized with the sol–gel technique and used to coat Ti-4 disks. The substrates were coated with the dip coating procedure and it was demonstrated that different amounts of calcium affects the film morphology and biocompatibility. FTIR analysis indicates that in the sol–gel materials the two components (Si and Ca) interact, recording a slight shift to lower wavenumbers of SiO4 tetrahedra bonds. SEM analysis shows the film morphology of various systems: SiO2 film is cracked, 0.3CaO·SiO2 and 0.4CaO·SiO2 are homogeneous and crack-free while 0.5CaO·SiO2 and 0.6CaO·SiO2 films show uncovered areas. This inhomogeneity has a negative influence on the biocompatibility. In fact the best results were obtained on compact films. In each case the coated substrates were more biocompatible than uncoated ones. All samples were shown to be bioactive, a fundamental property for their osseointegration after implantation. In conclusion, the results obtained suggest that the 0.3CaO·SiO2 and 0.4CaO·SiO2 systems are most suitable to modify the surface of titanium implants and improve their biocompatibility. Appendix A. Supplementary data Supplementary data to this article can be found online at http://dx. doi.org/10.1016/j.msec.2015.09.033. References [1] M. Catauro, F. Bollino, F. Papale, M. Gallicchio, S. Pacifico, J. Drug Delivery Sci. Technol. 24 (2014) 320–325. [2] M. Vallet-Regi, J. Chem. Soc. Dalton Trans. 97–108 (2001).
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Coatings of titanium substrates with xCaO·(1 − x)SiO2 sol–gel materials: characterization, bioactivity and biocompatibility evaluation M. Catauro ⁎, F. Papale, F. Bollino Department of Industrial and Information Engineering, Second University of Naples, Via Roma 29, 81031 Aversa, Italy
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Article history: Received 9 February 2015 Received in revised form 24 August 2015 Accepted 7 September 2015 Available online 12 September 2015 Keywords: Bioactivity Calcium silicate Glass Sol–gel method
a b s t r a c t The objective of this study has been to develop low temperature sol–gel coatings to modify the surface of commercially pure titanium grade 4 (a material generally used in dental application) and to evaluate their bioactivity and biocompatibility on the substrate. Glasses of composition expressed by the following general formula xCaO·(1 − x)SiO2 (0.0 b x b 0.60) have been prepared by means of the sol–gel route starting from tetraethyl orthosilicate and calcium nitrate tetrahydrate. Those materials, still in the sol phase, have been used to coat titanium substrates by means of the dip-coating technique. Fourier transform infrared spectroscopy (FTIR) and X-ray diffraction (XRD) allowed the materials to be characterized and a microstructural analysis of the coatings obtained was performed using scanning electron microscopy (SEM). The potential applications of the coatings in the biomedical field were evaluated by bioactivity and biocompatibility tests. The coated titanium was immersed in simulated body fluid (SBF) for 21 days and the hydroxyapatite deposition on its surface was subsequently evaluated via SEM–EDXS analysis, as an index of bone-bonding capability. To investigate cell-material interactions, mouse embryonic fibroblast cells (3 T3) were seeded onto the specimens and the cell viability was evaluated by a WST-8 assay. © 2015 Published by Elsevier B.V.
1. Introduction Continuous improvements in medicine and life quality are gradually leading to a senescence of the population in industrialized countries. This trend implies that a considerable portion of the population suffers from diseases caused by a progressive deterioration of the musculoskeletal system. The last decade has been characterized by many studies on biomaterials such as to allow the replacement the damaged tissues, to build long-life prostheses or to release drugs [1–4]. Metallic materials, thanks to their excellent mechanical properties, are used today in orthopedic and dental applications where the implants are subject to severe stresses. However, metallic materials have several significant disadvantages such as the tendency to release ions (toxic) when they come in contact with body fluids and the lack of integration in bone [5–9]. As a consequence, the formation of peri-implant fibrosis may occur, thus inducing prosthesis mobilization and then the reduction of its performance. Ceramic materials show a capability to bond spontaneously and to integrate with bones in living tissues, therefore stimulating osteogenesis [5,10–12]. The clinical use of such biomaterials is limited by their poor mechanical properties ⁎ Corresponding author. E-mail address: [email protected] (M. Catauro).
