Component effects of bioactive glass on corrosion resistance and in vitro biological properties of apatite-matrix coatings

Abstract

BACKGROUND:

Surface modification of metallic implants is critical for improving the clinical performance of the dental and orthopedic devices. Bioactive glasses exhibit different levels of cellular function and physicochemical behavior; however, there have been few previous studies on the effect of constituents of the bioactive glasses on the in vitro osteogenic activity and corrosion resistance of apatite-based coatings.

OBJECTIVE:

The objective of this work was to investigate the effect of SiO₂, CaO, Na₂O, and P₂O₅ on plasma-sprayed apatite coatings on Ti alloy substrates for tailoring the properties of implants making them suitable for clinical applications.

METHODS:

The corrosion potential and corrosion current of various coatings in simulated body fluid (SBF) were examined. MG63 cell proliferation, differentiation, and mineralization of plasma-sprayed apatite-matrix coatings were evaluated.

RESULTS:

The SiO₂ and CaO-containing HA (HSC) coating had a higher corrosion potential than the other three coatings, while SiO₂-containing HA (HS) coating displayed the highest corrosion current among all coatings. The effect of the oxides on cell functions followed the order SiO₂ > CaO > P₂O₅ > Na₂O in terms of cell attachment, proliferation, differentiation, and mineralization.

CONCLUSIONS:

The flexibility in oxide doping may allow for the tunable biological properties and corrosion-resistant ability of the apatite coatings.

Keywords

Hydroxyapatite bioactive glass component effect osteogenesis corrosion

1. Introduction

Much effort to design a new generation of dental and orthopedic implants focuses on bioactive surface coatings aiming at eliciting a favorable cellular reaction to ensure successful medical use [1–5]. Hydroxyapatite (HA)-coated Ti alloy has been accepted as one of the most promising implant materials for load-bearing applications because of its biocompatibility and mechanical properties [1], along with a protective layer shielding the release of metal ions. Nevertheless, the osteoconductive HA is a low bioresorbable ceramic, which makes it to remain in the body for a long time after implantation. In addition, the larger the solubility rate of the ceramic, the more pronounced is the enhanced effect of bone tissue growth because the local biological microenvironment can be influenced significantly upon its degradation [6]. This impels to develop alternative materials such as bioactive glasses in an effort to expedite the clinical practice, although the mechanical properties of bioactive glass lower than HA [7,8]. Several studies have been made to improve bioactivity and/or biodegradation of HA by incorporating bioactive glass [4,7,9].

Bioactive glasses consisting of a variety of oxide components in defined proportions, typically SiO₂, CaO, Na₂O, and P₂O₅, have attracted much attention. Bioactive glasses can be used in bone repair and regeneration applications with various types such as particle, scaffold, and coating. The bioactive glasses undergo an osseointegration process that involves biodegradation, apatite precipitation and the final bone formation onto the implant surface [6]. Oonishi et al. [10] reported that 45S5 bioactive glass granules promoted more rapid bone proliferation compared to HA in a rabbit femur mode. However, the highly degradable coating causes the instability of the metallic implant in the long-term use [11]. Therefore, the potential use of bioactive glass as an additive to HA has been developed. For example, a Ca-rich bioactive glass topcoat was introduced to increase in-vitro reactivity of HA coatings [4]. In vivo study showed that the bone-bonding ability of HA can be improved by adding silicate-based bioactive glass [12].

A crucial part of bioactive glass is the release of constituent ions upon implantation or exposure to the physiological environment, which could exhibit different levels of physicochemical behavior such as the degradation rate and cellular function [13]. The clinical success of some commercial bioactive glasses can be due to the controlled release of active ions including Ca and Si [9]. Investigations into the ion-related cell functions provide guidance for the design of future biomaterials that would enhance cell attachment, proliferation, differentiation, and bone mineralization. Studies have shown a higher proliferation rate and/or collagen secretion in presence of ionic products from bioactive glasses than calcium phosphates [14]. In a previous study, we found the component-dependent apatite formation rate of plasma-sprayed apatite-based coatings [2]. In this study, it is hypothesized that oxide dopants originating from components of bioactive glass in the apatite coatings can enhance coating osteoconductivity, which will favor implant lifetime in vivo. The objective of the present study was thus to investigate the effect of each individual component of generally used bioactive glass on the corrosion resistance and in vitro biological properties of apatite-based coatings plasma-sprayed on titanium alloy substrates.

