Electrochemical and biological characterization of coatings formed on Ti–15Mo alloy by plasma electrolytic oxidation

Materials Science and Engineering C 43 (2014) 172–181

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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec

Electrochemical and biological characterization of coatings formed on Ti–15Mo alloy by plasma electrolytic oxidation Alicja Kazek-Kęsik a, Małgorzata Krok-Borkowicz b, Elżbieta Pamuła b, Wojciech Simka a,⁎ a b

Faculty of Chemistry, Silesian University of Technology, B. Krzywoustego Street 6, 44-100 Gliwice, Poland Faculty of Materials Science and Ceramics, AGH University of Science and Technology, Mickiewicza Street 30, 30-059 Krakow, Poland

a r t i c l e

i n f o

Article history: Received 10 March 2014 Received in revised form 2 June 2014 Accepted 3 July 2014 Available online 10 July 2014 Keywords: Plasma electrolytic oxidation Titanium alloys Bioactive compounds Electrochemical analysis MG-63 osteoblast-like cells

a b s t r a c t β-Type titanium alloys are considered the future materials for bone implants. To improve the bioactivity of Ti–15Mo, the surface was modified using the plasma electrolytic oxidation (PEO) process. Tricalcium phosphate (TCP, Ca3PO4), wollastonite (CaSiO3) and silica (SiO2) were selected as additives in the anodizing bath to enhance the bioactivity of the coatings formed during the PEO process. Electrochemical analysis of the samples was performed in Ringer's solution at 37 °C. The open-circuit potential (EOCP) as a function of time, corrosion potential (ECORR), corrosion current density (jCORR) and polarization resistance (Rp) of the samples were determined. Surface modification improved the corrosion resistance of Ti–15Mo in Ringer's solution. In vitro studies with MG-63 osteoblast-like cells were performed for 1, 3 and 7 days. After 24 h, the cells were well adhered on the entire surfaces, and their number increased with increasing culture time. The coatings formed in basic solution with wollastonite exhibited better biological performance compared with the as-ground sample. © 2014 Elsevier B.V. All rights reserved.

1. Introduction The human adult skeleton is composed of 206 various bone types (with different shapes, sizes and densities). In addition to its mechanical functions, bone tissue participates in the production of blood components, storage of calcium and phosphorous compounds, regulation of blood pH and accumulation and excretion of heavy metals and toxic substances from organisms. Bone is mainly composed of collagen fibers and a primary inorganic bone mineral: hydroxyapatite (Ca10(PO4)6(OH)2, HA) [1,2]. With increasing age, an imbalance between bone resorption and formation leads to a reduction in skeletal bone mass. During accidents, the destruction of bone can occur, and materials for bone replacement become necessary [3]. Materials for bone implants are made from polymers, ceramics, metals or composites. Among biomaterials, metal implants are composed of 316L stainless steel, cobalt–chromium– molybdenum alloys, titanium or titanium alloys (e.g.,Ti–6Al–4V). These biomaterials are widely used in medical applications. However, new materials with better mechanical properties, biocompatibility and surface bioactivity are desirable. Titanium alloys, especially β-phase alloys, exhibit lower Young's modulus (Ti–13Nb–13Zr, Ti–15Mo—79–84 GPa) close to the value of natural bone (10–40 GPa) [4,5] compared with other materials such as tantalum (188 GPa), iron-based alloys (200–205 GPa), cobalt-based alloys (220–230 GPa), 316L stainless steel (193 GPa) and

⁎ Corresponding author. E-mail address: [email protected] (W. Simka).

http://dx.doi.org/10.1016/j.msec.2014.07.021 0928-4931/© 2014 Elsevier B.V. All rights reserved.

titanium (100–115 GPa) [6]. Materials used as implants should be multi-functional. Implants require superior mechanical properties (e.g., elasticity, yield stress, ductility, toughness, hardness, wear resistance), compatibility (tissue reaction, changes in mechanical, physical and chemical properties, degradation), and ease of manufacturing (fabrication methods, cost, quality of raw materials) for orthopedic applications [7,8]. To enhance the biocompatibility of metallic implants, their surfaces can be modified. Various methods are available for the bioactivation of metallic materials: sol–gel methods, plasma spraying, chemical or physical vapor deposition, electrophoretic deposition or sand-blasting. One of these methods, plasma electrolytic oxidation (PEO), is a simple and cost-effective method for surface modification. This process causes a porous oxide layer to be formed on the substrate [9–13]. The coating can be composed of bioactive compounds and can increase the corrosion resistance of the modified materials in Ringer's solution. Depending on the applied voltage and bath composition, oxide layers with different morphologies, oxide layer thicknesses, roughness, wettability and chemical compositions are formed on the substrate. Through the porous layer cells spread on the substrate. Titanium and its alloys are usually tested in vitro with the use of stromal steam cells, human osteosarcoma cells (MG-63, SAOS-2) or mouse osteosarcoma cells (MC3T3-E1). Many investigations have been conducted with different cell types to evaluate the influence of modified titanium alloy surfaces on cell adhesion and proliferation [14–19]. However, PEO-layers formed on titanium alloys in a solution containing bioactive additives are still not sufficiently characterized.