http://dx.doi.org/10.1016/j.msec.2015.09.033 0928-4931/© 2015 Published by Elsevier B.V.
which hamper monolithic glasses from being used in load-bearing applications, where metallic alloys are still the materials of choice. The coating of metal implants with ceramic materials has enabled corrosion [13] and ion release to be reduced, and bioactivity and biocompatibility [14] to be improved, maintaining the metal's mechanical properties [15]. A common technique used to obtain ceramic or glassy coatings on substrates, even with complex shapes, is the sol–gel dip coating process. Sol–gel is the method used to synthesize ceramic and/or glass materials at a relatively low temperature. Usually the chemical reagents used in the sol preparation are metal organic compounds such as metal alkoxides [M(OR)n], where M represents a network-forming element (such as Al, B, Si, Ti, Zr, etc), and R is an alkyl group. The conversion of the system from sol to gel is the result of hydrolysis and polycondensation reactions of the metal alkoxide precursors. The hydrolysis generates hydroxyl groups which tend to establish new bonds (by means of oxygen bridges) through polycondensation reactions. This process has several advantages when compared to traditional synthesis, such as the capability to obtain homogeneous materials with fine control of the chemical composition; it also allows the densification temperature of the layer to be decreased and it requires less equipment, and is thus cheaper. Materials prepared by the sol–gel process have shown to be more bioactive than those with the same composition but prepared with different methods. In fact, surface analysis of sol–gel materials
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and coatings has shown the presence of hydroxyl groups that can promote the nucleation of calcium and phosphate and, therefore, the osteointegration when those materials are implanted. The sol–gel technique leads easily to various coating techniques such as dip-coating, spin-coating and spray [16–18] . In the literature the use of the sol–gel dip coating technique in dental and orthopedic fields is reported for the surface modifications of metals (e.g. titanium-alloys) in order to reduce corrosion and ion release, and to improve their biological properties [6,19–21]. Bioactive and thin coatings (about 70 to 1000 nm depending on the number of coating layers) can promote an early attachment of the bone to the implant. This increased bond strength and accelerated osteoconductivity are very important features in clinical applications, as the success of an implant is also a function of the initial bone attachment. The dip coating method requires the substrate to be dipped into the ‘sol’ and withdrawn at a constant speed to allow the sol drainage and its instantaneous gelation. Many parameters influence the thickness and the morphology of the obtained coating, such as the withdrawal speed, sol viscosity and drying treatment [22]. The sol–gel dip coating process is an advantageous technique for coating preparation due to the higher purity and homogeneity, lower processing temperatures, reduced thickness, simple and cheap method of preparation and ability to form a physically and chemically uniform coating over complex shapes. The aim of this study is to prepare materials via the sol–gel method with the composition of the adapted glasses expressed by the following general formula: xCaO·(1 − x)SiO2 with x = 0.00; 0.30; 0.40; 0.50; 0.60 film on Titanium grade 4 (Ti-4), to improve their biocompatibility and bioactivity. Ti-4 is a metallic material widely used in implants for orthopedic, dental and orthodontic wires. The coating of these materials allow the main drawback to be overcome which limits their use as implants in many applications [23]. Although titanium and its alloys are able to form spontaneously a stable and protective TiO2 surface layer [24], corrosion can take place when this film is removed due to local mechanical abrasion [23]. Different strategies can be followed to avoid the early failure of the implants due to wear and corrosion and to extend the prosthese lifetime. Many studies describe the use of nitric acid passivation protocols in order to increase the thickness of the oxide films [23,25]. Other studies report that a protective coating can be applied on titanium substrates for this purpose. In particular, the use of sol–gel coatings for corrosion protection which also leads to an improvement of the substrate's bioactivity and biocompatibility was proposed [13,26]. The choice of the