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Tribological and biological behaviors of laser cladded Ti-based metallic glass composite coatings
Applied Surface Science 507 (2020) 145104
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Full Length Article
Tribological and biological behaviors of laser cladded Ti-based metallic glass composite coatings
T
Hong Wua,b,c, , Luxin Lianga, Xiaodong Lana, Yong Yina, Min Songa, Ruidi Lia, Yong Liua, ⁎⁎ Haiou Yangc, Ling Liud, , Anhui Caie, Qingxiang Lib, Weidong Huangc ⁎
a
State Key Laboratory of Powder Metallurgy, Central South University, Changsha 410083, PR China Shenzhen Zhong Jin Ling Nan Nonfemet Co., Ltd, Shenzhen 518040, PR China c School of Materials Science and Engineering, Northwestern Polytechnical University, Xi’an 710072, PR China d Hepatobiliary and Pancreatic Surgery Department, Xiangya Hospital, Central South University, Changsha 410008, PR China e College of Mechanical Engineering, Hunan Institute of Science and Technology, Yueyang 414000, PR China b
ARTICLE INFO
ABSTRACT
Keywords: Laser cladding Metallic glass composite coating Tribological behavior Biological behavior
In order to improve the wear resistance of pure Ti, three kinds of metallic glass composite coatings (MGCCs) with the compositions of Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3, were deposited on pure Ti by laser cladding. It is found that the order of amorphous phase percentage is Ti51Zr5Cu41Sn3 > Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9. Compared to the pure Ti substrate, the wear resistances of three coatings were significantly improved in simulated body fluid (SBF) solution. The wear resistance of MGCCs in SBF solution is in the order of Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9 ≥ Ti51Zr5Cu41Sn3, which is attributed to the combination of fine grain strengthening and well corrosion resistance. XPS results imply that these coatings enable to induce the deposition of calcium phosphate which is excellent responsible for the bone formation. The CCK-8 test and cell adhesion experiments show that compared to pure Ti, the Ti51Zr5Cu41Sn3 coating was free of cytotoxicity, while Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings have obvious cytotoxicity in vitro due to the high release rate of Cu and Ni ions from coatings. It is suggested that the Ti51Zr5Cu41Sn3 coating had not only well wear resistance, but also excellent biocompatibility.
1. Introduction Biomedical metals and alloys have been widely used in hard tissue (such as artificial bone, arthrosis and teeth) repair and implants [1]. Titanium (Ti) and its alloy have been extensively applied to orthopaedic implants for the good biocompatibility, together with excellent mechanical properties [2,55]. However, their service life is significantly limited due to the poor wear resistance. Loosening of the prosthesis due to sliding wear causes 80% of joint prostheses to require secondary surgery [1]. In addition, large number of wear debris and metallic ions such as Al and V for Ti-6Al-4V alloy are released into body tissue, which tend to provoke adverse reactions, like allergy and toxicity [3,4]. To improve wear resistance of Ti and its alloys, a variety of surface modification technologies were exploded nowadays, such as physical deposition methods by plasma spraying [5] and physical vapor deposition [6], chemical surface treatments like nitriding [7] and carburization [8]. However, the coatings prepared by plasma spraying and
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physical vapor deposition are easily exfoliated in case of frequent friction, because they are only mechanically bonded to the metallic substrate with low bonding strength [9]. Although nitriding and carburization can improve the bonding strength of the coating, both their complex operating processes and the easy deformation of the workpiece limit their widespread applications [9]. It is well-known that the atoms of bulk metallic glasses (BMGs) are disorderly arranged, which exhibits excellent performance on corrosion resistance, mechanical strength and wear resistance. These inspired properties make BMGs to possess great potential to be used as biomaterials [10-12]. Among the BMGs systems, Ti-based MG alloys have become a preferred candidate for structural materials or biomaterials because of their good wear resistance and excellent biocompatibility [12]. The BMGs exhibit much lower coefficient of friction (COF) and better anti-wear resistance than pure Ti [13]. More importantly, the results of in vivo implantation for 3 months present that the Ti41.5 Zr2.5 Hf5 Cu37.5 Ni7.5Si1 Sn5 BMGs are firmly bonded to the alveolar bone
Corresponding author at: State Key Laboratory of Powder Metallurgy, Central South University, Changsha 410083, PR China. Corresponding author. E-mail addresses: [email protected] (H. Wu), [email protected] (L. Liu).
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https://doi.org/10.1016/j.apsusc.2019.145104 Received 25 July 2019; Received in revised form 28 November 2019; Accepted 16 December 2019 Available online 19 December 2019 0169-4332/ © 2019 Elsevier B.V. All rights reserved.
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tissue [13,14]. Nickel ions are biologically toxic, that would cause series of harmful physiological and allergic reactions [15]. Therefore, Ni-free Ti-based amorphous alloys have been developed to meet the needs of biomedical applications [16,17]. The wear resistance as well as corrosion resistance of the Ti47Cu38Zr7.5Fe2.5Sn2Si1Ag2 BMG are significantly superior to that of the Ti6Al4V alloy either in air and in phosphate buffer saline (PBS) solutions [16]. However, the size of Tibased amorphous alloy fabricated by copper mold casting cannot meet the requirement of biological material because of processing technology. In recent years, laser cladding technique has been widely used in the preparation of BMGCs due to the advantages of rapid heating and cooling [18-20]. As laser cladding technique involves a repetitive cycle of heating and cooling, the temperature field of the current sedimentary layer will have annealing effect on the already deposited region, leading to the structural relaxation as well as crystallization of the deposited MG [21,54]. Zhang et al. [19]fabricated five layers Zr-based BMGCs by laser solid forming, in which the percentage of amorphous is still 81.5 vol%. Pauly et al. [20] also prepared large sized Fe-based MG with lower crystallization by selective laser melting. To overcome size limitations of amorphous alloy, Lan et al. [22] developed a process for depositing metallic glass composite coatings (MGCCs) on Ti alloys by laser cladding technique. Meanwhile, the metallurgical bonding strength between the clad layer and the metallic substrate is strong [19,22]. The microhardness and wear resistance of the coating can be significantly improved owing to the precipitation of the crystallite from the amorphous alloy [22,23]. In addition, it is reported that stannum (Sn) element is effectively improve the glass forming ability [24], plastic strain [25], corrosion resistance and biocompatibility of metal materials [26]. Therefore, it is promising to prepare Ni-free Ti-based MGCCs with Sn on pure titanium by laser cladding. The wear resistance is considered to be very important for laser clad MGCCs used for biomedical metals. It is well accepted that, the friction and wear behaviors are not the intrinsic properties of material, but strongly dependent on the wear conditions and testing parameters, such as sliding speeds, applied loads, sliding distances, and lubrication conditions, etc. [27,28]. In our previous study, it was found that the wear rate of the Zrbased BMG increased significantly with the increase of oxygen content in testing environment [27,28]. However, reported literatures on both tribological behaviors in SBF and biological properties of laser cladding Ti-based MGCCs are still inadequate. In the present study, three kinds of Ti-based MGCCs (Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3) were deposited on the surface of pure Ti by laser cladding. In order to evaluate whether they could meet the requirements of orthopedic implants, the tribological behaviors in SBF, corrosion resistance and biological properties of three kinds of MGCCs were systematically investigated.
1 mm, single beam pulse energy 143 J, pulse duration 6 ms, pulse frequency 20 Hz, scanning speed 5 mm/s, and overlap ratio 50%. The thickness of each layer was about 0.2 mm. 2.2. Coating characterization The coatings were cut into test samples by electrical discharge machine. The microstructure of the Ti-based MGCCs was observed by scanning electron microscopy (SEM, FEI Quanta FEG 250). The phases of laser cladded Ti-based MGCCs were analyzed by X ray diffractometer (XRD; Rigaku D/MAX-2250) with a Cu Kα radiation and scanning step of 0.02°. The thermal stability of the deposits was investigated by differential scanning calorimeter. The temperature of the DSC test ranges from room temperature to 1000 K. The experiments were carried out under argon atmosphere with a heating rate of 0.67 K/s. The distributions of relevant elements were analyzed by electron probe microanalyzer (EPMA; JXA-8230, Japan). The transmission electron microscope specimens were prepared using the focused ion beam (FIB). The specimens were characterized by transmission electron microscopy (TEM; Tecnai G2 F20 S-TWIN, USA). 2.3. Vickers microhardness The Vickers microhardness of various specimens was measured by the Buehler Micromet 5104 microhardness tester with a load of 100 g for 30 s. All experiments were repeated three times. 2.4. Tribological behaviors in SBF The tribological properties of pure Ti and Ti-based MGCCs were measured by tribotest device (UMT-3, USA) in SBF solution (CZ0400, China) at room temperature. The ion concentrations in SBF solution are showed in Table S1 in supplement information. Si3N4 balls with diameter of 6 mm were selected as counterpart. Loading conditions of 20 N was applied, with the linear speed being 50 mm/s and wear time being 30 min. The wear rate ω is determined by the following equation:
V = Mloss/P = V /(W L)
(1) (2)
where V is the wear volume, M is mass loss after the test (g), P is the density of the material (g/cm3), W is the load (N) and L is the total slip distance (m). 2.5. Electrochemical test The dynamic potential polarization curves were detected on Princeton electrochemical test system (Princeton Applied Research 4000). The pure Ti and Ti-based MGCCs were cut into square pieces of 10 mm × 10 mm × 1 mm and gradually ground by the metallographic grinding paper, and then they were cleaned in deionized water for ten minutes. After samples were dried at room temperature, electrochemical experiments were conducted. The working electrode was square sheets of pure Ti or Ti-based MGCCs. The alloy sheets were welded with copper wire and enclosed by the epoxy resin to make electrodes, with working area of 1 cm2. The saturated calomel electrode was used as reference electrode, and platinum sheet was auxiliary electrode. The voltage range was 0.25 V each around the initial voltage and the scan rate was 0.003 V/s. The dynamic potential polarization curves were tested in SBF solution at room temperature.
