ZrCuFeAlAg thin film metallic glass for potential dental applications

Intermetallics 86 (2017) 80e87

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ZrCuFeAlAg thin film metallic glass for potential dental applications Chu-Ning Cai a, Cheng Zhang a, *, Ying-Sui Sun b, Her-Hsiung Huang a, b, Cong Yang a, Lin Liu a a

The State Key Lab of Materials Processing and Die & Mould Technology, School of Materials Science and Engineering, Huazhong University of Science and Technology, 430074, Wuhan, China b Department of Dentistry, National Yang-Ming University, Taipei, Taiwan

a r t i c l e i n f o

a b s t r a c t

Article history: Received 9 December 2016 Received in revised form 27 February 2017 Accepted 18 March 2017

In this study, the application of a Ni-free Zr60.14Cu22.31Fe4.85Al9.7Ag3 thin film metallic glass (TFMG) was examined as an approach to retard the poor tribological properties of Ti-alloys for dental applications. The TFMGs were coated on biomedical Ti6Al4V substrate by single-target magnetron sputtering. The fretting resistance was assessed using a reciprocating tribo-tester against Si3N4 counterpart in air and in artificial saliva. Bio-corrosion resistance of TFMG-coated Ti6Al4V samples was tested via electrochemical polarizations and electrochemical impedance spectroscopy in artificial saliva. Biocompatibility of the TFMG was tested in vitro, in comparison with that of the Ti6Al4V alloy. The results showed that this TFMG not only possessed 1.8e2.8 times higher fretting resistance than the Ti6Al4V alloy under various tribological conditions, but also lower bio-corrosion rate and superior passive film. In-vitro assessments of cytotoxicity and cell adhesion indicated that the TFMG has no any cytotoxicity and well-flattened cell adhesion morphology, as good as the Ti6Al4V alloy. The present Ni-free Zr-based TFMG is expected to improve the lifetime and quality of biomedical implants or devices for dental applications. © 2017 Elsevier Ltd. All rights reserved.

Keywords: Zr-based thin film metallic glass Fretting Biocorrosion Biocompatibility Dental applications

1. Introduction The appropriate selection of a biomedical material is usually a complicated process because the biocompatibility must be assured together with sufficient mechanical durability. Ti6Al4V alloy is widely used as dental implant and orthodontic bracket owing to its high specific strength, good toughness, high biocorrosion resistance and good biocompatibility [1]. However, the Ti6Al4V alloy suffers from poor wear resistance, which significantly reduce the service life of implants made of Ti6Al4V [2]. Another incidental weakness related to its poor wear resistance is, for example, fretting wear may generate fine debris which release non-compatible metal ions (e.g. Vnþ and Al3þ) into the body, causing the wellknown “metal allergy” and granuloma [3]. To solve these drawbacks, surface modifications of Ti-alloys by coating wear-resistant films have been demonstrated to be an effective route [4]. Thin films including TiN, CrN, TiO2, Nb2O5 and diamond-like carbon are explored, which were demonstrated to successfully extend the tribological performance of Ti-alloys several times [5e7]. However,

* Corresponding author. E-mail address: [email protected] (C. Zhang). http://dx.doi.org/10.1016/j.intermet.2017.03.016 0966-9795/© 2017 Elsevier Ltd. All rights reserved.