A. Kazek-Kęsik et al. / Materials Science and Engineering C 43 (2014) 172–181 Table 1 The labels and treatment conditions of the samples. Sample no.

TM-200-TCP50 TM-250-TCP50 TM-300-TCP50 TM-350-TCP50 TM-300-W50 TM-300-W100 TM-300-W150 TM-300-S50 TM-300-S100 TM-300-S150

Bath composition

Voltage

Ca(H2PO2)2 M

Ca3PO4 g dm−3

CaSiO3 g dm−3

SiO2 g dm−3

0.1

50

– – – – 50 100 150 – – –

– – – – – – – 50 100 150

– – – – – –

200 250 300 350 300

In this paper, electrochemical and biological characterization of different PEO-layers formed on the surface of Ti–15Mo is presented. Coatings were formed in basic solution (0.1 M Ca(H2PO2)2) with tricalcium phosphate (Ca3(PO4)2), wollastonite (CaSiO3) or silica (SiO2) particles. Electrochemical analysis was performed in Ringer's solution at 37 °C. The cytocompatibility of the layers formed on the substrate was investigated using the osteoblast-like MG-63 cells after 1, 3 and 7 days of cell culture. The number of cells and their morphology on the coatings were monitored. 2. Materials and methods 2.1. Surface modifications The composition of the Ti–15Mo alloy (BIMO Metals, Wrocław, Poland) used in this investigation was as follows: 14.73–14.98 wt.% Mo, 0.016 wt.% N, 0.06 wt.% Fe, 0.08 wt.% C, 0.15 wt.% O, 0.01 wt.% H, and Ti balance. The sample was round and the surface area of the specimens was equal 0.8 cm2. The volume of an anodizing bath was 300 cm2, and the difference between the substrate and counter electrode during the PEO process was 5 cm. Before anodization, the surface of the samples was ground on abrasive paper (SiC) with 600 and 1000

173

granulation. The samples were etched in 1 M HF with 4 M H2SO4 solution and rinsed ultrasonically in deionized water for 5 min. The Ti–15Mo alloy samples were anodized in basic solution (0.1 M Ca(H2PO2)2, Alfa Aesar, Germany) with various compounds used as additives (Ca3(PO4)2, POCH, Poland; CaSiO3, Carl Jäger, Germany; SiO2, STANLAB, Poland). The process was performed in a cooled electrolyzer at an initial current density of 100 mA cm− 2 for 5 min using a DC power supply (PWR800H Kikusui, Japan). The voltage limits were 200, 250, 300, and 350 V. Titanium mesh was used as the cathode, and the treated sample was used as the anode [20]. The labels and treatment conditions of the samples are listed in Table 1. 2.2. Electrochemical analysis The corrosion resistance of the Ti–15Mo alloy samples was investigated using Ringer's solution, which is composed of 8.6 g dm−3 NaCl, 0.3 g dm− 3 KCl, and 0.48 g dm− 3 CaCl2·6H2O (Baxter, USA). The apparatus included a standard two-chamber electrolysis cell with three electrodes: a working electrode, a platinum auxiliary electrode and a Haber-Luggin capillary with a reference electrode (saturated calomel electrode—SCE). The electrolysis cell was powered by a potentiostat (PARSTAT 4000, AMETEK) equipped with Versa Studio software. The investigations included the following measurements: (a) recording the open-circuit potential (EOCP) as a function of time and (b) determining the log j = f(E) curve over a potential range of EOCP − 20 mV to EOCP + 20 mV (dE/dt = 1 mV s−1), which provides information regarding (1) the corrosion potential ECORR, in mV, (2) the corrosion current density jCORR, in A cm−2, (3) the polarization resistance Rp, in Ω cm2 and (4) the cyclic polarization curve (CV) over a potential range from EOCP—0.1 V to 3 V (dE/dt = 10 mV s−1). The corrosion characteristics, such as corrosion potential (ECORR), polarization resistance (Rp) and corrosion current density (iCORR), were extracted from the curves, calculated via the Stearn–Geary method:

jCORR ¼

βa βb B ¼ ; 2:303ðβa þ βb ÞRp Rp

Fig. 1. SEM images of samples anodized at 300 V in basic solution with 50 g dm−3 of a) tricalcium phosphate, b) wollastonite or c) silica.

ð1Þ

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Table 2 The results of the corrosion resistance investigation of the Ti–15Mo alloy in Ringer's solution. Sample no.