2. Materials and methods

2.1. Preparation of powder

In our previous work, we have reported the preparation of oxide-doped HA powders, as well as plasma sprayed coatings [2]. Briefly, the commercial HA powder (Merck A.G., Darmstadt, Germany) with particle size < 1 μm was used as matrix material, while reagent grade SiO₂, CaCO₃, Na₂O and NH₄H₂PO₄ powders were used as dopant phases. Chemical compositions of a series of HA-matrix powder specimens are listed in Table1. These HA-matrix powders were prepared using a sinter-granulation method [7] that involved mixing and ball-milling appropriate amounts of HA and various oxide powders in ethyl alcohol, followed by drying and sintering at 1300 °C for 24 h. The sintered granules were then crushed, milled and sieved to obtain particles with sizes 44–149 μm.

Table 1
Composition (wt%) of HA-matrix powders for plasma spray and corrosion parameters of respective coatings

Sample code Composition (wt%) Corrosion parameter

HA SiO₂ CaO P₂O₅ Na₂O E _corr(mV) I _corr(nA/cm²)

HS 95 5 0 0 0 −675 ± 20^a 698 ± 98^a

HSC 95 2.5 2.5 0 0 −543 ± 72^b 545 ± 91^a,b

HSCP 95 2.5 1.5 1 0 −700 ± 29^a 488 ± 66^b

HSCPN 95 2.5 1.5 0.5 0.5 −702 ± 26^a 445 ± 36^b

Sample code	Composition (wt%)	Corrosion parameter
HS	95	5	0	0	0	−675 ± 20^a	698 ± 98^a
HSC	95	2.5	2.5	0	0	−543 ± 72^b	545 ± 91^a,b
HSCP	95	2.5	1.5	1	0	−700 ± 29^a	488 ± 66^b
HSCPN	95	2.5	1.5	0.5	0.5	−702 ± 26^a	445 ± 36^b

Values are mean ± standard deviation. Mean values followed by the same superscript letter (e.g., a, b, c) were not significantly different (P > 0.05) according to Scheffé’s post hoc multiple comparisons.

2.2. Preparation of coating

Commercially available plates of Ti-6Al-4V alloy were used as the substrate material. Prior to plasma spray, the substrate surface was mechanically polished to #1200 grit level, cleaned and sandblasted with 450 μm SiC particles. The plasma spray coating was performed in air using a Plasma-Technik A-3000 system (Plasma Technik A.G., Wohlen, Switzerland). To obtain a uniform coating, the substrate was mounted on a disk which could rotate during plasma spray. For simplicity, the specimen code “HSCPN” stood for the plasma-sprayed coating derived from the sinter-granulated powder containing HA, SiO₂, CaO, P₂O₅, and Na₂O, while “HS” stood for the coating from the powder comprising HA and SiO₂.