the use of ceramic thin films in implants introduces new challenges, i.e., high internal stress, poor adhesive properties or high sensitivity to ambient conditions, which lead to spallation or interfacial separation under repeated loading [4,8,9]. Thus, thin films of better mechanics-compatibility with Ti-alloys have to be employed for dental applications. In comparison with conventional crystalline films, Zr-based thin film metallic glasses (TFMGs) have been gaining great attentions for bio-medical applications in recent years [10,11]. The TFMGs with atomically disordering structure had significantly increased the hardness, fatigue and anti-corrosion properties of Ti-alloys. For example, a 200 nm-thick Zr51Cu31Al13Ni5 TFMGs used by Lee et al. [12] was found to exhibit ~17% improvement in fatigue limit and 17-times enhancement in the fatigue life with respect to Ti6Al4V. Similarly, the Zr-Ti-Si TFMG developed by Ke et al. [13] had not only reduced modulus (i.e., ~110 GPa for TFMG and ~135 GPa for pure Ti) but also improved corrosion resistance in the simulated body fluid, compared to pure Ti. More importantly, unlike brittle ceramic coatings, nanometer-thick Zr-based TFMGs exhibited remarkable ductility at room temperature [14e16]. On the other hand, some Cu- and/or Ag-bearing Zr-based TFMGs were shown to exhibit nearly 100% antimicrobial rate against Escherichia coli and Staphylococcus aureus, which makes them promising candidates for bio-

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medical applications [17e19]. Nevertheless, studies on the tribological, bio-corrosion properties and biocompatibility under simulated physiological environment, for Ni-free Zr-based TFMGs, are still limited [20]. In this work, a newly-developed Zr60.14Cu22.31Fe4.85Al9.7Ag3 BMG system without toxic elements (i.e. Ni and Be) was deposited on Ti6Al4V substrate using single-target magnetic sputtering. This BMG system was chosen because of its large glass-formation ability (it can be casted into a 10 mm rod in diameter), excellent mechanical properties and good biocompatibility [21]. The most important requirements for dental applications such as adhesion, fretting resistance, biocorrosion and biocompatibility of the TFMG coated Ti6Al4V alloy were systematically investigated and compared to the commercial Ti6Al4V alloy.

electrochemical measurements in the artificial saliva, with a typical three electrode cell consisting of the corrosion sample as a working electrode, a platinum-foil counter electrode and a saturated calomel reference electrode (SCE). Potentiodynamic polarization behavior was characterized by recording a Tafel plot at a potential sweep rate of 0.5 mV s1 from 0.3 VSCE to 1.5 VSCE after immersing the specimens open to air for an hour, when the open-circuit potential (OCP) became almost steady. Each electrochemical test was repeated at least three times for repeatability. The stability of the passive film formed on the TFMG was studied by potentiostatic tests under a constant applied potential of 0 VSCE for 30 min. In addition, electrochemical impedance spectroscopy (EIS) was examined at OCP in artificial saliva with sinusoidal amplitude of 10 mV in the frequency ranging from 10 kHz to 10 mHz.

2. Materials and methods

2.4. Biocompatibility evaluation

2.1. Fabrication of thin film and structure characterization

The cytotoxicity assay of the film was carried out according to ISO 10993-5 specifications [24]. The L929 cells, mouse fibroblast cell line, were used to assess the cytotoxicity of the extracts from the tested specimens. The specimens were maintained in Dulbecco's modified Eagle's medium (DMEM) in a 5% CO2 incubator at 37  C for 72 h. The extracts of the test specimens were subsequently used to treat cell monolayers for 24 h, after which the cells were examined for morphological changes to verify the toxicity scores. Meanwhile, the media were collected to quantify the released Cu and Al ions, having potential biological side effects, from the test specimens. All of the tests were performed in triplicate. The ions were measured using an inductively coupled plasma mass spectrometer (ICP-MS) (Perkin-Elmer, PE Sciex Elan 6100 DRC) with the detection limits of 0.0009 ppm for Cu ion and 0.0036 ppm for Al ion. The cell viability was investigated using a 3-(4,5dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT) assay and a microplate photometer (wavelength ¼ 570 nm) to measure the optical density (OD). Higher OD values indicated greater cell viability. DMEM without extract was used as a blank control. Treatment with 10% dimethyl sulfoxide in DMEM served as a positive control (PC). Cell adhesion test was conducted using human bone marrow mesenchymal stem cells (hMSCs), which were transduced using the gene for green fluorescent protein (GFP) through retroviral delivery. The GFP-labeled hMSCs were cultured on test specimens at a density of 5  104 cells/specimen. The culture medium contained RPMI-1640 medium supplemented with 5% fetal bovine serum and 10% horse serum. During 24 h cell incubation, cells adhering to the test specimens were observed in situ using a fluorescence microscope (AZ100, Nikon, Japan).