EOCP mV

ECORR mV

Rp Ω cm−2

Ti-15Mo TM-200-TCP50 TM-250-TCP50 TM-300-TCP50 TM-350-TCP50 TM-300-W50 TM-300-W100 TM-300-W150 TM-300-S50 TM-300-S100 TM-300-S150

−317.4 226.3 252.0 224.0 206.3 221.1 210.0 220.1 124.6 185.0 152.3

−315.6 197.5 221.9 206.3 191.8 209.1 198.3 208.6 109.76 174.7 143.9

1.2 1.5 1.1 3.1 1.2 7.5 8.0 4.7 1.8 1.8 2.5

where B is the Stern–Geary constant (V), Tafel slopes (anodic βa and cathodic βc) were calculated from Tafel extrapolation (V/decade); and Rp of a corroding metal is defined as the slope of a potential versus current density plot at j = 0:

Rp ¼

  ∂ΔE : ∂j j¼0; dE=dt→0

× × × × × × × × × × ×

105 105 106 105 105 104 105 104 105 104 104

jCORR A cm−2 1.9 2.4 3.3 1.1 3.0 2.9 2.7 4.6 1.2 1.9 1.4

× × × × × × × × × × ×

10−7 10−8 10−8 10−7 10−7 10−7 10−7 10−7 10−7 10−6 10−6

Contact angle ° 47.6 ± 5.71 Strongly hydrophilic 43.7 ± 16.3 32.7 ± 11.5 41.4 ± 10.3 55.6 ± 10.4 36.9 ± 7.2 51.6 ± 9.3 55.0 ± 8.9 33.6 ± 8.6 51.6 ± 25.9

2.3. Wettability and biological investigations The contact angles were determined using a drop shape analysis system (DSA 10Mk2, KRÜSS). The drop image was recorded using a video camera, and an image analysis system (DSA software) calculated

ð2Þ

Fig. 2. Variation of OCP with time for ground samples and samples anodized in basic solution containing a) tricalcium phosphate, b) wollastonite, and c) silica powders.

Fig. 3. The polarization curves for the ground samples and samples anodized in basic solution containing a) tricalcium phosphate, b) wollastonite, and c) silica powders.

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Fig. 4. The cyclic voltammograms recorded for the ground Ti–15Mo alloy (left axis) and the samples anodized in basic solution (right axis) containing a) tricalcium phosphate, b) wollastonite, and c) silica powders.

the contact angle from the shape of the drop. Ultra-high-purity water (0.20 μL) at room temperature was used for the drop shape tests. Ten drops were analyzed for each specimen. The cytocompatibility of the modified titanium alloy samples was determined using a cell viability assay (MTT, 5 g dm−3, (3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyltetrazolium bromide)) (Sigma Aldrich, Germany). MG-63 osteoblast-like cells (European Collection of Cell Cultures, Salisbury, UK) were cultured on the samples for 1, 3 and 7 days. As a control, tissue culture polystyrene (TCPS, Nunclon) was used. The experimental procedures included the following steps: (i) the samples were sterilized in an autoclave at 121 °C for 1 h and moved into 24-well plates, (ii) the cells were cultured at an initial density of 3.0 × 105 cells per well in 1 mL of EMEM cell culture medium (ATTC, USA) supplemented with 10% FBS, 1% penicillin/streptomycin and 0.1% amino acids and sodium pyruvate (PAA, Germany) at 37 °C under a humidified atmosphere with 5.0% CO2, (iii) at the end of each culture time, 1 cm3 of MTT was added into the culture medium and incubated with the samples for 4 h, and (iv) the resulting intracellular formazan was solubilized in dimethylsulfoxide (DMSO, Sigma Aldrich, Germany) and quantified spectrophotometrically at λ = 540 nm using a Multiscan FC microplate photometer (Thermo Scientific, USA). The results of the number of metabolically active cells were calculated

based on calibration curve and were expressed as mean and standard error of the mean (S.E.M.) from three independent samples. Statistical analysis was performed using Student's t-test. Significant differences were assumed at *p b 0.05, **p b 0.01 and ***p b 0.001. The morphology of the adhered cells was examined under a fluorescence microscope (Zeiss Axiovert 40, Carl Zeiss, Germany). The MG-63 cells were fixed in cold methanol for 30 min and then washed with phosphate buffered saline (PBS, PAA, Austria). Then, 0.5 mL of acridine orange (1 g dm−3) was added for 7 min and the samples were washed once again in PBS prior to microscopic observations. 3. Results and discussion Investigations on the modification of the Ti–15Mo alloy in various suspensions using the plasma electrolytic oxidation (PEO) process were earlier presented [20]. All the anodized samples were porous with incorporated bioactive compounds. Fig. 1. presents representative SEM images of the samples anodized in basic solution with tricalcium phosphate, wollastonite or silica at 300 V. The concentration of calcium, phosphorous or silica compounds was dependent on the applied voltage and anodizing bath. The Ca/P atomic ratios for the samples anodized in TCP50 solution at 200, 250, 300, and 350 V were 0.036,

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Fig. 5. Number of cells after 24 h and 3 and 7 days of culturing on the reference TCPS and anodized Ti–15Mo alloy surfaces. The data are expressed as the mean ± S.E.M. of three similar experiments performed in triplicate. The asterisks indicate a statistical significance from the control TCPS group (*) and from the non-modified surface Ti–15Mo group (#): *, #p b 0.05, **, ##p b 0.01, ***, ###p b 0.001.