2.3. Corrosion measurement

The corrosion measurements included open circuit potential (OCP) time methods and potentiodynamic polarization in a non-deaerated simulated body fluid (SBF) solution, using a CHI 660A electrochemical system (CH Instrument, Austin, TX, USA). The SBF solution, the ionic composition of which is similar to that of human blood plasma, consisted of 7.9949 g NaCl, 0.3528 g NaHCO₃, 0.2235 g KCl, 0.147 g K₂HPO₄, 0.305 g MgCl₂⋅ 6H₂O, 0.2775 g CaCl₂, and 0.071 g Na₂SO₄ in 1000 mL distilled H₂O and was buffered to pH 7.4 with hydrochloric acid (HCl) and tris-hydroxymethyl aminomethane (Tris, (CH₂OH)₃CNH₂) [15]. All chemicals used were of reagent grade and used as obtained. For OCP measurement, only two electrodes (working electrode and reference electrode) were involved, whereas for the potentiodynamic polarization method, a conventional three-electrode cell was used. A saturated calomel reference electrode (SCE) and a platinum counter electrode were employed. The sample surface was cleaned by distilled water. The evaluation of potentiodynamic polarization was started after immersion in SBF for 1 h. The scanned potential range varied from −1 to 1 V toward the anodic direction at a sweep rate of 1 mV/s in the Tafel mode. The current was recorded in the absence of stirring or gas bubbling into the electrolyte. The corrosion potential (E _corr) and corrosion current density (I _corr) were provided after being analyzed by the software. The results were obtained from five separate experiments.

2.4. Cell culture

The mouse fibroblast cell line L929 (BCRC RM60091, Hsinchu, Taiwan) and human osteoblast-like cells line MG63 (BCRC 60279, Hsinchu, Taiwan) were used to evaluate cytotoxicity and osteogenic activity of the coating implants, respectively. The cells were suspended in Dulbecco’s modified Eagle medium (DMEM; Gibco, Langley, OK) containing 10% fetal bovine serum (FBS) (Gibco) and 1% penicillin/streptomycin solution (Gibco) in 5% CO₂ at 37 °C. Before cell incubation, specimens were sterilized by soaking in a 75% ethanol solution and exposure to UV light overnight.

2.5. Cytotoxicity

L929 cells were incubated on the specimens for 12, 24, and 48 h. Cell suspensions (10⁴ cells per well) were seeded over each of the specimens in a 24-well plate. The Ti alloy without the coating was used as a negative control, while the 10% dimethylsulfoxide (DMSO; Sigma-Aldrich) was used as a positive control. After the established L929-cell incubation period, the cytotoxicity was assayed using the MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide; Sigma-Aldrich) assay, in which tetrazolium salt was reduced to formazan crystals by the mitochondrial dehydrogenase of living cells. Briefly, 3 h before the end of the incubation period, 20 μL of MTT solution and 180 μL of DMEM containing 1% penicillin/streptomycin were added to each well. Upon removal of the MTT solution, 200 μL of isopropanol were also added to each well. The plates were then shaken until the formazan crystals had dissolved, and 150 μL of the solution from each well was transferred to a new 96-well plate. Plates were read using a Sunrise microplate reader (Tecan, Salzburg, Austria) at 570 nm, with a reference wavelength of 650 nm. The absorbance results were recorded for three independent measurements. The cell viability was normalized to the negative control (the pristine Ti alloy plate) in terms of absorbance.

2.6. Cell attachment and proliferation

MG63 cell viability was examined using the MTT assay. To assess attachment, cells were cultured for 6 h, 12 h, and 1 day, while cell proliferation was assessed at days 3 and 7. Cell suspensions were seeded on specimen surfaces in a 24-well plate. Before the end of the incubation period, 100 μL of MTT solution and 900 μL of DMEM containing 1% penicillin/streptomycin were added to each well followed by 3 h of incubation. After incubation, supernatants were removed, and the 500 μL of DMSO (Sigma-Aldrich) were added to each well. The plates were then shaken until the formazan crystals had dissolved, and 150 μL of the solution from each well was transferred to a new 96-well plate. Plates were read in a Sunrise microplate reader at 570 nm, with a reference wavelength of 650 nm. The results were reported in terms of absorbance form three independent measurements.

2.7. Alkaline phosphatase activity

To evaluate the effect of dopants on early cell differentiation, ALP activity of MG63 at a density of 5000 cells per well on the various specimens was measured after 7 and 14 days of incubation. ALP catalysed the hydrolysis of the colourless organic phosphate ester substrate (p-nitrophenyl phosphate; pNPP) to p-nitrophenol (a yellow product) and phosphate. ALP activity was measured using the TRACP & ALP assay kit (Takara, Shiga, Japan) according to the manufacturer’s instructions. Six measurements were carried out.