A disk-shaped crystalline alloy-target with a nominal composition of Zr60.14Cu22.31Fe4.85Al9.7Ag3 (at. %) were fabricated by vacuum sintering with pure elemental powders (purity>99.99%). Thin film was deposited on as-polished Ti6Al4V substrate by direct-current magnetron sputtering in pure Ar (purity>99.99%) gas. The detailed sputtering parameters are: base pressure of 2.0  104 Pa; working pressure of 0.65 Pa; sputtering power of 120 W; and bias voltage of 150 V. The microstructure of the films prepared was examined by field emission scanning electron microscope (FESEM, sirion 200), atomic force microscopy (AFM, SPM9700), X-ray diffraction (XRD) using Cu-Ka radiation and transmission electron microscopy (TEM, JEOL 2100). For TEM sample preparation, a NaCl crystal plate was used as the substrate, from which free-standing thin film (~50 nm in thickness) was obtained by dissolving NaCl substrate in deionized water. 2.2. Mechanical property measurements The hardness and elastic modulus of the film was measured by a nanoindentation system (Hysitron) using a Berkovich 142.3 indenter and in the displacement control model. Hardness value was calculated from a depth of less than 10% of film thickness to exclude any substrate effect [10] and the average hardness was obtained over at least 16 measurements. The adhesion of thin film was evaluated by micro-scratching using a 100 mm-radius diamond probe on a tribo-meter (UMT-3). Fretting resistance of the film was studied by ball-on-disk reciprocating sliding in air and in artificial saliva (with a composition of 0.4 g/L NaCl, 0.4 g/L KCl, 0.91 g/L CaCl2$2H2O, 0.69 g/L NaH2PO4$2H2O, 1.0 g/L CO(NH2)2, 0.005 g/L Na2S$9H2O, and pH ¼ 6.8 [22]). The fretting frequency was set to 1 Hz with a stroke of 0.5 mm and total duration of 2000 s. Si3N4 balls with a diameter of 6.35 mm were chosen as the counterpart because of its high hardness and good chemical stability. The normal load applied is 120 mN, corresponding to the contact pressure generates in the dental implant-supported system [23]. Each test was repeated at least five times to ensure data reliability. 3D optical measurement system (Keyence VHX-2000) and SEM/ EDX were used to determine the worn volume and wear morphologies after the fretting tests. Wear rate (Q, mm3 N1 m1) was then calculated using the equation of Q ¼ V/FzS, where V is the wear volume, Fz and S means the applied load and the total sliding distance, respectively. 2.3. Biocorrosion studies The biocorrosion behavior of the TFMG was evaluated by

3. Results and discussion 3.1. Microstructure of the thin film Fig. 1 (a) shows the cross-sectional view of a 530-nm thick TFMG prepared by magnetic sputtering, which has a composition of Zr63.59Cu20.75Fe3.12Al9.36Ag3.18 (at%). The composition determined by EDX is almost the same to its nominal composition (Zr60.14Cu22.31Fe4.85Al9.7Ag3). The AFM micrograph shown in Fig. 1(b) reveals uniform topography of the TFMG, with an average roughness Ra down to ~1.66 nm. By comparison, the as-polished Ti6Al4V alloy has a higher Ra of ~4.37 nm (not shown here). The amorphous state of the thin film was checked by XRD and TEM. The broad peak without any crystalline Bragg peaks in the XRD pattern (Fig. 1(c)), and the featureless TEM image as well as the diffusive ring pattern in selected area electron diffraction (SAED) pattern confirm the monolithic amorphous structure of the TFMG fabricated.