2. Materials and methods 2.1. Coating fabrication The circular bulks used as substrates was pure Ti with a size of Ф20 mm × 5 mm. The nominal components of gas-atomized powders for laser cladding were Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 (at%), respectively. The size of powder was less than 75 μm. Laser cladding was conducted in glove boxes under argon atmosphere (oxygen content less than 10 ppm) in order to prevent any oxidation reaction. Before laser cladding, the surface of pure Ti samples was polished with 800 mesh sand papers. During the cladding, a powder layer with the thickness of 0.2 mm was firstly laid on the pure Ti (Ф 20 mm), then the powder layer was melted point by pulsed laser beam and solidified to form a solid deposited layer. The specific experimental procedures are consistent with the literature [22]. Briefly, the 300 W Nd: YAG pulsed laser was used for multiple clad. Laser processing parameters were wavelength 1.06 μm, laser beam diameter
2.6. Analysis of MGCCs after immersed in SBF The surface chemical states of the specimens immersed in SBF solution for 7 days were evaluated using SEM, EDS and the X-ray photoelectron spectroscopy (XPS; K-ALPHA, Thermo Fisher Scientific, 2
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Britain). In order to further understand the form of elements in the coatings, Ca and P peaks of the coating surfaces were emphatically analyzed through XPS. The test results were analyzed and processed by XPSPEAK4.1 software.
Fig. 1(a). Three coatings displayed broad diffraction maxima at 2θ = 41°, corresponding to amorphous phase and several sharp diffraction peaks are related to crystalline phases. The crystalline phases contain Ti2CuNi phase for Ti45Zr5Cu41Ni9 coating and Ti45Zr5Cu41Ni6Sn3 coating, and TiCu phase for Ti51Zr5Cu41Sn3 coating. With the addition of Sn element, the peak of amorphous phases of Ti51Zr5Cu41Sn3 coating were broader than that of Ti45Zr5Cu41Ni6Sn3 coating and Ti45Zr5Cu41Ni9 coating. The crystallinities of three coatings calculated by the RulandVonk method are in the order of Ti45Zr5Cu41Ni9 (51.45%) > Ti45Zr5Cu41Ni6Sn3 (39.48%) > Ti51Zr5Cu41Sn3 (29.21%). These results are consistent with our previous results [22]. No oxide was detected on any of the coating surfaces, indicating that the oxygen content of the coating is low. In order to further determine the existence of an amorphous phase in coating Ti51Zr5Cu41Sn3, the DSC traces revealing the glass transition temperature Tg and the onset temperature of crystallization Tx was provided, as indicated in Fig. 1(b). The deposited coating of Ti51Zr5Cu41Sn3 exhibited ΔT = 30 K with Tg = 673 K and Tx = 703 K. To further analyze the microstructure, the surface and cross sections of different coatings were observed by SEM in Fig. 1(c) and supplementary information Fig. S1. The structures of coating surfaces were homogeneous with some flower-like dendrites. Obviously, the dendrite size in the MGCCs was in the order of Ti45Zr5Cu41Ni9 > Ti45Zr5Cu41Ni6Sn3 > Ti51Zr5Cu41Sn3, which is consistent with crystallinity. The elemental distributions of Ti, Zr, Cu, Ni and Sn are exhibited in Fig. 1(c). The distribution of elements in the amorphous zone is relatively homogeneous, without obvious segregation of components exists for various coatings. It is necessary to point out that the content of Ti, Ni and Sn element in the dendrite are higher than those in the amorphous zone for Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3. The concentration of Ti element in the plate-shaped dendrites is higher than amorphous for Ti51Zr5Cu41Sn3. In our previous work, high resolution transmission electron microscopy and selected area electron diffraction were employed to characterize the Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3, indicating that the Ti2CuSn nanocrystals are dispersed in the Ti2CuNi dendrites in Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 [22]. In order to investigate the dendrite phase in Ti51Zr5Cu41Sn3 coating, the microstructure of the specimen was characterized using TEM, as shown in Fig. 2. The amorphous phase S1 in Fig. 2(a) was further confirmed by the typical halo ring patterns in Fig. 2(b). Simultaneously, some nanocrystal phases embedded in the amorphous matrix are identified as TiCu phases [30]. The dendrite phase S2 (Fig. 2(a)) is identified as TiCu phase using selected area electron diffraction pattern in Fig. 2(c). The TEM results are consistent with the XRD patterns. Additionally, as displayed in Fig. S1, the microstructure of the coating was epitaxial to the substrate by columnar grains, implying a good fusion bond between the substrate and coating.
2.7. Metal ions release To evaluate the releasement of Cu and Ni ions, the specimens of coatings were cut into the size of 10 mm × 10 mm × 1 mm by wire electrical discharge machining, and polished with 2000 mesh sand papers. Then the coatings were immersed in 10 ml phosphate buffered saline (PBS) for various days (1, 3, 7 and 14 days) in constant temperature incubator with 37 °C. The solutions containing Cu and Ni ions were detected by the inductively coupled plasma atomic emission spectrometry (ICP-AES). 2.8. SaOS-2 cell culture and proliferation Human osteoblastic cells (SaOS-2), which were acquired from China Infrastructure of Cell Line Sources, were cultured in McCoy’s medium containing 15% fetal bovine serum and 1% penicillin/streptomycin solution. Cells were cultured in incubator with 5% CO2 and 100% humidity at 37 °C. To assess cell proliferation behavior, the SaOS-2 cell suspension with a concentration of 30,000 cells per well were introduced on the coating surfaces for 1, 3 and 5 days. Before the cells were seeded in 24well plates, all disks with the size of 10 mm × 10 mm × 1 mm were polished with 800 mesh sand papers, and were sterilized using ultraviolet rays for 48 h. Fresh medium and CCK-8 solution with a volume ratio of 10:1 was added into each well. After incubated for 1 h at 37 °C, 100 μL of the incubation solution was added to a 96-well plate, and the absorbance was detected by a microplate reader at 450 nm. 2.9. Live cell staining The live cell staining was executed according to manufacturer’s instructions. Briefly, SaOS-2 cells were seeded on the surface of coatings with a concentration of 30,000 cells per well, and cultured for 1, 3 and 5 days. Cells were washed with PBS 3 times. 2 μM calcein-AM was added in each well, and cells were cultured at cell incubator for 15 min. Fluorescence microscopy (DFC420C, Leica, Germany) was used to observe and photograph. 2.10. Cell morphology For the observation of morphology, SaOS-2 cell suspension with a concentration of 30,000 cells per well were cultured on the various coating surfaces and cultured for 1 day. To prepare the samples for SEM, the specific preparation method was consistent with the literature [29]. Briefly, the cells were first fixed with 2.5% glutaraldehyde, and dehydrated with different concentration of alcohol gradient. Then ethanol was replaced with different concentration of tert-butanol, and finally the samples were freeze-dried.
3.2. Microhardness The microhardness for the surface of specimens is shown in Fig. S2. The Vickers hardness for Ti is only HV100 156.9, while the microhardness for Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 coatings are HV100 590.1, 663.6 and 515.5, respectively. The microhardness of Ti is much lower than that of MGCCs due to the disorder arrangement of atoms and strengthening of nanocrystals [31]. Previous studies indicated that the hardness of Ti2CuNi phase is higher than that of amorphous alloy, and the hardness of TiCu is lower than that of amorphous alloy [32]. Therefore, the hardness of coating Ti45Zr5Cu41Ni9 is higher than that of Ti51Zr5Cu41Sn3 coating. We previously proved that there are fine Ti2CuSn nanocrystals dispersed in the Ti2CuNi phase in the Ti45Zr5Cu41Ni6Sn3 coating. The nanocrystals have a strengthening effect on dendritic Ti2CuNi, so the hardness of Ti45Zr5Cu41Ni6Sn3 coating is higher than that of Ti45Zr5Cu41Ni9 coating [22].