2.8. Calcium quantification

The mineralized matrix synthesis was analysed using an Alizarin Red S staining method, which identifies calcium deposits. After culturing for 7 and 14 days, MG63 cells were washed with PBS and fixed in 4% paraformaldehyde (Sigma-Aldrich) for 10 min at 4 °C. This was followed by staining for 10 min in 0.5% Alizarin Red S (Sigma-Aldrich) in PBS at room temperature. The stained cells were completely washed with PBS to reduce nonspecific Alizarin Red S stain. To quantify matrix mineralization, the calcium mineral precipitate was destained with 10% cetylpyridinium chloride (Sigma-Aldrich) in PBS for 30 min at room temperature. The absorbance of Alizarin Red S extracts was measured at 560 nm using a Sunrise microplate reader. Mean absorbance values were obtained from six independent experiments.

2.9. Statistical analysis

One-way analysis of variance (ANOVA) statistical analysis was used to evaluate the significance of the differences between the mean values. Scheffé’s multiple comparison testing was used to determine the significance of the deviations in the data for each specimen. In all cases, the results were considered statistically significant at a P value of less than 0.05.

3. Results and discussion

3.1. Corrosion behavior

The outermost surface of the metallic implant is modified to promote the interaction between the implant and its human host for improving the clinical performance of the medical devices [16,17]. The role of bioinorganic ions in the formation, regulation and maintenance of bone is imperative for being unraveled [18]. Understanding the efficacy of glass constituents can be advantageous to enhance the physicochemical and biological properties of the bioactive glasses. First of all, in order to clarify the corrosion behavior of the present coating implants, the OCP over time of the samples in SBF is demonstrated in Fig. 1. The OCP of the all samples shifted towards a steady state after soaking in SBF. Compared to the Ti control, all coating samples showed a higher initial potential and attained more noble potential values, indicating a superior corrosion behavior. Among the coating samples, the SiO₂- and CaO-containing HA (HSC) coating surface exhibited a better corrosion resistance during the OCP measurements than the other coating surfaces by virtue of more noble potential. Interestingly, the HSC coating approached an active direction at the initial test time, which can be due to the dissolution-precipitation that occurred at the coating surface. In the earlier study, we found that not only the immersed HSC coating had the highest rate of apatite precipitation among all coatings, but also the HSC-immersed solution reached its maximal Ca concentration the earliest when soaked in an SBF solution [2]. The faster apatite precipitation might be a protective layer to reduce the occurrence of corrosion, which can explain the fact of a higher OCP for HSC coating than the other coating samples at a later stage of the OCP examination.

Fig. 1.

Open circuit potential (OCP) of Ti alloy substrates with and without coatings in SBF.

Typical potentiodynamic polarization curves are presented in Fig. 2. The measured corrosion potential and corrosion current density of all coating samples were obtained using Tafel extrapolation methods, as summarized in Table1. For the corrosion potential, the E _corr value of the pristine Ti alloy substrate was −859 mV versus SCE, while the E _corr values of the coating samples ranged from −543 to −702 mV. The statistical analysis revealed a significant difference (P < 0.05) between the Ti alloy and the coating samples. Similar to the findings of the OCP examination, the HSC group had a corrosion potential (−543 mV) significantly (P < 0.05) higher than the other coating samples (−699 to −702 mV). The shifting of corrosion potential of the materials towards more positive values demonstrated a greater corrosion-resistant ability [19]. The corrosion rate of a coating implant is related to the corrosion current density (I _corr). The higher the I _corr value, the greater the corrosion rate was on the coating when exposed to SBF. The comparative study elicited that the average I _corr values of the coating samples between 445 and 698 nA/cm² were dependent on the components of the coating samples. The HS coating had a significant (P < 0.05) higher I _corr value (698 nA/cm²) in comparison with the Ti alloy substrate (50 nA/cm²), the HSCP coating (488 nA/cm²), and HSCPN coating (445 nA/cm²), possibly because of the release of Si ions during corrosion examination in SBF. There was not significant difference in I _corr between HSC coating and HC coating.