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Fig. 1. (a) A cross-section SEM image and EDX result of as-deposited ZrCuFeAlAg TFMG; (b) AFM morphology; (c) XRD pattern; and (d) bright-field TEM image and selective area electron diffraction (SAED) of the TFMG.

3.2. Hardness, elastic modulus and adhesion Hardness and elastic modulus of the TFMG coated Ti6Al4V substrate were measured by nanoindenter, which yields an average hardness (H) of 7.0 ± 0.3 GPa for the TFMG and 4.6 ± 0.3 GPa for the Ti6Al4V alloy, as shown in Fig. 2(a). Meanwhile, the elastic modulus (E) of the TFMG is around 116.1 ± 1.8 GPa, which is comparable to the Ti6Al4V alloy (E ¼ 126.6 ± 2.7 GPa) as well as those as-sputtered hydroxyapatite coatings (E ¼ 120 GPa) [25]. Adhesion strength between the thin film and Ti6Al4V substrate was evaluated by using a micro-scratch method, wherein the critical load leading to the sudden increase of coefficient of friction (COF) is regarded as the adhesion strength of the film [26]. The normal load (Fz) and COF as a function of the scratch distance is plotted in Fig. 2 (b), from which the adhesion strength for the 530 nm-thick TFMG is determined to be about 250 mN. This value compares the reported adhesion in Zr-Cu-Al-Ag TFMGs (~210 mN) [27]. Noted that good adhesion was achieved for the samples deposited using a relative high negative bias voltage (150 V) during sputtering, in which higher negative bias voltage yields a higher kinetic bombardment energy of the ions arriving on the substrate [28]. Those thin films fabricated using lower bias voltage (i.e., 50 V and 100 V) or without bias voltage have an adhesion strength lower than 110 mN. 3.3. Tribological performance of the TFMG in air and in artificial saliva In order to assess the tribological performance of the ZrCuFeAlAg TFMG, fretting wear tests were performed using ball-on-plate method (as schematically illustrated in the inset of Fig. 3(a)) in

air and in artificial saliva (simulate mouth environment). For comparison, the Ti6Al4V alloy was tested under the same experimental conditions. Fig. 3(a) shows the change of coefficient of friction (COF) with fretting duration for the TFMG and Ti6Al4V alloy under dry-sliding condition. As seen, the COF values of the two specimens attains the steady-state regime after a very short running-in period. The TFMG shows relatively higher steady-state COF (~0.52) but smaller COF fluctuation than Ti6Al4V alloy. The corresponding wear rate (Q) was calculated based on the asmeasured worn volume. Although the TFMG exhibits a higher COF, its wear rate is only 4.48  104 mm3 N1 m1, about 2.8-times lower than that of the Ti6Al4V alloy. Fig. 3(b) illustrates the COFs and wear rates under wet-sliding condition. The two specimens display similar COFs of around 0.75 in artificial saliva, both of which are higher than the COF values tested in air. It is generally believed that medium could have a lubrication effect on the COF, i.e., smaller COF under wet sliding. The higher COF observed here could be related to much smaller normal load (Fz ¼ 120 mN) used in this study as compared to most reported studies (Fz > 1 N). As the COF equals to Ft/Fz (Ft is the tangential force), the change of Ft is significant if an ultra-small Fz was used. This is exactly the case under wet fretting, wherein the solution resistance and passivation/repassivation resist the lateral movement of the indenter (i.e., Ft) and contribute to the increase of COF. Similar phenomenon was also observed by Seo et al. [29]. The mean wear rates under wet-sliding were 7.12  104 mm3 N1 m1 and 1.25  103 mm3 N1 m1 for the TFMG and Ti6Al4V alloy, respectively. Again, the TFMG exhibits approximately 1.8-times increase in the fretting resistance over Ti6Al4V alloy in the artificial saliva. There is no distinct correlation between COF and wear rate, i.e., even though the TFMG has higher

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Fig. 2. (a) Hardness and elastic modulus for the TFMG measured by nanoindention, as compared to the Ti6Al4V alloy. (b) Determination of adhesion strength of the TFMG with the Ti6Al4V substrate by a micro-scratch testing, in which a steep increase of coefficient of friction (COF) suggests the breakdown of the film. The critical load (Fz) is therefore regarded as the adhesion strength.