material's constituents used to make the coating has been targeted to obtain layers with high bioactivity and biocompatibility. It is known [27] that calcium silicate amorphous glasses, as bulks, are able to form strong bonds to living bones and tissues. They stimulate osteogenesis via both their dissolution and degradation products. Indeed, soluble silica and calcium ions are proposed to both activate and stimulate osteoprogenitor cells at the implant site, promoting bone tissue growth. Different formulations of such materials, which differ in Ca-content, were prepared in order to study the influence of the calcium amount on the biological properties of the coating. In the literature [28,29], indeed, it has been proven that the presence of cations in a material can stimulate hydroxyapatite nucleation when it is soaked in simulated body fluid (SBF). In the present work, the bioactivity and biocompatibility of the prepared coatings have been evaluated by bioactivity tests and WST-8 assays on murine fibroblasts (3 T3) seeded on the layers. In order to assess that the good biological properties detected for calcium silicate have been retained when these materials have been used as coatings and have been transferred to the Ti-4 substrates, a comparison between the biological response of the uncoated and coated substrates has been carried out. Moreover, particular attention has been focused on the
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study of the influence of the Ca-content on the material's biological properties. Therefore, all results have been reported as a function of the Ca content. 2. Experimental section 2.1. Sol–gel synthesis and coatings procedure Glasses were synthesized starting from tetraethyl orthosilicate (TEOS, Si(OC2H5)4, Sigma Aldrich) and calcium nitrate tetrahydrate (Ca(NO3)2·4H2O, Sigma Aldrich) as sources of SiO2 and CaO, respectively. To synthesize Ca-free gel, Si(OC2H5)4 was mixed, under stirring, in ethanol 99.8% (Sigma–Aldrich), nitric acid (HNO 3 ≥ 65%, Sigma–Aldrich) and water, according to the following molar ratios: TEOS:EtOH:H2O:HNO3 = 1:5:1:1. The other systems were obtained by adding a solution of calcium nitrate tetrahydrate in ethanol 99.8% (molar ratio Ca(NO3) 2·4H2O:EtOH = 1:22) . Different volumes of the two solutions were mixed to obtain glasses with various molar percentages in calcium, expressed by the following general formula: xCaO·(1 − x)SiO2 with x = 0–0.3–0.4–0.5–0.6 . In Table 1 the label and compositions of the synthesized materials are given. Mixtures were left stirring for 1 h and the resulting sols were uniform and transparent. Subsequently, samples on the dip coater arm (KSV LM, Stockholm, Sweden) were used as substrates for coating. Ti-4 disks were ultrasonically cleaned with acetone and passivated with HNO3 ≥ 65% (Sigma–Aldrich) for 60 min, and finally stored in ethanol 99.8%. Substrates were dried before use in an oven at 50 °C, immersed in the sols and left, protected from dust, for 24 h at room temperature. Titanium grade 4 (Ti-4) disks (Sweden & Martina, Padua, Italy) of 8 mm diameter and 2.20 mm thickness, equipped with a pin for attachment to the gel for 24 h. After various trials, the extraction speed (Ve) of disks from the sols was set at 15 mm/min to obtain a homogeneous layer. The coated substrates were heat-treated at 45 °C for 1 h to promote film densification. After coating, the residual sols were left to gel at room temperature. As a result of hydrolysis of precursor molecules and condensation of the oligomers obtained, after 10–30 days (depending on Ca-content) the gelation of all sols occurred and also the obtained gels were heattreated at 45 °C for 1 h. Glassy and transparent materials were obtained. In Fig. 1 is shown 0.4CaO·SiO2 as an example. The solid samples obtained were subsequently crushed to a powder in a mortar and the resulting powders were used for chemical characterization. 2.2. Chemical and structural characterization of the materials and coatings The pulverized sol–gel glasses were analyzed using Fourier transform infrared spectroscopy (FTIR) to characterize the chemical interactions between the material components. Disks containing 2 mg of sample powder and 198 mg of KBr were made. Transmittance FTIR spectra were recorded in the 400–4000 cm−1 region using a Prestige 21 spectrophotometer (Shimadzu, Tokyo, Japan) equipped with DTGS KBr (Deuterated Tryglycine Sulfate with potassium bromide windows) detector, with a resolution of 2 cm− 1 (45 scans). FTIR spectra were analyzed by Prestige software (IRsolution).