2.11. Statistical analysis All experiments were repeated three times and the data were expressed as mean ± standard deviation. To compare the differences in significance between the study groups, the two samples t-test was applied. The value of p < 0.05 was defined as statistically significant. 3. Results 3.1. Surface microstructure analysis The XRD patterns of the deposited coatings surfaces are presented in 3
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Fig. 1. (a) XRD patterns of coatings surfaces. (b) DSC curve of the coating Ti51Zr5Cu41Sn3. (c) SEM BSE and EPMA micrographs showing the surface microstructures of the coatings.
0.2288 × 10−3 mm3 N−1 m−1, and 0.5101 × 10−3 mm3 N−1 m−1, respectively. The three coatings exhibited lower COF than control group (pure Ti) in the SBF solution. The wear loss of three coatings were an order of magnitude less than the control group. The SEM morphologies of surfaces after wear in SBF solution are presented in Fig. 3(c–f). The worn surface of the pure Ti are consisted of numerous deep grooves and wear debris in the worn scar, revealing the local severe deformation. Compared to the pure Ti, the worn surfaces of the MGCCs in Fig. 3(d–f) display much shallower
3.3. Tribological behaviors in SBF solution As shown in Fig. 3(a, b), the friction coefficients of coatings Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 were stable under the wear time of 1800 s in SBF solution. However, under wet sliding condition in SBF solution, the wear rates of pure Ti, Ti45Zr5Cu41Ni9 coating, Ti45Zr5Cu41Ni6Sn3 coating and Ti51Zr5Cu41Sn3 coating were 5.9877 × 10−3 mm3 N−1 m−1, 0.4667 × 10−3 mm3 N−1 m−1, 4
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Fig. 2. The TEM macrographs (a) of Ti51Zr5Cu41Sn3, together with SAED patterns obtained from amorphous matrix S1 (b) and dendrite S2 (c).
addition, Cu2+ ions released from Ti51Zr5Cu41Sn3 coatings and Ti45Zr5Cu41Ni6Sn3 coatings are basically equal, while Cu2+ and Ni2+ ions released from Ti45Zr5Cu41Ni9 coating are significantly higher that from the coating Ti45Zr5Cu41Ni6Sn3 after immersion for 14 days.
grooves. For the Ti45Zr5Cu41Ni6Sn3 coating, the plastic deformation grooves were the shallowest. The surface of Ti45Zr5Cu41Ni9 coating was corroded after wear in SBF. The wear resistance of MGCCs is in the order of Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9/Ti51Zr5Cu41Sn3 > pure Ti. The SEM images of the wear debris are presented in Fig. 4. The EDS scanner was employed to analyze the elements of the wear debris (marked as Sa, Sb and Sc), which is displayed in Table 1. The results showed that all the wear debris contained oxygen element, indicating the formation of complex metallic oxides during the wet frictional process. Interestingly, oxygen content of wear debris was the highest in Sa of Ti45Zr5Cu41Ni9 coating, and the minimum oxygen was found in Sb of Ti51Zr5Cu41Sn3 coating. The debris exhibited a flat and slip surface without plastic deformation. It may be attributed to the deterioration of the combination between the oxides layer and the metal matrix in PBS solution, which promotes the refinement of wear debris [17]. In general, the tribological behavior of MGCCs in SBF solution may be influenced by the precipitation of the crystalline phase from the amorphous matrix [33,34], and solution corrosion [17]. In a word, the main wear mechanisms of pure titanium were adhesive wear and corrosion wear, while the MGCCs were dominated by abrasive wear and corrosion wear.
3.5. Analysis of coatings after immersed in SBF The surface morphology and chemical composition of the coatings after immersed in SBF for 7 d are shown in Fig. 7. Many small particles can be found on the surface of various coatings (Fig. 7(a–c)). EDS analyses show that the chemical composition of the precipitates contain Ca and P. More detailed analyses on the surface composition were conducted by XPS (Fig. 7(d–f)). As presented in Fig. 7(d), XPS spectra of the surface indicates that there are a few elements such as Ca, P, Ti and Cu in the coating surface. In order to further clarify the form of elements in the coatings, Ca peaks, P peaks of the coating surfaces were emphatically analyzed through XPS. Fig. 7(e) presents Ca2p (2p3/2 and 2p1/2) peaks in the surface, respectively located at 347.0 eV ± 0.2 and 350.6 eV ± 0.2, and Fig. 7(f) shows the P2p peaks in the surface located at 133.0 eV ± 0.2, indicating the presence of calcium phosphates [37,38]. These results agree well with the findings of previous study that Cu promote the calcium phosphate deposition [39].
3.4. Corrosion behavior and ion release of coating surfaces
3.6. Cell proliferation and adhesion
Fig. 5 displays the typical polarization curves of different coatings in SBF solution. The corrosion potentials (Ecorr) were in the order of coating > Ti45Zr5Cu41Ni6Sn3 coating > Ti51Zr5Cu41Sn3 Ti45Zr5Cu41Ni9 coating > pure Ti. It could be clearly seen that the corrosion current density (Icorr) of MGCCs samples was in the order of 10−5 A/cm2 in SBF solution, which is approximately equal to that of pure Ti, indicating that the corrosion resistance of three coatings is stronger than that of pure Ti. The Ti45Zr5Cu41Ni9 coating had the worst corrosion resistance of the three coatings. MGCCs are liable to corrosion in the complex humoral environment, resulting in the release of Ni2+ ions and Cu2+ ions. Researchers shown that the release of large amounts of Ni2+ ions and Cu2+ ions can cause cytotoxicity, tissue and organ disorders [35,36]. The amounts of Ni2+ ions and Cu2+ ions leached from the three coatings were tested by ICPAES, as shown in Fig. 6. The concentration of cumulative Ni2+ ions and Cu2+ ions increased with the extension of immersion duration. In
Proliferation of SaoS-2 cultured on the surface of coatings and pure Ti was assessed using Calcein-AM staining and CCK-8 reagent. As displayed in Fig. 8(a–b), the cell number on the surface of Ti45Zr5Cu41Ni6Sn3 coating and Ti45Zr5Cu41Ni9 coating was significantly less than that on the pure Ti for day 1, day 3 and day 5, while the cells number on surface of Ti51Zr5Cu41Sn3 coating and pure Ti was comparable. The cells number on various surfaces of coating increased from day 1 to day 5. Fig. 8(c) presents that the quantitative results detected by CCK-8 reagent is in accordance with the fluorescence images described in Fig. 8(a). According to the current ISO standards, cell viability higher than 75% could be considered with no toxic risks for medical devices, implying that the Ti51Zr5Cu41Sn3 and Ti45Zr5Cu41Ni6Sn3 coatings were almost non-toxic for SaOS-2 cells, while Ti45Zr5Cu41Ni9 coating had obvious toxicity. The morphology of SaOS-2 cells on different surfaces of MGCCs for 5
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Fig. 3. (a) Friction coefficient curves of coatings immersed in SBF under a normal loads of 20 N. (b) Mean values of friction coefficients of three coatings. The worn surfaces of Ti (c), Ti45Zr5Cu41Ni9 coating (d), Ti45Zr5Cu41Ni6Sn3 coating (e), and Ti51Zr5Cu41Sn3 coating (f). ** P＜0.01, compared to pure Ti. # P＜0.05, compared to Ti45Zr5Cu41Ni9 coating.
day 1, 3 and 5 was further observed by SEM. As shown in Fig. 9, after 1 day, SaOS-2 cells grown on surface of various coatings exhibited similar spreading area compared to pure Ti. The lamellipodium formation and filopodium extension were obviously observed on the surfaces of deferent coating. At day 3 and 5, a lot of cells grown on surfaces of the Ti45Zr5Cu41Ni9 coating showed obvious atrophic morphology, and some of the cells grown on surfaces of the Ti45Zr5Cu41Ni6Sn3 coating showed atrophic morphology. Almost no cells showed atrophic morphology on the surface of the Ti51Zr5Cu41Sn3 coating. These observations indicate that Ti51Zr5Cu41Sn3 coating show excellent biocompatibility, but
Ti45Zr5Cu41Ni9 coating has obvious toxicity for SaOS-2 cells. 4. Discussion 4.1. Phase composition In the present study, three kinds of Ti-based MGCCs, Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 were deposited by laser cladding. With the addition of Sn, the amorphous phase content increases. The order of amorphous phase percentage in as-designed three MGCCs is 6
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Fig. 4. SEM images of collected MGCCs wear debris: (a) Ti45Zr5Cu41Ni9 coating, (b) Ti45Zr5Cu41Ni6Sn3 coating and (c) Ti51Zr5Cu41Sn3 coating.
relatively large, which increases the mismatch of the atomic size which may delay the rearrangement of the atoms required for crystallization, thus enhancing the glass forming ability [40,41]. Secondly, appropriate negative heat of mixing between the main elements is beneficial to the deep suppression of the free energy curve of the supercooled liquid phase [42]. As the negative mixing enthalpy between Sn and Cu, Ti and Zr are higher than that between Ni and Cu, Ti and Zr, which contribute to inhibit the free energy of supercooled liquid phase, resulting in a smaller free energy to drive crystallization [43]. Zhang et al. [44] found that the addition of 1 at% Sn in Cu60Zr30Ti10 alloy could inhibit the trend of crystallization due to the large difference in atomic size of Sn, Cu, Zr and Ti. Similarly, in the present study, the glass forming ability of the coating was improved when Ni was gradually substituted by Sn and Ti. Nevertheless, the process of laser cladding technique is a cyclic heating and cooling process, and the temperature field of the current sedimentary layer has annealing effect on the deposited region, leading to the crystallization of the deposited MG [21,22]. Hence, the formation of Ti2CuNi and TiCu phases in the MGCCs is ascribed to the annealing effect during laser cladding.