COF, it exhibits better fretting resistance. Similar phenomenon was also reported in the Zr61Ti2Cu25Al12 BMG system [30]. To understand the wear mechanism of the TFMG in detail, worn scar after fretting tests were further examined by SEM/EDX, as shown in Fig. 4. Evidently, wear traces on the TFMG are much slight than those on the Ti6Al4V in both dry-sliding and wet-sliding conditions, which is in good accordance with the obtained wear rate values (see Fig. 3). Regarding fretting in air, heavy abrasive furrow and oxygen traces (see EDX result in Fig. 4(a)) was observed in Ti6Al4V alloy, typical features of abrasive and oxidational wear mechanisms. On the contrary, no visible furrow except some debris (white-contrast phases) was observed on the TFMG after fretting. EDX analysis revealed that those debris contain high oxygen content (~35 at%) and major metals of 24 at% Zr and 24 at% Ti (Ti was mainly from the substrate), in addition to some other minor metals of Cu, Al, Fe, all of which were listed in the inset of Fig. 4 (b). The asformed debris on the TFMG were therefore mainly Zr-oxides, indicating that the wear behavior of this Zr-based TFMG is dominated by an oxidation-wear mechanism during dry sliding. In such case, the wear resistance is related to the generation and breakdown of an oxide layer formed on the outermost surface during

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Fig. 3. The COFs as a function of sliding time for the TFMG and Ti6Al4V alloy. (a) dry sliding in air and (b) wet sliding in simulated saliva. Insets show the schematic illustration of fretting experiment and as-calculated wear rates for the TFMG and Ti6Al4V alloy after the fretting tests.

reciprocating friction. Considering the high chemical affinity of Zr with oxygen (DHmixz 1097 kJ/mol at 298e900 K [31]), it is not supervising Zr-oxides were preferentially formed on the TFMG during fretting, which were broken and delaminated as debris in the continuous fretting. A similar oxidation-controlled wear mechanism was also reported for Zr42Ti15.5Cu14.5Ni3.5Be24.5 BMG [32]. Regarding the fretting mechanism in artificial saliva, a cracked tribo-layer in white-contrast was clearly observed on the Ti6Al4V, as shown in Fig. 4(c). The tribo-layer contains higher concentration of oxygen in addition to a slight Ca from the solution, which was different from the oxide layer formed during dry sliding. It has been reported that those repassivated tribo-layers on Ti6Al4V had a lower hardness (<250 Hv) than that naturally formed oxide layer [33]. Therefore, the tribo-layer is easy to cracking due to its low hardness and weak resistance to shear deformation. By contrast, there is no visible cracks on the surface of TFMGs, indicating superior protectiveness of the tribo-layer formed. The chemical composition of the tribo-layer is very similar to that formed in air, suggesting the same wear mechanism (i.e., oxidative wear) dominated the fretting behavior of the present Zr-based TFMG in artificial saliva.

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Fig. 4. Worn surface morphologies for the Ti6Al4V alloy after fretting tests in air and in artificial saliva. (a) Ti6Al4V and (b) TFMG after dry sliding; (c) Ti6Al4V and (d) TFMG after wet sliding. Insets show EDX results on the wear tracks.

Fig. 5. (a) Potentiodynamic polarization curves of the TFMG and Ti6Al4V alloy in the artificial saliva; (b) potentiostatic polarization plots under a constant applied potential of 0 V(SCE). Inset is the double-log plots of current-time curves for the two specimens. (c) EIS spectra diagrams for the TFMG and the Ti6Al4V alloy in artificial saliva.