Table 1 Labels and composition of the synthesized gels. Label
Compositions
SiO2 0.3CaO·SiO2 0.4CaO·SiO2 0.5CaO·SiO2 0.6CaO·SiO2
0.0CaO·SiO2 0.3CaO·0.7SiO2 0.4CaO·0.6SiO2 0.5CaO·0.5SiO2 0.6CaO·0.4SiO2
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Fig. 1. 0.4CaO·SiO2 gel after drying.
Fig. 2. FTIR spectra of a) SiO2; b) 0.3CaO·SiO2; c) 0.4CaO·SiO2; d) 0.5CaO·SiO2; e) 0.6CaO·SiO2; f) Ca(NO3)2·4H2O.
The nature of the pulverized SiO2 and of the xCaO·(1 − x)SiO2 glasses was ascertained by X-ray diffraction (XRD) analysis using a Philips diffractometer. Powdered samples were scanned from 2Θ = 5° to 60° using CuKα radiation. Scanning electron microscopy (SEM, Quanta 200, FEI, The Netherlands) was used to investigate the films' morphologies and the influence of calcium. Coated Ti-4 disks were fixed on aluminium stubs with colloidal graphite; as Ti-4 disks are conductive, it was possible to avoid the coating of the samples with a conductive layer.
represent 100% viability. The cells were allowed to grow for 24 h on each sample. Afterwards they were washed 3 times with PBS (phosphate buffered saline) and again incubated with 10% v/v of WST-8 in a fresh medium for 2 h. WST-8 is a colorimetric assay and the number of viable cells is directly proportional to the absorbance value (450 nm). The absorbance value was measured with a UV-visible spectrophotometer (Biomate 3, Thermo Scientific). The statistical analysis of the results was performed using Student's t-test; significance was at the 0.05 level.
2.3. Biological characterization
3. Results and discussion
The assessment of the in vitro bioactivity response was carried out by soaking the un-coated and coated disks for 21 days in Simulated Body Fluid (SBF) as proposed by Kokubo et al. [30]. SBF is an acellular aqueous solution with an inorganic ion composition almost equal to human plasma. It was prepared by dissolving NaCl, NaHCO3, KCl, MgCl2 · 6H2O, CaCl2, Na2HPO4, Na2SO4 (Sigma–Aldrich) in ultra-pure water and buffered at physiological pH 7.40 using 4-(2-hydroxyethyl)piperazine1-ethanesulfonic acid hemisodium salt (HEPES, Sigma–Aldrich) and NaOH. As the formation of the hydroxyapatite layer is influenced by the ratio of the surface sample to SBF volume, the ratio was kept equal to 10 mm2/mL as reported in the literature [30]. The disks and the SBF were placed in a polystyrene bottle. During the experiment the temperature is kept constant at 37 °C ± 0.5 °C by means of a water bath. The solution was exchanged every 2 days to avoid depletion of the ionic species. After 7 and 21 days the samples were removed from SBF, gently rinsed with distilled water and dried in a desiccator. The surface of the films was analyzed with SEM equipped with Energy Dispersive X-ray Spectroscopy (EDXS). The bioactive response of coatings was evaluated in terms of the hydroxyl-apatite layer formation on the surface of the samples. NIH 3 T3 murine fibroblasts cells (ATCC, VA, USA) were used to evaluate the biocompatibility of the films. Cells were seeded onto coated Ti-4 substrates and their viability was tested with the WST-8 Assay (Dojindo Molecular Technologies Inc., MD, USA) [31]. Cells were proliferated in DMEM medium (Gibco, CA, USA) with 10% (v/v) fetal bovine serum, 1% pen-strep, in a humidified incubator, at 37 °C and 5% CO2. The samples were sterilized with 70% ethanol and UV for 6 h and afterwards placed on the bottom of a polystyrene 24-well plate. For each system 3 disks were used and 5000 cells were plated onto each sample. As a negative control, uncoated titanium disks were used and the cells grown on the polystyrene wells have been considered to
3.1. Chemical characterization and coating morphology The infrared spectra of xCaO·(1 − x)SiO2 with x = 0;0.3;0.4;0.5;0.6; (a, b,c,d,e) and Ca(NO3)2·4H2O (curve f) powders are compared in Fig. 2. In all spectra, a broad band at 3440 cm−1 and a peak at 1640 cm−1 are visible which are due to the stretching and bending of water, respectively. The SiO2 spectrum (curve a) is characterized by a strong band at 1080 cm−1 associated with bands at 800 and 460 cm− 1 which are assigned to the symmetric and asymmetric stretching and bending
Fig. 3. XRD spectra of synthesized materials.