Table 1 Chemical compositions (at.%) of white dot on the debris in Fig. 4. Spots
Ti
Zr
Cu
Ni
Sn
O
Sa Sb Sc
22.29 23.00 26.83
4.34 3.40 4.48
20.06 21.76 20.09
5.11 3.67 –
– 1.54 2.27
48.21 46.63 46.32
4.2. Tribological behaviors The wear loss of all the coatings are an order of magnitude less than pure Ti in wet sliding process. The main reason is that the wear mechanisms of pure Ti and MGCCs are different. The underlying wear mechanism of pure Ti is adhesive wear and corrosion wear, while that of the MGCCs is abrasive wear and corrosion wear. In addition, owing to unique structures of MGCCs, many researchers have proved that MGCCs show better wear resistance than that of the traditional crystalline materials [31,45]. The mean COF values and wear rates for three Ti-based MGCCs are as the following order: Ti51Zr5Cu41Sn3 ≥ Ti45Zr5Cu41Ni9 coating > Ti45Zr5Cu41Ni6Sn3 coating, as shown in Fig. 3(a, b). There is no significant difference between Ti51Zr5Cu41Sn3 and Ti45Zr5Cu41Ni9
Fig. 5. Potentiodynamic polarization curves (Tafel curve) of composite coatings and pure Ti immersed in SBF.
Ti51Zr5Cu41Sn3 > Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9. The size of dendrites in the MGCCs was inversely proportional to amorphous content. There are two possible reasons. Firstly, the radius of the Sn atom is
Fig. 6. The concentrations of Cu ions (a) and Ni ions (b) released form coatings immersed in PBS. 7
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Fig. 7. SEM micrographs of difference coating after immersed in SBF for 7 days: (a) Ti45Zr5Cu41Ni9 coating, (b) Ti45Zr5Cu41Ni6Sn3 coating, and (c) Ti51Zr5Cu41Sn3 coating, corresponding to the EDS pattern (a1), (b1) and (c1), ,respectively. XPS spectra of difference coating after immersed in SBF for 7 days: (d) General scanning, (e) Ca peaks, (f) P peaks.
coatings. On the one hand, the microhardness of specimens is in the order: Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9 coating > Ti51Zr5Cu41Sn3coating > Ti. Generally, the wear resistance of the material is proportional to the hardness. Qiao et al. [31] also pointed out that the wear resistance of amorphous alloy was much worse than that of MGCCs. On the other hand, the Ti45Zr5Cu41Ni9 coating contains high content and large size Ti2CuNi dendrites (Fig. 1(c)), which lead to the higher electrochemical corrosion in SBF solution (Fig. 5), causing accelerated corrosion of wear, and resultant increased wear rate. In addition, as shown in Fig. 4, oxides tend be formed in the friction process in SBF solution owing to the high chemical activity of MGCCs. Oxide has higher hardness than metallic matrix, which causes better wear resistance and lower friction coefficient [13,46]. Espallargas et al. [47] suggested that Zr-Cu-Ni-Al BMG accelerated corrosion and increases wear rate due to electrical coupling effects in SBF. Wang et al. [48]studied the friction and wear behavior of Zr- based BMG in simulated physiological media, and suggested that wear deterioration was a typical tribological corrosion, which was mainly determined by the synergistic interaction between abrasive and corrosion wear. Therefore, the antagonism between fine grain strengthening and corrosion performance would be also responsible for the no significant difference in COF and wear rate between Ti51Zr5Cu41Sn3 coating and Ti45Zr5Cu41Ni9 coating.
4.3. Biological behaviors Previous studies have shown that calcium phosphate is responsible for bone formation [49,50]. Thus, the three MGCCs are promising candidates for orthopaedic implants and dental materials. However, SaOS-2 cells exhibit different proliferation and biological activity on different coatings compare to pure Ti (Figs. 8 and 9). The Ti51Zr5Cu41Sn3 coating showed excellent biocompatibility, but Ti45Zr5Cu41Ni9 coating and Ti45Zr5Cu41Ni6Sn3 coating had obvious cytotoxicity to SaOS-2 cells. The reason could be that the amorphous content is relatively high in the Ti51Zr5Cu41Sn3 coating, which has good corrosion resistance (Fig. 5), and possess lower release rate of the concentration of Cu ions (Fig. 6). Recently, Huang et al. [51] suggested that the Cu-doping TiO2 coatings facilitated cell proliferation and promote macrophage-mediated osteogenesis. Due to the high content of Cu and Ni in the Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings, Cu ions and Ni ions released into the medium would be significantly toxic to cells. Schedle et al. [52] reported that toxicity level of Cu ions to human mast cell line HMC-1 was about 6 mg/L. Wataha et al. [15] reported that 25% toxicity of Ni ion to L929 cell line was about 0.22 mmol/L, which was about 12.98 mg/L. As shown in Fig. 6, the concentration of Cu and Ni ions released from day 1 to 14 was significantly lower than the previously reported safe ion concentration [51,52]. It might be due to the fact that the Cu ions and Ni ions concentrations on the surface of the material are significantly higher 8
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Fig. 8. (a) The fluorescence microscope of the live cells stained with green. (b) The relative number of live cells. (c) Cell proliferation was measured by CCK-8; * P < 0.05 compared to pure Ti. ** P < 0.01 compared to pure Ti. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
than the concentrations in the medium [53], which are toxic to the cells. However, the ion concentration on the surface of the material has not been accurately detected so far. Although the Ti41.5Zr2.5Hf5Cu37.5Ni7.5Si1Sn5 BMG implied low cell viability in vitro, it showed good biocompatibility in vivo evaluation for 30 days implantation in beagle's mandible due to dynamic environment in vivo [13]. The surface of Ti51Zr5Cu41Sn3 coating showed excellent biocompatibility, while the surfaces of Ti45Zr5Cu41Ni9 coating and Ti45Zr5Cu41Ni6Sn3 coating had obvious cytotoxicity for SaOS-2 cells in vitro due to the high Cu and Ni ions concentrations in surface. Therefore, it is necessary to conduct long-term implantation to further study the biological properties.
successfully fabricated on the pure Ti by laser cladding. (1) The Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings are mainly consisted of amorphous phases and Ti2CuNi reinforced phases, while Ti51Zr5Cu41Sn3 coating mainly is composed of amorphous phases and TiCu reinforced phases. The order of amorphous phase percentage in the three MGCCs is Ti51Zr5Cu41Sn3 > Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9. (2) Compared to pure Ti, MGCC Ti51Zr5Cu41Sn3 exhibited excellent tribological properties in SBF solution. The wear resistance of MGCCs are in the order: Ti45Zr5Cu41Ni6Sn3 coating > Ti45Zr5Cu41Ni9/Ti51Zr5Cu41Sn3 coating > pure Ti. (3) The pure Ti and Ti51Zr5Cu41Sn3 coating present no cytotoxicity, while Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings have obvious cytotoxicity in vitro. XPS shows that three kinds of coatings were able to induce hydroxyapatite depositions resulting in bone formation.
5. Conclusions Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 Ti51Zr5Cu41Sn3 MGCCs were 9
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Fig. 9. The morphology of adherent cells on difference sample surfaces after 1, 3 and 5 day for different coatings.
Appendix A. Supplementary material
(4) Although Ti51Zr5Cu41Sn3 coating displayed good biocompatibility in vitro, high levels of Cu ions caused a low rate of cell growth. Therefore, it is necessary to conduct long-term implantation to further study the biological properties. Hopefully, the excellent properties of Ni-free MGCCs make them promising candidates for orthopaedic implants and dental applications.