Fig. 2.

Typical potentiodynamic polarization curves of Ti alloy substrates with and without coatings in SBF.

It has been pointed out that the SiO₂ component in the silicate-based bioactive glass is the glass network former, while CaO and Na₂O components are network modifiers. The ratio between the network former and the network modifier in a glass determines its bioactivity and resorbability [20]. The Na₂O-SiO₂ glass can form the Si-OH groups because the Na⁺ ions exchange with H₃O⁺ ions in SBF to form silanol (Si-OH) groups, thereby producing apatite layer on its surface [21]. Na ions impart degradability upon silicate glass networks providing a certain amount of control over constituent ion release [22]. On the other hand, phosphate-based bioactive glasses are potential bioabsorbable materials and offer a more controlled rate of dissolution because of their solubility [23], as compared to silicate based glasses. This might explain why HSCP and HSCPN coatings had a lower corrosion potential compared to the HSC coating.

3.2. Cytotoxicity

Unraveling the effect by which cells respond to ions will be beneficial for materials design in their eventual clinical use. On that account, the biological roles of the soluble species from the bioactive glasses are necessary to be clearly understood. The present study shed light on the effects by which the ions influenced biological functions, in addition to the corrosion resistance. Before the promising coating implants for in vivo evaluation, the cytotoxicity of the coating samples should be demonstrated. The results of the MTT assay for cytotoxicity are shown in Fig. 3. Lower viability values verified higher cytotoxic potential for the test sample. The cells on the positive control (10% DMSO) showed a high degree of cytotoxicity with respect to the Ti control at all culture time, which L929 cell viability significantly (P < 0.05) decreased with the increasing culture time. In contrast to the findings, when the cells were seeded on different coating surfaces, L929 cell viability more than 100% was found. Regarding cytotoxicity, according to ISO 10993-5 standard definition, a viability of more than 70% is considered non-cytotoxic [24]. Of particular note, the HS coating showed the higher viability at all culture time points between the coating samples, while the HSCPN coating had the lower viability.

Fig. 3.

Cytotoxicity of various test samples seeded with L929 cells at various time-points. 10% DMSO was used as the positive control. The cell viability was normalized to the pristine Ti alloy in terms of absorbance. ^∗Statistically significant difference (P < 0.05) from the HS coating.

3.3. Cell attachment and proliferation

To elucidate the effects of various oxide components on in vitro osteogenic activities, the biological functions of MG63 cells cultured on coating surfaces were evaluated. The numbers of initially attached cells for the control and coatings were different after culture for 6 h, as shown in Fig. 4. It can be clearly seen that the attachment of MG63 cells cultured on HS coating surfaces was significantly (P < 0.05) higher than those on the surfaces of the other coating samples at all culture time points, whereas HSCP and HSCPN coatings had the lower cell attachment. Regarding cell proliferation, the steadily increased absorbance value revealed increasing numbers of viable cells for all of the samples on Days 3 to 7. On day 7, the absorbance value for HS coating was approximately 15% higher than that of the HSCPN coating.

Fig. 4.

MTT assay of MG63 cells cultured on the samples to demonstrate cell attachment and proliferation at various culture time points. ^∗Statistically significant difference (P < 0.05) from the HS coating.

3.4. Cell differentiation

The intracellular ALP level was measured to observe the early differentiation activity of MG63 cells. Not surprisingly, all coating samples had the higher ALP level than the Ti control at all incubation times (Fig. 5). Both HS and HSC coatings had a similar ALP expression, but they were higher than HSCP and HSCPN coatings.

Fig. 5.

ALP assay on MG63 cells presented as absorbance for cell differentiation on various samples after 7 and 14 days of culture. ^∗Statistically significant difference (P < 0.05) from the HS coating.