0.038, 0.083, 0.201 and 0.346, respectively. For the samples anodized in basic solution with 50, 100 and 150 g dm−3 of wollastonite at 300 V, the Ca/P values were 0.150, 0.133, and 0.241, respectively, and for the samples anodized in a solution with 50, 100, and 150 g dm−3 of silica, the Ca/P values were 0.208, 0.056, and 0.042, respectively. The highest Ra parameter was determined for sample TM-350-TCP50 (1.64 μm). No significant difference was observed among the samples anodized at 300 V in solutions composed of different concentrations of the same additive (wollastonite or silica). Among the samples anodized at 300 V, the highest Ra parameter was determined for the sample anodized in basic solution with 150 g dm−3 of wollastonite (Ra = 1.50 μm). All the samples were examined using thin layer X-ray diffraction (TL-XRD) analysis. All the coatings formed in basic solution with wollastonite or silica only exhibited signals for titanium, with the exception of the TM-300-S150 sample, which also exhibited anatase peaks. A small rutile signal was detected for the TM-350-TCP sample; for the other samples anodized in the TCP50 solution, signals from titanium and anatase were detected. No traces of Ca- and P-containing phases were detected by XRD analysis. XPS measurements revealed that in coatings of the samples anodized in basic solution with 50 g dm− 3 of tricalcium phosphate, [PO4]-type ceramic coatings were formed during the PEO process. The valence band structure at 0–14 eV was observed to be similar for all regimes and very close to the calculated total density of the electronic states of hydroxyapatite. The XPS survey spectra of TM-300-W50, TM-300-W100 and TM-300-W150 samples revealed the presence of Ca 2s, Ca 2p, P 2s, P 2p, Si 2s and Si 2p signals. The XPS Ca2p measurements indicated that Ca was mainly present in the form of CaHPO4, with a binding energy of approximately 347.5– 347.6 eV. In coatings anodized in basic solution with silica, Ca, P and Si species were incorporated into the oxide layer formed on the Ti– 15Mo alloy surfaces. The typical O 1s spectra at 532.8 eV for SiO2 was detected for all the samples anodized in basic solution with various concentrations of silica [20]. 3.1. Electrochemical analysis The results of the electrochemical analysis of the Ti–15Mo alloy samples are presented in Table 2. The open circuit potential (EOCP) for all the samples was measured for 60 min in Ringer's solution at 37 °C and the results are shown in Fig. 2. The EOCP values of the modified samples were much higher than that of the as-ground sample. The corrosion potential of the anodized samples increased

from − 317.4 mV to 124.6–252.0 mV. The results indicated that the anodized titanium alloy samples were more chemically stable in Ringer's solution at 37 °C, due to oxide layer which was formed on the substrate during the PEO process. The highest EOCP was recorded for TM-250-TCP50, and the lowest value was recorded for TM-300-S50. The EOCP of the samples anodized in basic solution with 50 g dm−3 of TCP and for TM-300-W100, TM-300-S50 and TM-300-S100 decreased slightly before stabilizing during the first 600 s, which may indicate that activity occurred on the samples in Ringer's solution. The measured EOCP for the samples anodized in basic solution with silica was lower compared to other samples. The coatings formed on the TM-300-S100 and TM-300-S150 samples were thinner than samples anodized at the same voltage in basic solution with tricalcium phosphate or wollastonite [20]. The potentiodynamic polarization curves obtained for the all samples are presented in Fig. 3. The polarization resistance (Rp) and corrosion current density (iCORR) were calculated according to the Stearn–Geary method. For TM-250-TCP50, the sample polarization resistance increased from 1.2 × 105 Ω cm2 to 1.1 × 106 Ω cm2, and the current density decreased from 1.9 × 10−7 to 3.3 × 10−8 A cm−2 and to 2.4 × 10−8 A cm−2 for TM-200-TCP50. The results implies that the surface modification improved the corrosion resistance of the substrate in Ringer's solution. For the other samples, the polarization resistance and current density were similar to the values recorded for the as-ground sample. However, the surface areas of the anodized samples were much higher due to presence of pores, than that of the non-modified titanium alloy surface, which indicated that the current density of these samples could be lower compared with the as-ground sample. It should be taken into account that the compact, welladherent barrier-type oxide layers, was formed on all the anodized samples. This coating protects the substrate against aggressive chemical environment [21–23]. For the samples anodized in basic solution with silica, TM-300-S50 exhibited the highest corrosion resistance in Ringer's solution. The oxide layer formed on this sample was thicker and more compact compared with the other samples, which could explain the better chemical stability during electrochemical analysis. The voltammetric curves for the ground and modified samples of the Ti–15Mo alloy are presented in Fig. 4. In the curve obtained for the ground sample, an increase in the anodic current density was observed above 0.9 V. An intensive rise in the current density begins at 1.4 V, and a maximum was observed at 2.9 V. The observed shape of curve for non-modified Ti–15Mo alloy sample was typical for titanium and other titanium-based alloys [24–26]. The voltammetric profiles