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modes of SiO4 tetrahedra, respectively [32]. Moreover, the band at 950 cm− 1 is attributed to surface silanols groups [33] and the sharp peak at 1382 cm−1 is ascribed to N\\O stretching modes [34,35]. The last peak shows that nitrate ions are present in the obtained materials, resulting from HNO3 used for the synthesis. The spectra of the xCaO·(1 − x)SiO2 samples (from curve b to e) show all the described peaks, but those related to SiO4 tetrahedra appear to be slightly shifted to lower wavenumbers (1075 cm−1 and 795 cm−1). That shift can be due to the interaction between Ca2+ ions and the negative charge of SiO4 units. Moreover, in all xCaO·(1 − x)SiO2 sample spectra some peaks typical of calcium nitrate (spectrum f) are visible, such as the strong bands at 1425 and 1357 cm−1 and a sharp peak at 823 cm− 1 due to asymmetric and symmetric stretching vibrations and the bending mode of nitrate ions [36,37]. The nitrate peaks have high intensity in all spectra, whereas the intensity of the main Si\\O\\Si band, at 1075 cm−1, decreases when the Ca2+ content increases. This observation suggests that a part of the calcium nitrate added during the
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synthesis process isn't incorporated into the silica network but coats it. The hypothesis is in agreement with Martin et al. [38]. The authors in their review reported that the entire removal of nitrate ions in calcium silicate materials obtained via sol–gel can only occur after heat-treatment of the gels at temperatures above 350 °C. The diffractograms of all xCaO·(1 − x)SiO2 powders are shown in Fig. 3. They show no Bragg diffraction peaks, indicating the absence of crystalline phases. Hence, the pristine sol–gel prepared glasses are amorphous in nature, having long-range structural disorder. The coated Ti-4 disks were observed by SEM after heat treatment and the thin film micrographs are shown in Fig. 4. In the SiO2 (x = 0) film shown in panel A cracks are evident in the coating. Panel B and C represent 0.3CaO·SiO2 and 0.4CaO·SiO2. The films are homogeneous and without cracks. 0.5CaO·SiO2 and 0.6CaO·SiO2 are shown in panels D and E. Here the surface shows large uncovered areas. This suggests that a high calcium percentage may interfere with the dip-coating process, changing the values of the
Fig. 4. SEM micrographs of A) SiO2; B) 0.3CaO·SiO2; C) 0.4CaO·SiO2; D) 0.5CaO·SiO2 E) 0.6CaO·SiO2 coatings.
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Fig. 5. SEM micrographs of un-coated Ti-4 (panel A) and coated (panel B) samples after 7 days of soaking to SBF. Panel C shows coated samples after 21 days. Panel D: EDXS analysis of globular crystals.