3.5. Mineralization

Quantification of calcium mineral deposits by the Alizarin Red S assay showed that on days 7 and 14, less mineral deposition was found in MG63 cells cultured on the Ti control, while HS coating had the greatest mineral deposition (Fig. 6). With increasing culture time, mineral deposition increased for the cells cultured on all samples. By day 14, significant 60% increment (P < 0.05) of Ca deposits was measured for the HS coating compared to the Ti control. The effect of the coatings on cell mineralization followed the order HS > HSC > HSCP > HSCPN. In vitro cell culture showed no marked difference in the proliferation and differentiation of MG63 cells between HC and HSC coating.

Fig. 6.

Quantification of calcium mineral deposits by Alizarin Red S assay of MG63 cells cultured on various samples after culture for 7 and 14 days. ^∗Statistically significant difference (P < 0.05) from the HS coating.

Cell viability and functions associated with a biomaterial are closely related to the physical, chemical and biological characteristics of the materials used [25,26]. The introduction of bioactive glass to HA has been shown to tune the dissolution of the final system that could enhance its biological response [27]. The present investigations verified a different role for silicon, calcium, phosphate, and sodium components in controlling osteoblast viability, proliferation and differentiation. Ca is a potent regulator of cell behavior and has significant effects on the proliferation and differentiation of osteoblasts. In tissue culture models, elevation in Ca concentration increases osteoblast chemotaxis and proliferation and alters the levels of expression of some differentiation markers [28]. Si ions can enhance attachment, proliferation, differentiation and mineralization of osteoblast-like cells and mesenchymal stem cells [29], as well as osteogenesis in vivo [30]. Tang et al. [31] reported that 0.8 wt% Si content in the plasma-sprayed SiHA coatings enhanced cell growth with compared to the pure HA coating. The osteoblasts in the presence of ionic products from 60% SiO₂-containing bioactive glass dissolution showed high proliferation and collagen secretion when compared to biphasic calcium phosphate, which may be related to silicon contents of bioactive glass [32]. Si has been shown to play an integral role in in vivo bone formation. So et al. [12] found 1 wt% SiO₄-based glass-containing HA can bond to bone more rapidly and firmly than the HA control in a rabbit model. It is believed that a combination of some of the released ions from the dissolution triggers cells to produce new bone [33], in particular the Ca and Si ions are thought to be critical role, consistent with the data of the present study.

Inorganic Phosphate (P) was reported to stimulate expression of matrix Gla protein, a key regulator for bone formation [34]. Low levels of phosphate have been shown to stimulate osteoclastic resorption, as well as osteoblastic differentiation [35]. In contrast, Meleti and colleagues [36] found the cell death occurred through apoptosis when primary human osteoblast-like cells were treated with 7 mM inorganic phosphate. On the other hand, the Na ion in bioactive glass compositions fails to enhance biological effects itself, but to impart degradability and control over the release of other constituent ions [37]. Furthermore, the glasses of higher Na₂O content were associated with a cytotoxic response [21]. The effect of reactivity of individual components in the bioactive glass on osteoblastic function might be to demonstrate their unique role in bone bonding processes. Thus, these might explain why the Si and Ca components in the coatings would induce more superior biological functions than P and Na components.

4. Conclusions

In this study, the constituents of the bioactive glass have a remarkable effect on its ability to support the cell functions in vitro and reflect the corrosion resistance. The Si and Ca components indicated the higher promotion of cell growth and differentiation compared to P and Na components, in addition to greater corrosion-resistant ability. The flexibility in oxide doping nay allow for the tunable biological properties and corrosion-resistant ability of the apatite-matrix coatings. Further investigations such as an in vivo study are needed to elucidate the constituent effects.

Footnotes

Acknowledgements

The authors acknowledge the financial support of Antai Medical Care Cooperation Antai Tian-Sheng Memorial Hospital and Chung Shan Medical University under contract no. CSMU-TSMH-107-01.

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