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Fig. 6. Fluorescence microscope images of MG-63 cells on 24 h (a–d) and 3 days (e–h) of culture on samples anodized in basic solution containing 50 g dm−3 of TCP at 200 V (a, e), 250 V (b, f), 300 V (c, g) and 350 V (d, h), bar = 50 μm.

were characteristic of the formation and grown of a passivating film on the substrate with no occurrence of transpassivation in the measured potential range. In the region, where a current plateau characterizes the growth of oxide layer (passivation current). The shape of curve was also in agreement with the behavior found for pure titanium, with the formation of titanium oxides at a lower oxidation state (TiOOH, Ti2O3). At higher potential (0.1 V) Ti3+ was oxidized to Ti4+ and the most stable titanium oxides were formed [27–31]. The anodic peak, near a potential of 1.5 V vs. SCE, implies that oxygen evolution reaction took place [32]. Oliviera and Guastaldi [33] reported that the molybdenum present in titanium alloys improves the resistivity of the anodic oxide film on the substrate. The shapes of the voltammetric

curves for the anodized samples were similar. The recorded current density values were considerably lower due to the presence of a thicker oxide layer, which was a consequence of the anodic oxidation. The lowest current density was recorded for the samples anodized in basic solution with wollastonite. Among the samples anodized in basic solution with β-TCP, lower current densities were recorded for the TM-200-TCP50 and TM-350-TCP50 samples compared with the TM-250-TCP50 and TM-300-TCP50 samples. For the samples anodized in basic solution with wollastonite, the current density recorded for the TM-300-W100 sample was slightly higher than that of the other samples. However, the smoothest curves were obtained for TM-300W50 and TM-300-W150. Similar smooth curves were observed for

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Fig. 7. Fluorescence microscope images of MG-63 cells on 24 h (a–c) and 3 days (d–f) of culture on samples anodized in basic solution containing (a, d) 50 g dm−3, (b, e) 100 g dm−3, and (c, f) 150 g dm−3 of wollastonite, bar = 50 μm.

the samples anodized in basic solution with 50 and 100 g dm−3 of silica; however, no significant difference in the current density was observed for these samples. For all the anodized samples, breakdown of the oxide layer was not observed, indicating that pitting corrosion did not commence. Electrochemical analysis of titanium alloys with modified surfaces is often performed in physiological solution at 36–37 °C. The results of the investigation of the corrosion resistance of the Ti–15Mo alloy in Ringer's solution were similar to the results reported for other β-phase titanium alloys [34–37]. A significant difference in the corrosion behavior between the anodized samples and the substrate was observed. The oxide layers formed on the modified titanium alloy samples offered protection against the harmful effects of aggressive environments. Among the anodized samples of the Ti–15Mo alloy, the highest open circuit potential, lowest current density and highest polarization resistance were recorded for the TM-250-TCP50 sample. On this coating, the osteoblast-like cells were well adhered after each incubation period, and the cells grew and covered the anodized surface. However, better biological properties were observed for the sample anodized in basic solution containing 50 g dm− 3 of wollastonite (TM-300-W50 sample). The wettability of this sample was similar to that of TM-300-W150, TM-300-S50 and TM-300-S150. However, better cell proliferation was observed for this sample after 7 days of cell incubation. The TM-300-W50 sample also exhibited better corrosion resistance in physiological solution compared with the non-modified Ti–15Mo alloy surface.