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CRediT authorship contribution statement Hong Wu: Conceptualization, Funding acquisition and Writing review & editing. Luxin Liang: Data curation and Writing - original draft. Xiaodong Lan: Methodology. Yong Yin: Methodology. Min Song: Data curation. Ruidi Li: Data curation. Yong Liu: Formal analysis. Haiou Yang: Formal analysis. Ling Liu: Writing - review & editing. Anhui Cai: Writing - review & editing. Qingxiang Li: Validation. Weidong Huang: Validation. Declaration of Competing Interest The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. Acknowledgment This work was supported by the National Key Research and Development Program (Grant No. 2016YFB1100103), National Natural Science Foundation of China (Grant No. 51771233), Key Research and Development Program of Hunan Province (Grant No. 2016JC2003), China Postdoctoral Science Foundation (Grant No. 2018M633164), Fundamental Research Funds for the Central Universities of Central South University (Grant No. 2018ZZTS127). 10
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Full Length Article
Tribological and biological behaviors of laser cladded Ti-based metallic glass composite coatings
T
Hong Wua,b,c, , Luxin Lianga, Xiaodong Lana, Yong Yina, Min Songa, Ruidi Lia, Yong Liua, ⁎⁎ Haiou Yangc, Ling Liud, , Anhui Caie, Qingxiang Lib, Weidong Huangc ⁎
a
State Key Laboratory of Powder Metallurgy, Central South University, Changsha 410083, PR China Shenzhen Zhong Jin Ling Nan Nonfemet Co., Ltd, Shenzhen 518040, PR China c School of Materials Science and Engineering, Northwestern Polytechnical University, Xi’an 710072, PR China d Hepatobiliary and Pancreatic Surgery Department, Xiangya Hospital, Central South University, Changsha 410008, PR China e College of Mechanical Engineering, Hunan Institute of Science and Technology, Yueyang 414000, PR China b
ARTICLE INFO
ABSTRACT
Keywords: Laser cladding Metallic glass composite coating Tribological behavior Biological behavior
In order to improve the wear resistance of pure Ti, three kinds of metallic glass composite coatings (MGCCs) with the compositions of Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3, were deposited on pure Ti by laser cladding. It is found that the order of amorphous phase percentage is Ti51Zr5Cu41Sn3 > Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9. Compared to the pure Ti substrate, the wear resistances of three coatings were significantly improved in simulated body fluid (SBF) solution. The wear resistance of MGCCs in SBF solution is in the order of Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9 ≥ Ti51Zr5Cu41Sn3, which is attributed to the combination of fine grain strengthening and well corrosion resistance. XPS results imply that these coatings enable to induce the deposition of calcium phosphate which is excellent responsible for the bone formation. The CCK-8 test and cell adhesion experiments show that compared to pure Ti, the Ti51Zr5Cu41Sn3 coating was free of cytotoxicity, while Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings have obvious cytotoxicity in vitro due to the high release rate of Cu and Ni ions from coatings. It is suggested that the Ti51Zr5Cu41Sn3 coating had not only well wear resistance, but also excellent biocompatibility.
1. Introduction Biomedical metals and alloys have been widely used in hard tissue (such as artificial bone, arthrosis and teeth) repair and implants [1]. Titanium (Ti) and its alloy have been extensively applied to orthopaedic implants for the good biocompatibility, together with excellent mechanical properties [2,55]. However, their service life is significantly limited due to the poor wear resistance. Loosening of the prosthesis due to sliding wear causes 80% of joint prostheses to require secondary surgery [1]. In addition, large number of wear debris and metallic ions such as Al and V for Ti-6Al-4V alloy are released into body tissue, which tend to provoke adverse reactions, like allergy and toxicity [3,4]. To improve wear resistance of Ti and its alloys, a variety of surface modification technologies were exploded nowadays, such as physical deposition methods by plasma spraying [5] and physical vapor deposition [6], chemical surface treatments like nitriding [7] and carburization [8]. However, the coatings prepared by plasma spraying and
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physical vapor deposition are easily exfoliated in case of frequent friction, because they are only mechanically bonded to the metallic substrate with low bonding strength [9]. Although nitriding and carburization can improve the bonding strength of the coating, both their complex operating processes and the easy deformation of the workpiece limit their widespread applications [9]. It is well-known that the atoms of bulk metallic glasses (BMGs) are disorderly arranged, which exhibits excellent performance on corrosion resistance, mechanical strength and wear resistance. These inspired properties make BMGs to possess great potential to be used as biomaterials [10-12]. Among the BMGs systems, Ti-based MG alloys have become a preferred candidate for structural materials or biomaterials because of their good wear resistance and excellent biocompatibility [12]. The BMGs exhibit much lower coefficient of friction (COF) and better anti-wear resistance than pure Ti [13]. More importantly, the results of in vivo implantation for 3 months present that the Ti41.5 Zr2.5 Hf5 Cu37.5 Ni7.5Si1 Sn5 BMGs are firmly bonded to the alveolar bone
Corresponding author at: State Key Laboratory of Powder Metallurgy, Central South University, Changsha 410083, PR China. Corresponding author. E-mail addresses: [email protected] (H. Wu), [email protected] (L. Liu).
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https://doi.org/10.1016/j.apsusc.2019.145104 Received 25 July 2019; Received in revised form 28 November 2019; Accepted 16 December 2019 Available online 19 December 2019 0169-4332/ © 2019 Elsevier B.V. All rights reserved.
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tissue [13,14]. Nickel ions are biologically toxic, that would cause series of harmful physiological and allergic reactions [15]. Therefore, Ni-free Ti-based amorphous alloys have been developed to meet the needs of biomedical applications [16,17]. The wear resistance as well as corrosion resistance of the Ti47Cu38Zr7.5Fe2.5Sn2Si1Ag2 BMG are significantly superior to that of the Ti6Al4V alloy either in air and in phosphate buffer saline (PBS) solutions [16]. However, the size of Tibased amorphous alloy fabricated by copper mold casting cannot meet the requirement of biological material because of processing technology. In recent years, laser cladding technique has been widely used in the preparation of BMGCs due to the advantages of rapid heating and cooling [18-20]. As laser cladding technique involves a repetitive cycle of heating and cooling, the temperature field of the current sedimentary layer will have annealing effect on the already deposited region, leading to the structural relaxation as well as crystallization of the deposited MG [21,54]. Zhang et al. [19]fabricated five layers Zr-based BMGCs by laser solid forming, in which the percentage of amorphous is still 81.5 vol%. Pauly et al. [20] also prepared large sized Fe-based MG with lower crystallization by selective laser melting. To overcome size limitations of amorphous alloy, Lan et al. [22] developed a process for depositing metallic glass composite coatings (MGCCs) on Ti alloys by laser cladding technique. Meanwhile, the metallurgical bonding strength between the clad layer and the metallic substrate is strong [19,22]. The microhardness and wear resistance of the coating can be significantly improved owing to the precipitation of the crystallite from the amorphous alloy [22,23]. In addition, it is reported that stannum (Sn) element is effectively improve the glass forming ability [24], plastic strain [25], corrosion resistance and biocompatibility of metal materials [26]. Therefore, it is promising to prepare Ni-free Ti-based MGCCs with Sn on pure titanium by laser cladding. The wear resistance is considered to be very important for laser clad MGCCs used for biomedical metals. It is well accepted that, the friction and wear behaviors are not the intrinsic properties of material, but strongly dependent on the wear conditions and testing parameters, such as sliding speeds, applied loads, sliding distances, and lubrication conditions, etc. [27,28]. In our previous study, it was found that the wear rate of the Zrbased BMG increased significantly with the increase of oxygen content in testing environment [27,28]. However, reported literatures on both tribological behaviors in SBF and biological properties of laser cladding Ti-based MGCCs are still inadequate. In the present study, three kinds of Ti-based MGCCs (Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3) were deposited on the surface of pure Ti by laser cladding. In order to evaluate whether they could meet the requirements of orthopedic implants, the tribological behaviors in SBF, corrosion resistance and biological properties of three kinds of MGCCs were systematically investigated.