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Fig. 6. (a) Cytotoxicity assay results showing the morphology of L929 cell after 72 h of incubation in extract-containing media; (b) cell viability of L929 cells co-cultured with extracts obtained from the Ti6Al4V alloy and Zr-based TFMG specimens.

3.4. Biocorrosion behavior in artificial saliva The human mouth presents an aggressive environment for metals used for therapy. Therefore, electrochemical polarization was conducted to study the biocorrosion resistance of the TFMG in artificial saliva, as compared to the Ti6Al4V alloy. As seen in Fig. 5(a), the corrosion current density (icorr) of the TFMG is nearly one order of magnitude lower than that of Ti6Al4V alloy, indicating a lower electrochemical activity. Corrosion rates were calculated by fitting Tafel regions, which yield 0.02 ± 0.01 mm/y and 0.30 ± 0.04 mm/y for the TFMG and Ti6Al4V alloy, respectively. This demonstrates a better general corrosion resistance of the TFMG over the Ti6Al4V in artificial saliva. Moreover, the TFMG is spontaneously passivated with very low passive current density on the order of 108e105 A cm2 and a distinct passive region (with a width of ~1.56 VSCE) before pitting. For the Ti6Al4V alloy, pitting corrosion did not occur, indicative of high pitting resistance. It is

worthy to be noted that the pitting potential (Epit) of this Zr-based TFMG is higher than 1.0 VSCE in Cl-containing solution, which far exceeds the values reported in Zr-based BMGs tested in the same composition of artificial saliva (i.e., Epit~0.125 VSCE for Zr50Cu43Al7 system [34] and Epit~0.787 VSCE for Zr62.5Cu22.5Fe5Al10 system [35]). The superior pitting resistance for as-sputtering TFMG could be ascribed to its homogenous chemistry, low surface roughness and uniform microstructure in the absence of defects and/or nanocrystalline phases, which eliminate the microgalvanic corrosion effect. By contrast, factors such as casting defects or Cu-rich nanocrystals would trigger pit initiations on Cu-bearing Zr-based BMGs [35,36]. The stability and protectiveness of the passive film formed were evaluated by potentiostatic polarization and EIS testing. The current (i)-time (t) curves of the two specimens tested at a constant potential of 0 VSCE are shown in Fig. 5(b). Compared to Ti6Al4V, the TFMG shows a lower passive current density (~108e107 A cm2)