3.2. Surface wettability and osteoblast-like cell response on the surface of modified Ti–15Mo alloys In Table 2, the measured contact angle of the Ti–15Mo alloy and PEO-coatings formed on the substrate are presented. It is known that surface wettability strongly affects osteoblast cell attachment and differentiation on materials [38–41] of culturing osteoblasts on various materials confirmed that the cells exhibit better adhesion and growth on hydrophilic surfaces [42]. In our case, all the coatings formed on the Ti–15Mo alloy surface were hydrophilic. The measured contact angle was between 32.7 and 55.0°, whereas for the non-modified titanium alloy surface, the contact angle was 47.6°. Based on the microstructure, chemical composition, wettability and surface roughness of the coatings formed on the Ti–15Mo alloy surface during the PEO process, biological investigations were planned. The adhesion and number of MG-63 osteoblast-like cells on the samples after 24 h and 3 and 7 days of culture were investigated. The number of metabolically active cells after each incubation time for all the samples is shown in Fig. 5. After 24 h of cell incubation, the highest number of cells was observed on the surface of the TM-300-W50 sample, whereas, the lowest number of cells was observed on the surface of the TM-300-S50 sample. After 3 days of incubation, the number of cells increased on all the investigated samples. The cells were well-adhered and covered the surface of all the PEO-coatings. Better proliferation of cells on the surface of the TM-300-W100 samples

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Fig. 8. Fluorescence microscope images of MG-63 cells on 24 h (a–c) and 3 days (d–f) of culture on samples anodized in basic solution containing (a, d) 50 g dm−3, (b, e) 100 g dm−3, and (c, f) 150 g dm−3 of silica, bar = 50 μm.

was observed compared with the other samples. No significant difference was observed in the number of cells among the coatings anodized in basic solution with various concentrations of silica after 3 days of cell incubation. After 7 days of incubation, the cells continued to grow and were well adhered on all the samples (the fluorescence microscope images are not presented). However, significant cell proliferation on the surface of the TM-300-W50 sample compared with the other samples was observed. The MTT assay indicated that the lowest cell proliferation occurred on the surface of the TM-300-S50 sample. On the coatings anodized in basic solution with tricalcium phosphate, the number of cells was similar, except for the TM-350-TCP50 sample, where the number of cells was much lower. Figs. 6–9 presents the morphology of the cells after 1 and 3 days of culture on the anodized samples and on non-modified Ti–15Mo alloy surfaces. Cells were well adhered on the surfaces of all the modified samples after 24 h. The growth of cells which was visible on the fluorescence images and the results were in agreement with results obtained from MTT assay. Among samples anodized in basic solution with TCP, higher cell number was observed on the TM-300-TCP150 and TM-350TCP50 samples compared with other samples. However, the highest number of adhered cells was observed on the TM-300-W50 sample (Fig. 7b) after 3 days of culture. No significant differences of cell numbers and their morphology between the samples anodized in basic solution with silica were observed (Fig. 8). On all the coatings cells spread through the coatings after each time of culture without any obvious

dysfunction in their morphology. The cells on the surface of the substrate were elongated (Fig. 9a, c), whereas the cells on all the coatings formed during the PEO process on the substrate were well distributed and spread over the formed coatings. Cell adhesion and proliferation strongly depend on the surface microstructure, chemical composition, wettability and roughness. The coatings formed on the TM-250-TCP50, TM-300-TCP50, TM-350-TCP50 and TM-300-W50 samples exhibited better cell adhesion and proliferation compared with the non-modified Ti–15Mo alloy surface. After 3 days of cell culturing, the samples anodized in basic solution with 50–150 g dm− 3 of silica and the TM-300-W100 samples exhibited better cytocompatibility than the as-ground sample. After 7 days of cell incubation, the cell number was only significantly higher on the coating formed in basic solution with 50 g dm−3 of wollastonite compared with the non-modified titanium alloy surface. Due to β-type alloy properties (e.g., a low Young's modulus), the in vitro cell activity of the ground and modified surfaces of these alloys has often been investigated. H.-T. Chen et al. [43] modified the Ti–13Cr–3Al–1Fe alloy surface using the PEO process in 0.05 M NaH2PO4 solution at 350 and 450 V. The coatings were composed of the anatase or rutile phase, depending on the applied voltage. Investigations with the MC3T3-E1 murine preosteoblast cell line revealed better cell adhesion and proliferation on both layers formed on the substrate than on the raw material. It was also reported that the layer composed of the rutile phase exhibited better bioactivity properties than the layer composed of the anatase

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Fig. 9. Fluorescence microscope images of MG-63 cells on 24 h (a, b) and 3 days (c, d) of culture on (a, c) non-modified Ti–15Mo samples and (b, d) TCPS, bar = 50 μm.