1 mm, single beam pulse energy 143 J, pulse duration 6 ms, pulse frequency 20 Hz, scanning speed 5 mm/s, and overlap ratio 50%. The thickness of each layer was about 0.2 mm. 2.2. Coating characterization The coatings were cut into test samples by electrical discharge machine. The microstructure of the Ti-based MGCCs was observed by scanning electron microscopy (SEM, FEI Quanta FEG 250). The phases of laser cladded Ti-based MGCCs were analyzed by X ray diffractometer (XRD; Rigaku D/MAX-2250) with a Cu Kα radiation and scanning step of 0.02°. The thermal stability of the deposits was investigated by differential scanning calorimeter. The temperature of the DSC test ranges from room temperature to 1000 K. The experiments were carried out under argon atmosphere with a heating rate of 0.67 K/s. The distributions of relevant elements were analyzed by electron probe microanalyzer (EPMA; JXA-8230, Japan). The transmission electron microscope specimens were prepared using the focused ion beam (FIB). The specimens were characterized by transmission electron microscopy (TEM; Tecnai G2 F20 S-TWIN, USA). 2.3. Vickers microhardness The Vickers microhardness of various specimens was measured by the Buehler Micromet 5104 microhardness tester with a load of 100 g for 30 s. All experiments were repeated three times. 2.4. Tribological behaviors in SBF The tribological properties of pure Ti and Ti-based MGCCs were measured by tribotest device (UMT-3, USA) in SBF solution (CZ0400, China) at room temperature. The ion concentrations in SBF solution are showed in Table S1 in supplement information. Si3N4 balls with diameter of 6 mm were selected as counterpart. Loading conditions of 20 N was applied, with the linear speed being 50 mm/s and wear time being 30 min. The wear rate ω is determined by the following equation:
V = Mloss/P = V /(W L)
(1) (2)
where V is the wear volume, M is mass loss after the test (g), P is the density of the material (g/cm3), W is the load (N) and L is the total slip distance (m). 2.5. Electrochemical test The dynamic potential polarization curves were detected on Princeton electrochemical test system (Princeton Applied Research 4000). The pure Ti and Ti-based MGCCs were cut into square pieces of 10 mm × 10 mm × 1 mm and gradually ground by the metallographic grinding paper, and then they were cleaned in deionized water for ten minutes. After samples were dried at room temperature, electrochemical experiments were conducted. The working electrode was square sheets of pure Ti or Ti-based MGCCs. The alloy sheets were welded with copper wire and enclosed by the epoxy resin to make electrodes, with working area of 1 cm2. The saturated calomel electrode was used as reference electrode, and platinum sheet was auxiliary electrode. The voltage range was 0.25 V each around the initial voltage and the scan rate was 0.003 V/s. The dynamic potential polarization curves were tested in SBF solution at room temperature.
2. Materials and methods 2.1. Coating fabrication The circular bulks used as substrates was pure Ti with a size of Ф20 mm × 5 mm. The nominal components of gas-atomized powders for laser cladding were Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 (at%), respectively. The size of powder was less than 75 μm. Laser cladding was conducted in glove boxes under argon atmosphere (oxygen content less than 10 ppm) in order to prevent any oxidation reaction. Before laser cladding, the surface of pure Ti samples was polished with 800 mesh sand papers. During the cladding, a powder layer with the thickness of 0.2 mm was firstly laid on the pure Ti (Ф 20 mm), then the powder layer was melted point by pulsed laser beam and solidified to form a solid deposited layer. The specific experimental procedures are consistent with the literature [22]. Briefly, the 300 W Nd: YAG pulsed laser was used for multiple clad. Laser processing parameters were wavelength 1.06 μm, laser beam diameter
2.6. Analysis of MGCCs after immersed in SBF The surface chemical states of the specimens immersed in SBF solution for 7 days were evaluated using SEM, EDS and the X-ray photoelectron spectroscopy (XPS; K-ALPHA, Thermo Fisher Scientific, 2
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Britain). In order to further understand the form of elements in the coatings, Ca and P peaks of the coating surfaces were emphatically analyzed through XPS. The test results were analyzed and processed by XPSPEAK4.1 software.
Fig. 1(a). Three coatings displayed broad diffraction maxima at 2θ = 41°, corresponding to amorphous phase and several sharp diffraction peaks are related to crystalline phases. The crystalline phases contain Ti2CuNi phase for Ti45Zr5Cu41Ni9 coating and Ti45Zr5Cu41Ni6Sn3 coating, and TiCu phase for Ti51Zr5Cu41Sn3 coating. With the addition of Sn element, the peak of amorphous phases of Ti51Zr5Cu41Sn3 coating were broader than that of Ti45Zr5Cu41Ni6Sn3 coating and Ti45Zr5Cu41Ni9 coating. The crystallinities of three coatings calculated by the RulandVonk method are in the order of Ti45Zr5Cu41Ni9 (51.45%) > Ti45Zr5Cu41Ni6Sn3 (39.48%) > Ti51Zr5Cu41Sn3 (29.21%). These results are consistent with our previous results [22]. No oxide was detected on any of the coating surfaces, indicating that the oxygen content of the coating is low. In order to further determine the existence of an amorphous phase in coating Ti51Zr5Cu41Sn3, the DSC traces revealing the glass transition temperature Tg and the onset temperature of crystallization Tx was provided, as indicated in Fig. 1(b). The deposited coating of Ti51Zr5Cu41Sn3 exhibited ΔT = 30 K with Tg = 673 K and Tx = 703 K. To further analyze the microstructure, the surface and cross sections of different coatings were observed by SEM in Fig. 1(c) and supplementary information Fig. S1. The structures of coating surfaces were homogeneous with some flower-like dendrites. Obviously, the dendrite size in the MGCCs was in the order of Ti45Zr5Cu41Ni9 > Ti45Zr5Cu41Ni6Sn3 > Ti51Zr5Cu41Sn3, which is consistent with crystallinity. The elemental distributions of Ti, Zr, Cu, Ni and Sn are exhibited in Fig. 1(c). The distribution of elements in the amorphous zone is relatively homogeneous, without obvious segregation of components exists for various coatings. It is necessary to point out that the content of Ti, Ni and Sn element in the dendrite are higher than those in the amorphous zone for Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3. The concentration of Ti element in the plate-shaped dendrites is higher than amorphous for Ti51Zr5Cu41Sn3. In our previous work, high resolution transmission electron microscopy and selected area electron diffraction were employed to characterize the Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3, indicating that the Ti2CuSn nanocrystals are dispersed in the Ti2CuNi dendrites in Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 [22]. In order to investigate the dendrite phase in Ti51Zr5Cu41Sn3 coating, the microstructure of the specimen was characterized using TEM, as shown in Fig. 2. The amorphous phase S1 in Fig. 2(a) was further confirmed by the typical halo ring patterns in Fig. 2(b). Simultaneously, some nanocrystal phases embedded in the amorphous matrix are identified as TiCu phases [30]. The dendrite phase S2 (Fig. 2(a)) is identified as TiCu phase using selected area electron diffraction pattern in Fig. 2(c). The TEM results are consistent with the XRD patterns. Additionally, as displayed in Fig. S1, the microstructure of the coating was epitaxial to the substrate by columnar grains, implying a good fusion bond between the substrate and coating.
2.7. Metal ions release To evaluate the releasement of Cu and Ni ions, the specimens of coatings were cut into the size of 10 mm × 10 mm × 1 mm by wire electrical discharge machining, and polished with 2000 mesh sand papers. Then the coatings were immersed in 10 ml phosphate buffered saline (PBS) for various days (1, 3, 7 and 14 days) in constant temperature incubator with 37 °C. The solutions containing Cu and Ni ions were detected by the inductively coupled plasma atomic emission spectrometry (ICP-AES). 2.8. SaOS-2 cell culture and proliferation Human osteoblastic cells (SaOS-2), which were acquired from China Infrastructure of Cell Line Sources, were cultured in McCoy’s medium containing 15% fetal bovine serum and 1% penicillin/streptomycin solution. Cells were cultured in incubator with 5% CO2 and 100% humidity at 37 °C. To assess cell proliferation behavior, the SaOS-2 cell suspension with a concentration of 30,000 cells per well were introduced on the coating surfaces for 1, 3 and 5 days. Before the cells were seeded in 24well plates, all disks with the size of 10 mm × 10 mm × 1 mm were polished with 800 mesh sand papers, and were sterilized using ultraviolet rays for 48 h. Fresh medium and CCK-8 solution with a volume ratio of 10:1 was added into each well. After incubated for 1 h at 37 °C, 100 μL of the incubation solution was added to a 96-well plate, and the absorbance was detected by a microplate reader at 450 nm. 2.9. Live cell staining The live cell staining was executed according to manufacturer’s instructions. Briefly, SaOS-2 cells were seeded on the surface of coatings with a concentration of 30,000 cells per well, and cultured for 1, 3 and 5 days. Cells were washed with PBS 3 times. 2 μM calcein-AM was added in each well, and cells were cultured at cell incubator for 15 min. Fluorescence microscopy (DFC420C, Leica, Germany) was used to observe and photograph. 2.10. Cell morphology For the observation of morphology, SaOS-2 cell suspension with a concentration of 30,000 cells per well were cultured on the various coating surfaces and cultured for 1 day. To prepare the samples for SEM, the specific preparation method was consistent with the literature [29]. Briefly, the cells were first fixed with 2.5% glutaraldehyde, and dehydrated with different concentration of alcohol gradient. Then ethanol was replaced with different concentration of tert-butanol, and finally the samples were freeze-dried.