viscosity, surface tension and density. These modifications make sols unsuitable for the formation of thin films. 3.2. Films' biological properties Un-coated (Fig. 5A) and coated (Fig. 5B) disks were observed by SEM after 7 days. Un-coated Ti-4 disks show few globular grains in comparison to all coated samples at same time. The amount of this precipitate increases with the exposure time of the samples to SBF solution. Coated disks, after 21 days in SBF, show the surface completely covered by globular grains (Fig. 5C). As all samples show a similar distribution of globules, only a 0.4CaO·SiO2 sample is shown in Fig. 5 as an example. This globular morphology is typical of the bone-like apatite which precipitates on bioactive materials when they are soaked in SBF solution, as widely reported in the literature [30,39–41]. In order to confirm that the chemical composition of the observed grains was in agreement with the hydroxyl-apatite formula [Ca10(PO4)6(OH)2], EDXS analysis of the precipitate has been carried out (Fig. 5D). An atomic content
ratio Ca/P b 1.67 (in the range 1.56–1.62, percentage report in Table 2) has been recorded for all samples. This result suggests that the precipitate doesn't consist of stoichiometric hydroxyapatite but of Ca-deficient type apatite. The apatite formed in SBF is, indeed, similar to bone apatite because it is generally a Ca-deficient type apatite with lower Ca/P atomic ratio than stoichiometric one (ISO 23317:2014). According to Kokubo et al. [30] materials able to induce the hydroxyapatite nucleation when soaked in SBF, will be able to bond the living bone when implanted in vivo by the formation of a bone-like apatite layer on their surface. Therefore, this in vitro test is widely used in the
Table 2 EDXS percentages of hydroxyapatite after 21 days. Atomic content (%)
SiO2 0.3CaO·SiO2 0.4CaO·SiO2 0.5CaO·SiO2 0.6CaO· SiO2
Ca
P
Ca/P
18.16 19.10 18.65 18.38 19.69
11.21 12.14 11.73 11.78 12.23
1.62 1.57 1.59 1.56 1.61
Fig. 6. WST-8 assay value: cells grown on polystyrene were considered 100% viable, uncoated titanium (Ti-4) show the lowest viability value. * denotes significance.
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biomaterials field in order to predict the osseointegration ability of new materials in vivo. A part of the scientific community doesn't share the use of this test because it can lead to false positive and false negative results [42]. However, the data obtained on the bioactivity of the synthesized materials are in accordance with the results of other researcher groups [43] which proved that calcium silicate-based materials can be osseointegrated when implanted in vivo. An in vitro study [44] demonstrated that these materials can release calcium and silicate ions which induces the proliferation of osteoblasts by means of gene activation. The results of a cell viability test are shown in Fig. 6. The value of cells grown on polystyrene was considered as 100%. The cells grown on uncoated disks (Ti-4) show the lowest viability while the results improved using all coated disks. The data obtained from 0.3CaO·SiO2 are significantly better than that of 0.5CaO·SiO2 and 0.6CaO·SiO2. No statistically significant differences were found among the SiO2 group and 0.5CaO·SiO2 − 0.6CaO·SiO2 groups. This suggests that, in the latter systems, some cells grow in direct contact with the less biocompatible titanium substrate because of the non-homogeneous films (as observed by SEM), while a uniform film allows the best results of cell viability to be obtained. The results are in agreement with data reported in other studies where a decrease of cell viability is recorded on no homogeneous coatings [3,45]. 4. Conclusions Glass materials, composed of silica and calcium in different molar ratios, were synthesized with the sol–gel technique and used to coat Ti-4 disks. The substrates were coated with the dip coating procedure and it was demonstrated that different amounts of calcium affects the film morphology and biocompatibility. FTIR analysis indicates that in the sol–gel materials the two components (Si and Ca) interact, recording a slight shift to lower wavenumbers of SiO4 tetrahedra bonds. SEM analysis shows the film morphology of various systems: SiO2 film is cracked, 0.3CaO·SiO2 and 0.4CaO·SiO2 are homogeneous and crack-free while 0.5CaO·SiO2 and 0.6CaO·SiO2 films show uncovered areas. This inhomogeneity has a negative influence on the biocompatibility. In fact the best results were obtained on compact films. In each case the coated substrates were more biocompatible than uncoated ones. All samples were shown to be bioactive, a fundamental property for their osseointegration after implantation. In conclusion, the results obtained suggest that the 0.3CaO·SiO2 and 0.4CaO·SiO2 systems are most suitable to modify the surface of titanium implants and improve their biocompatibility. Appendix A. Supplementary data Supplementary data to this article can be found online at http://dx. doi.org/10.1016/j.msec.2015.09.033. References [1] M. Catauro, F. Bollino, F. Papale, M. Gallicchio, S. Pacifico, J. Drug Delivery Sci. Technol. 24 (2014) 320–325. [2] M. Vallet-Regi, J. Chem. Soc. Dalton Trans. 97–108 (2001).
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