phase. In our case, the oxide layers formed on the Ti–15Mo alloy surface contained mainly the anatase phase, and our surface modification significantly affected the cell adhesion and proliferation, especially on the layers containing wollastonite. Anodization in suspensions resulted in oxide coatings enriched in tricalcium phosphate, wollastonite or silica particles. These compounds exhibited bioactive properties. The effects of β-TCP on the osteoinductive and osteoconductive properties of primary human osteoblast (HOB) cells were investigated by Lu and Zreiqat [44]. After 24 h and 7 days of cell culturing, the expression levels of osteoblastic genes (Cbfa-1, osteopontin, osteocalcin, bone sialoprotein) in HOBs on β-TCP scaffolds were significantly higher compared to the levels of the cells cultured on TCPS. The investigation revealed increasing of the gene expression of α2 integrin (receptors for collagen type I) and decreasing of α5 integrin (receptors for fibronectin) after 24 h of cell culturing on scaffolds composed of β-TCP. After 7 days of cell culture, the concentrations of these genes were similar. It was hypothesized that β-TCP regulates the α2β1 integrin signaling pathway and affects the osteogenic gene expression of HOB cells on the surface [45]. Thus, one of our surface modifications of the Ti–15Mo alloy was performed in suspensions containing β-TCP. Our investigations confirmed cytocompatibility of porous oxide layers incorporating TCP particles formed on β-type titanium alloys during the PEO process. Silicon is an essential element present in the bone and it influences the growth and development of the human skeletal system [46]. Q. Wang et al. [47] performed investigations of the bioactivity of TiO 2 coatings incorporating silicon from a mixture of an organic solution (glycerophosphate sodium pentahydrate) and sodium silicate (Na2SiO3·9 H2O) formed on a titanium surface using the PEO process. Surface investigations using MC3T3-E1 preosteoblast cells revealed that coatings formed on the substrate with and without silicon promote cell attachment to the surface. After 4 and 7 days of cell culturing, the coatings containing the silicon compounds exhibited improved cell proliferation ability compared with the titanium and TiO2 coating samples. Depending on the cell culturing time, the coating containing silicon compounds exhibited enhanced gene expression (alkaline phosphatase—ALP; Runt-related transcription factor 2—Runx-2; bone sialoprotein—BSP; bone-specific gene—OCN). These results indicated that PEO-layers containing silicon compounds formed on titanium

surfaces could stimulate MC3T3-E1 preosteoblast cell proliferation and develop bone minerals on cell membranes. In our case, the cells cultured on coatings with incorporated silica particles were well adhered; however, the number of cells was not higher than that on the as-ground sample. After 3 days of incubation, the cell behaviors on TM-300-S50, TM-300-S100 and TM-300-150 samples were similar and did not depend on the silica concentration in the oxide layers. After 7 days of cell culturing, the number of cells was higher on the surface of the TM-300-S100 sample than on the other coatings formed in basic solution with silica. Better cell proliferation on TM-300-W50 was observed. In this coating, larger pores were formed compared with the pores in the oxide layers formed in basic solution with silica on the substrate during the PEO process. In addition, the coating formed on the TM-300-W50 sample contained a higher concentration of calcium compounds. However, compared with the coatings formed in basic solution containing β-TCP, the morphology and thickness of the coating formed on the TM-300-W50 sample (d = 3.40–4.60 μm) were similar to the coating formed on the TM-300-TCP50 sample (d = 4.15–4.60 μm). The Ra parameter of these layers was also similar (1.30 μm for TM300-W50 and 1.42 μm for TM-300-TCP50). However, the measured contact angle of the TM-300-W50 sample was higher by approximately 23°, and this coating exhibited appropriate conditions for cell adhesion and proliferation.

4. Conclusions Bioactive coatings were formed on β-phase titanium alloy (Ti–15Mo) surfaces using plasma electrolytic oxidation in solutions containing tricalcium phosphate, wollastonite or silica powders. The results confirmed that the Ti–15Mo alloy surface modification by the PEO process significantly increases its corrosion resistance in the solution simulating the physiological media. The greatest corrosion resistance was observed for the sample anodized in basic solution with TCP at 250 V. The current densities calculated for the other samples were a bit higher, which can be associated with the higher surface area due to their porous microstructure. The coatings formed in basic solution with silica were thinner and the open-circuit potentials were lower compared