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over the testing period, verifying more stable passive film formed. The passivation kinetic was studied in i-t double logarithmic curves, as shown in the inset of Fig. 5(b). According to the equation of i ¼ At n [37], where A is a constant, t is the polarization duration and n is the passivation rate, n ¼ 1 generally signifies the formation of a compact, highly protective passive film under high electronic-field controlled growth mode (typical for valve metals). As seen, both specimens show two-stages of growth of the passive film. In the initial stage, the values of n are 1 for TFMG and 0.1 for Ti6Al4V, indicating a faster formation rate of the passive film for the former. Later (after 100 s), the passive film on the TFMG is slowed down; while the passive film on Ti6Al4V exhibits an accelerated growth with n ¼ 0.7. Fig. 5(c) shows the Nyquist plots of the TFMG and Ti6Al4V alloy in artificial saliva. As seen, the TFMG exhibited larger diameter of semi-circles in the Nyquist plot, again verifying the superior barrier resistance over the Ti6Al4V. These results agree well with the polarization curves. 3.5. Cytotoxicity and cell adhesion assessments Optical microscopy images of the TFMG and uncoated Ti6Al4V alloy in contact with L929 cells co-cultured with extracts are shown in Fig. 6(a). For comparison, a blank group and a positive control (PC) group are also presented in the figure. In the PC group, the L929 mouse fibroblast cells indicated a positive cytotoxic reaction, wherein the cells showed grainy and lack normal cytoplasmic spacing. In addition, the large open areas between the cells are an indication of extensive cell lysis (disintegration). Conversely, the blank group appeared a confluent monolayer of well-defined L929 mouse fibroblast cells displaying cell-to-cell contact. This appearance is an indication of a non-cytotoxic response. L929 cells cultured in extract-containing media (TFMG and Ti6Al4V) for 72 h were not morphologically different, demonstrating similar cell viability as cells cultured without the extract (blank control). The results indicated that the present Zr-based TFMG belongs to noncytotoxic materials, similar to the well-known biocompatible Ti6Al4V alloy. The cell viability of four groups assessed by measured the viability of L929 cells that were cultured in the extractcontaining media for 72 h is shown in Fig. 6(b). It can be seen that the TFMG group exhibits the same level of viability (100%) as Ti6Al4V and blank groups, demonstrating the good in vitro biocompatibility of the present Zr-based TFMG. After the 72 h cytotoxicity test, the accumulated amounts of Cu and Al ions in the media were approximately 5.569 ppm and not detectable, respectively. It has been reported that the maximum Cu content in the metallic glass alloy for better corrosion resistant is suggested below 5 at% [38]. Although the Zr-based TFMG used in this study contained 22.3 at% Cu, the accumulated amount of Cu ion release after 72 h cytotoxicity test was approximately 5.569 ppm, i.e., 1.856 ppm per day, which was still far below the safe level (10e12 mg for adult per day) suggested by the World Health Organization [39]. To further investigate the cell adhesion morphology of the test specimens, the GFP-labeled hMSCs were seeded onto the TFMG and Ti6Al4V alloy. Fig. 7 shows the morphologies of green fluorescent protein of hMSCs attached onto the Ti6Al4V alloy (Fig. 7(a)) and TFMG (Fig. 7(b)) after cell incubation for 24 h. Fluorescent micrographs revealed the efficient attachment and good extension morphology of the hMSCs on the TFMG and Ti6Al4V alloy, indicating a good cell adhesion to the material surface. For the two specimens, GFP signal increased over the course of the cell incubation period, indicating that cell growth was taking place. Compared to the Ti6Al4V alloy, the cell density on the TFMG is similar; while a more flattened cell adhesion morphology and overlapping of adjacent cell is observed on the TFMG. This signifies a better adhesion and healthy growth of cells on the present

Fig. 7. The cell morphologies of GFP-labeled hMSCs on (a) Ti6Al4V alloy and (b) TFMG after cell incubation for 24 h.

biocompatible TFMG than on the Ti6Al4V alloy. 4. Conclusions A Ni-free Zr60.14Cu22.31Fe4.85Al9.7Ag3 TFMG was synthesized using single-target magnetron sputtering on biomedical Ti6Al4V substrate. The as-deposited TFMG consists of a single glassy structure, with a hardness of ~7 GPa and modulus of ~116 GPa. The fretting wear rate of the TFMG is lower than Ti6Al4V alloy, i.e., 2.8times lower in dry-sliding condition and 1.8-times lower in wetsliding condition. The TFMG also exhibit excellent biocorrosion resistance in artificial saliva, with a corrosion rate lower than Ti6Al4V and a high pitting potential of above 1.0 VSCE. The high biocorrosion resistance of the TFMG is attributed to its fast formation rate and high stability of the passive film formed. The TFMG shows no any cytotoxicity and good cell adhesion with hMSCs, as good as Ti6Al4V alloy, confirming the good biocompatibility of the TFMG. Owing to the achievement of favorable mechanical properties, corrosion resistance and good biocompatibility, the Zr60.14Cu22.31Fe4.85Al9.7Ag3 TFMG can be potential candidate as novel biomedical metallic coating for dental applications. Acknowledgement This work was financially supported by the National Nature Science Foundation of China (Grant No. 51301072, 51471074 and 51531003) and State Key Laboratory of Materials Processing and Die

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