3.2. Microhardness The microhardness for the surface of specimens is shown in Fig. S2. The Vickers hardness for Ti is only HV100 156.9, while the microhardness for Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 coatings are HV100 590.1, 663.6 and 515.5, respectively. The microhardness of Ti is much lower than that of MGCCs due to the disorder arrangement of atoms and strengthening of nanocrystals [31]. Previous studies indicated that the hardness of Ti2CuNi phase is higher than that of amorphous alloy, and the hardness of TiCu is lower than that of amorphous alloy [32]. Therefore, the hardness of coating Ti45Zr5Cu41Ni9 is higher than that of Ti51Zr5Cu41Sn3 coating. We previously proved that there are fine Ti2CuSn nanocrystals dispersed in the Ti2CuNi phase in the Ti45Zr5Cu41Ni6Sn3 coating. The nanocrystals have a strengthening effect on dendritic Ti2CuNi, so the hardness of Ti45Zr5Cu41Ni6Sn3 coating is higher than that of Ti45Zr5Cu41Ni9 coating [22].
2.11. Statistical analysis All experiments were repeated three times and the data were expressed as mean ± standard deviation. To compare the differences in significance between the study groups, the two samples t-test was applied. The value of p < 0.05 was defined as statistically significant. 3. Results 3.1. Surface microstructure analysis The XRD patterns of the deposited coatings surfaces are presented in 3
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Fig. 1. (a) XRD patterns of coatings surfaces. (b) DSC curve of the coating Ti51Zr5Cu41Sn3. (c) SEM BSE and EPMA micrographs showing the surface microstructures of the coatings.
0.2288 × 10−3 mm3 N−1 m−1, and 0.5101 × 10−3 mm3 N−1 m−1, respectively. The three coatings exhibited lower COF than control group (pure Ti) in the SBF solution. The wear loss of three coatings were an order of magnitude less than the control group. The SEM morphologies of surfaces after wear in SBF solution are presented in Fig. 3(c–f). The worn surface of the pure Ti are consisted of numerous deep grooves and wear debris in the worn scar, revealing the local severe deformation. Compared to the pure Ti, the worn surfaces of the MGCCs in Fig. 3(d–f) display much shallower
3.3. Tribological behaviors in SBF solution As shown in Fig. 3(a, b), the friction coefficients of coatings Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 were stable under the wear time of 1800 s in SBF solution. However, under wet sliding condition in SBF solution, the wear rates of pure Ti, Ti45Zr5Cu41Ni9 coating, Ti45Zr5Cu41Ni6Sn3 coating and Ti51Zr5Cu41Sn3 coating were 5.9877 × 10−3 mm3 N−1 m−1, 0.4667 × 10−3 mm3 N−1 m−1, 4
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Fig. 2. The TEM macrographs (a) of Ti51Zr5Cu41Sn3, together with SAED patterns obtained from amorphous matrix S1 (b) and dendrite S2 (c).
addition, Cu2+ ions released from Ti51Zr5Cu41Sn3 coatings and Ti45Zr5Cu41Ni6Sn3 coatings are basically equal, while Cu2+ and Ni2+ ions released from Ti45Zr5Cu41Ni9 coating are significantly higher that from the coating Ti45Zr5Cu41Ni6Sn3 after immersion for 14 days.
grooves. For the Ti45Zr5Cu41Ni6Sn3 coating, the plastic deformation grooves were the shallowest. The surface of Ti45Zr5Cu41Ni9 coating was corroded after wear in SBF. The wear resistance of MGCCs is in the order of Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9/Ti51Zr5Cu41Sn3 > pure Ti. The SEM images of the wear debris are presented in Fig. 4. The EDS scanner was employed to analyze the elements of the wear debris (marked as Sa, Sb and Sc), which is displayed in Table 1. The results showed that all the wear debris contained oxygen element, indicating the formation of complex metallic oxides during the wet frictional process. Interestingly, oxygen content of wear debris was the highest in Sa of Ti45Zr5Cu41Ni9 coating, and the minimum oxygen was found in Sb of Ti51Zr5Cu41Sn3 coating. The debris exhibited a flat and slip surface without plastic deformation. It may be attributed to the deterioration of the combination between the oxides layer and the metal matrix in PBS solution, which promotes the refinement of wear debris [17]. In general, the tribological behavior of MGCCs in SBF solution may be influenced by the precipitation of the crystalline phase from the amorphous matrix [33,34], and solution corrosion [17]. In a word, the main wear mechanisms of pure titanium were adhesive wear and corrosion wear, while the MGCCs were dominated by abrasive wear and corrosion wear.
3.5. Analysis of coatings after immersed in SBF The surface morphology and chemical composition of the coatings after immersed in SBF for 7 d are shown in Fig. 7. Many small particles can be found on the surface of various coatings (Fig. 7(a–c)). EDS analyses show that the chemical composition of the precipitates contain Ca and P. More detailed analyses on the surface composition were conducted by XPS (Fig. 7(d–f)). As presented in Fig. 7(d), XPS spectra of the surface indicates that there are a few elements such as Ca, P, Ti and Cu in the coating surface. In order to further clarify the form of elements in the coatings, Ca peaks, P peaks of the coating surfaces were emphatically analyzed through XPS. Fig. 7(e) presents Ca2p (2p3/2 and 2p1/2) peaks in the surface, respectively located at 347.0 eV ± 0.2 and 350.6 eV ± 0.2, and Fig. 7(f) shows the P2p peaks in the surface located at 133.0 eV ± 0.2, indicating the presence of calcium phosphates [37,38]. These results agree well with the findings of previous study that Cu promote the calcium phosphate deposition [39].
3.4. Corrosion behavior and ion release of coating surfaces
3.6. Cell proliferation and adhesion
Fig. 5 displays the typical polarization curves of different coatings in SBF solution. The corrosion potentials (Ecorr) were in the order of coating > Ti45Zr5Cu41Ni6Sn3 coating > Ti51Zr5Cu41Sn3 Ti45Zr5Cu41Ni9 coating > pure Ti. It could be clearly seen that the corrosion current density (Icorr) of MGCCs samples was in the order of 10−5 A/cm2 in SBF solution, which is approximately equal to that of pure Ti, indicating that the corrosion resistance of three coatings is stronger than that of pure Ti. The Ti45Zr5Cu41Ni9 coating had the worst corrosion resistance of the three coatings. MGCCs are liable to corrosion in the complex humoral environment, resulting in the release of Ni2+ ions and Cu2+ ions. Researchers shown that the release of large amounts of Ni2+ ions and Cu2+ ions can cause cytotoxicity, tissue and organ disorders [35,36]. The amounts of Ni2+ ions and Cu2+ ions leached from the three coatings were tested by ICPAES, as shown in Fig. 6. The concentration of cumulative Ni2+ ions and Cu2+ ions increased with the extension of immersion duration. In
Proliferation of SaoS-2 cultured on the surface of coatings and pure Ti was assessed using Calcein-AM staining and CCK-8 reagent. As displayed in Fig. 8(a–b), the cell number on the surface of Ti45Zr5Cu41Ni6Sn3 coating and Ti45Zr5Cu41Ni9 coating was significantly less than that on the pure Ti for day 1, day 3 and day 5, while the cells number on surface of Ti51Zr5Cu41Sn3 coating and pure Ti was comparable. The cells number on various surfaces of coating increased from day 1 to day 5. Fig. 8(c) presents that the quantitative results detected by CCK-8 reagent is in accordance with the fluorescence images described in Fig. 8(a). According to the current ISO standards, cell viability higher than 75% could be considered with no toxic risks for medical devices, implying that the Ti51Zr5Cu41Sn3 and Ti45Zr5Cu41Ni6Sn3 coatings were almost non-toxic for SaOS-2 cells, while Ti45Zr5Cu41Ni9 coating had obvious toxicity. The morphology of SaOS-2 cells on different surfaces of MGCCs for 5
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Fig. 3. (a) Friction coefficient curves of coatings immersed in SBF under a normal loads of 20 N. (b) Mean values of friction coefficients of three coatings. The worn surfaces of Ti (c), Ti45Zr5Cu41Ni9 coating (d), Ti45Zr5Cu41Ni6Sn3 coating (e), and Ti51Zr5Cu41Sn3 coating (f). ** P＜0.01, compared to pure Ti. # P＜0.05, compared to Ti45Zr5Cu41Ni9 coating.
day 1, 3 and 5 was further observed by SEM. As shown in Fig. 9, after 1 day, SaOS-2 cells grown on surface of various coatings exhibited similar spreading area compared to pure Ti. The lamellipodium formation and filopodium extension were obviously observed on the surfaces of deferent coating. At day 3 and 5, a lot of cells grown on surfaces of the Ti45Zr5Cu41Ni9 coating showed obvious atrophic morphology, and some of the cells grown on surfaces of the Ti45Zr5Cu41Ni6Sn3 coating showed atrophic morphology. Almost no cells showed atrophic morphology on the surface of the Ti51Zr5Cu41Sn3 coating. These observations indicate that Ti51Zr5Cu41Sn3 coating show excellent biocompatibility, but
Ti45Zr5Cu41Ni9 coating has obvious toxicity for SaOS-2 cells. 4. Discussion 4.1. Phase composition In the present study, three kinds of Ti-based MGCCs, Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 and Ti51Zr5Cu41Sn3 were deposited by laser cladding. With the addition of Sn, the amorphous phase content increases. The order of amorphous phase percentage in as-designed three MGCCs is 6
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Fig. 4. SEM images of collected MGCCs wear debris: (a) Ti45Zr5Cu41Ni9 coating, (b) Ti45Zr5Cu41Ni6Sn3 coating and (c) Ti51Zr5Cu41Sn3 coating.