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to other. However, on all the anodized samples the barrier-type oxide layer was formed and protects against corrosion in Ringer's solution. The coatings were hydrophilic, and after 24 h of incubation, MG-63 osteoblast-like cells were well adhered on all the PEO-coated surfaces. The cells on the surface of the substrate were elongated, whereas on all the PEO-coatings spread over the formed layers. The number of cells increased with increasing incubation time. The cells proliferation strongly depends on chemical composition of coating formed during the PEO process on the substrate. Addition of wollastonite into anodizing bath is more favorable for the growth of cells than tricalcium phosphate or silica. After 7 days, the highest number of cells was observed on the sample TM-300-W50 sample. However, all the coated titanium alloy samples were cytocompatible. Acknowledgments This work was supported by the Polish Ministry of Science and Education under research project no. IP 2012 0459 72. References [1] P. Tate, Seeley's Principles of Anatomy & Physiology, (Chapter 7) Second ed. McGraw-Hill, United States, 2012. [2] J.Y. Rho, L. Kuhn-Spearing, P. Zioupos, Mechanical properties and the hierarchical structure of bone, J. Med. Eng. Phys. 20 (1998) 92–102. [3] D. Green, D. Walsh, S. Mann, R.O.C. Oreffo, The potential of biomimesis in bone tissue engineering: lessons from the design and synthesis of invertebrate skeletons, Bone 30 (2002) 810–815. [4] K. Miura, N. Yamada, S. Hanada, T.-K. Jung, E. Itoi, The bone tissue compatibility of a new Ti–Nb–Sn alloy with a low Young's modulus, Acta Biomater. 7 (2011) 2320–2326. [5] M. Long, H.J. Rack, Titanium alloys in total joint replacement—a materials science perspective, Biomaterials 19 (1998) 1621–1639. [6] S. Wu, X. Liu, K.W.K. Yeung, H. Guo, P. Li, T. Hu, C.Y. Chung, P.K. Chu, Surface nanoarchitectures and their effects on the mechanical properties and corrosion behavior of Ti-based orthopedic implants, Surf. Coat. Technol. 233 (2013) 13–26. [7] B.M. Holzapfel, J.C. Reichert, J.-T. Schantz, U. Gbureck, L. Rackwitz, U. Nöth, F. Jakob, M. Rudert, J. Groll, D.W. Hutmacher, How smart do biomaterials need to be? A translational science and clinical point of view, Adv. Drug Deliv. Rev. 65 (2013) 581–603. [8] S.H. Teoh, Fatigue of biomaterials: a review, Int. J. Fatigue 22 (2000) 825–837. [9] P.K. Chu, Surface engineering and modification of biomaterials, Thin Solid Films 528 (2013) 93–105. [10] S. Minagar, C.C. Berndt, J. Wang, E. Ivanova, C. Wen, A review of the application of anodization for the fabrication of nanotubes on metal implant surfaces, Acta Biomater. 8 (2012) 2875–2888. [11] P.K. Chu, J.Y. Chen, L.P. Wang, N. Huang, Plasma-surface modification of biomaterials, Mater. Sci. Eng. R 36 (2002) 143–206. [12] T.R. Rautray, R. Narayanan, K.H. Kim, Ion implantation of titanium based biomaterials, Prog. Mater. Sci. 56 (2011) 1137–1177. [13] T. Akatsua, Y. Yamada, Y. Hoshikawa, T. Onoki, Y. Shinoda, F. Wakai, Multifunctional porous titanium oxide coating with apatite forming ability and photocatalytic activity on a titanium substrate formed by plasma electrolytic oxidation, Mater. Sci. Eng. C 33 (2013) 4871–4875. [14] S. Yu, Z. Yu, G. Wang, J. Han, X. Ma, M.S. Dargusch, Biocompatibility and osteoconduction of active porous calcium-phosphate films, on a novel Ti–3Zr–2Sn–3Mo–25Nb biomedical alloy, Colloids Surf. B 85 (2011) 103–115. [15] Z. Zhang, B. Gu, W. Zhang, G. Kan, The enhanced characteristics of osteoblast adhesion to porous Zinc–TiO2 coating prepared by plasma electrolytic oxidation, J. Sun Appl. Surf. Sci. 258 (2012) 6504–6511. [16] H. Cimenoglu, M. Gunyuz, G.T. Kose, M. Baydogan, F. Uğurlu, C. Sener, Micro-arc oxidation of Ti6Al4V and Ti6Al7Nb alloys for biomedical applications, Mater. Charact. 62 (2011) 304–311. [17] W.K. Yeung, G.C. Reilly, A. Matthews, A. Yerokhin, In vitro biological response of plasma electrolytically oxidized and plasma-sprayed hydroxyapatite coatings on Ti–6Al–4V alloy, J. Biomed. Mater. Res. B 101 (2013) 939–949. [18] H. Hu, Y. Qiao, F. Meng, X. Liu, C. Ding, Enhanced apatite-forming ability and cytocompatibility of porous and nanostructured TiO2/CaSiO3 coating on titanium, Colloids Surf. B 101 (2013) 83–90. [19] A. Krząkała, A. Kazek-Kęsik, W. Simka, Application of plasma electrolytic oxidation to bioactive surface formation on titanium and its alloys, RSC Adv. 3 (2013) 19725–19743. [20] A. Kazek-Kęsik, G. Dercz, I. Kalemba, K. Suchanek, A.I. Kukharenko, D.M. Korotin, J. Michalska, A. Krząkała, J. Piotrowski, E.Z. Kurmaev, S.O. Cholakh, W. Simka, Surface characterisation of Ti–15Mo alloy modified by a PEO process in various suspensions, Mater. Sci. Eng. C (2014), http://dx.doi.org/10.1016/j.msec.2014.03.008 (in press).

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