relatively large, which increases the mismatch of the atomic size which may delay the rearrangement of the atoms required for crystallization, thus enhancing the glass forming ability [40,41]. Secondly, appropriate negative heat of mixing between the main elements is beneficial to the deep suppression of the free energy curve of the supercooled liquid phase [42]. As the negative mixing enthalpy between Sn and Cu, Ti and Zr are higher than that between Ni and Cu, Ti and Zr, which contribute to inhibit the free energy of supercooled liquid phase, resulting in a smaller free energy to drive crystallization [43]. Zhang et al. [44] found that the addition of 1 at% Sn in Cu60Zr30Ti10 alloy could inhibit the trend of crystallization due to the large difference in atomic size of Sn, Cu, Zr and Ti. Similarly, in the present study, the glass forming ability of the coating was improved when Ni was gradually substituted by Sn and Ti. Nevertheless, the process of laser cladding technique is a cyclic heating and cooling process, and the temperature field of the current sedimentary layer has annealing effect on the deposited region, leading to the crystallization of the deposited MG [21,22]. Hence, the formation of Ti2CuNi and TiCu phases in the MGCCs is ascribed to the annealing effect during laser cladding.
Table 1 Chemical compositions (at.%) of white dot on the debris in Fig. 4. Spots
Ti
Zr
Cu
Ni
Sn
O
Sa Sb Sc
22.29 23.00 26.83
4.34 3.40 4.48
20.06 21.76 20.09
5.11 3.67 –
– 1.54 2.27
48.21 46.63 46.32
4.2. Tribological behaviors The wear loss of all the coatings are an order of magnitude less than pure Ti in wet sliding process. The main reason is that the wear mechanisms of pure Ti and MGCCs are different. The underlying wear mechanism of pure Ti is adhesive wear and corrosion wear, while that of the MGCCs is abrasive wear and corrosion wear. In addition, owing to unique structures of MGCCs, many researchers have proved that MGCCs show better wear resistance than that of the traditional crystalline materials [31,45]. The mean COF values and wear rates for three Ti-based MGCCs are as the following order: Ti51Zr5Cu41Sn3 ≥ Ti45Zr5Cu41Ni9 coating > Ti45Zr5Cu41Ni6Sn3 coating, as shown in Fig. 3(a, b). There is no significant difference between Ti51Zr5Cu41Sn3 and Ti45Zr5Cu41Ni9
Fig. 5. Potentiodynamic polarization curves (Tafel curve) of composite coatings and pure Ti immersed in SBF.
Ti51Zr5Cu41Sn3 > Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9. The size of dendrites in the MGCCs was inversely proportional to amorphous content. There are two possible reasons. Firstly, the radius of the Sn atom is
Fig. 6. The concentrations of Cu ions (a) and Ni ions (b) released form coatings immersed in PBS. 7
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Fig. 7. SEM micrographs of difference coating after immersed in SBF for 7 days: (a) Ti45Zr5Cu41Ni9 coating, (b) Ti45Zr5Cu41Ni6Sn3 coating, and (c) Ti51Zr5Cu41Sn3 coating, corresponding to the EDS pattern (a1), (b1) and (c1), ,respectively. XPS spectra of difference coating after immersed in SBF for 7 days: (d) General scanning, (e) Ca peaks, (f) P peaks.
coatings. On the one hand, the microhardness of specimens is in the order: Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9 coating > Ti51Zr5Cu41Sn3coating > Ti. Generally, the wear resistance of the material is proportional to the hardness. Qiao et al. [31] also pointed out that the wear resistance of amorphous alloy was much worse than that of MGCCs. On the other hand, the Ti45Zr5Cu41Ni9 coating contains high content and large size Ti2CuNi dendrites (Fig. 1(c)), which lead to the higher electrochemical corrosion in SBF solution (Fig. 5), causing accelerated corrosion of wear, and resultant increased wear rate. In addition, as shown in Fig. 4, oxides tend be formed in the friction process in SBF solution owing to the high chemical activity of MGCCs. Oxide has higher hardness than metallic matrix, which causes better wear resistance and lower friction coefficient [13,46]. Espallargas et al. [47] suggested that Zr-Cu-Ni-Al BMG accelerated corrosion and increases wear rate due to electrical coupling effects in SBF. Wang et al. [48]studied the friction and wear behavior of Zr- based BMG in simulated physiological media, and suggested that wear deterioration was a typical tribological corrosion, which was mainly determined by the synergistic interaction between abrasive and corrosion wear. Therefore, the antagonism between fine grain strengthening and corrosion performance would be also responsible for the no significant difference in COF and wear rate between Ti51Zr5Cu41Sn3 coating and Ti45Zr5Cu41Ni9 coating.
4.3. Biological behaviors Previous studies have shown that calcium phosphate is responsible for bone formation [49,50]. Thus, the three MGCCs are promising candidates for orthopaedic implants and dental materials. However, SaOS-2 cells exhibit different proliferation and biological activity on different coatings compare to pure Ti (Figs. 8 and 9). The Ti51Zr5Cu41Sn3 coating showed excellent biocompatibility, but Ti45Zr5Cu41Ni9 coating and Ti45Zr5Cu41Ni6Sn3 coating had obvious cytotoxicity to SaOS-2 cells. The reason could be that the amorphous content is relatively high in the Ti51Zr5Cu41Sn3 coating, which has good corrosion resistance (Fig. 5), and possess lower release rate of the concentration of Cu ions (Fig. 6). Recently, Huang et al. [51] suggested that the Cu-doping TiO2 coatings facilitated cell proliferation and promote macrophage-mediated osteogenesis. Due to the high content of Cu and Ni in the Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings, Cu ions and Ni ions released into the medium would be significantly toxic to cells. Schedle et al. [52] reported that toxicity level of Cu ions to human mast cell line HMC-1 was about 6 mg/L. Wataha et al. [15] reported that 25% toxicity of Ni ion to L929 cell line was about 0.22 mmol/L, which was about 12.98 mg/L. As shown in Fig. 6, the concentration of Cu and Ni ions released from day 1 to 14 was significantly lower than the previously reported safe ion concentration [51,52]. It might be due to the fact that the Cu ions and Ni ions concentrations on the surface of the material are significantly higher 8
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Fig. 8. (a) The fluorescence microscope of the live cells stained with green. (b) The relative number of live cells. (c) Cell proliferation was measured by CCK-8; * P < 0.05 compared to pure Ti. ** P < 0.01 compared to pure Ti. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
than the concentrations in the medium [53], which are toxic to the cells. However, the ion concentration on the surface of the material has not been accurately detected so far. Although the Ti41.5Zr2.5Hf5Cu37.5Ni7.5Si1Sn5 BMG implied low cell viability in vitro, it showed good biocompatibility in vivo evaluation for 30 days implantation in beagle's mandible due to dynamic environment in vivo [13]. The surface of Ti51Zr5Cu41Sn3 coating showed excellent biocompatibility, while the surfaces of Ti45Zr5Cu41Ni9 coating and Ti45Zr5Cu41Ni6Sn3 coating had obvious cytotoxicity for SaOS-2 cells in vitro due to the high Cu and Ni ions concentrations in surface. Therefore, it is necessary to conduct long-term implantation to further study the biological properties.
successfully fabricated on the pure Ti by laser cladding. (1) The Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings are mainly consisted of amorphous phases and Ti2CuNi reinforced phases, while Ti51Zr5Cu41Sn3 coating mainly is composed of amorphous phases and TiCu reinforced phases. The order of amorphous phase percentage in the three MGCCs is Ti51Zr5Cu41Sn3 > Ti45Zr5Cu41Ni6Sn3 > Ti45Zr5Cu41Ni9. (2) Compared to pure Ti, MGCC Ti51Zr5Cu41Sn3 exhibited excellent tribological properties in SBF solution. The wear resistance of MGCCs are in the order: Ti45Zr5Cu41Ni6Sn3 coating > Ti45Zr5Cu41Ni9/Ti51Zr5Cu41Sn3 coating > pure Ti. (3) The pure Ti and Ti51Zr5Cu41Sn3 coating present no cytotoxicity, while Ti45Zr5Cu41Ni9 and Ti45Zr5Cu41Ni6Sn3 coatings have obvious cytotoxicity in vitro. XPS shows that three kinds of coatings were able to induce hydroxyapatite depositions resulting in bone formation.
5. Conclusions Ti45Zr5Cu41Ni9, Ti45Zr5Cu41Ni6Sn3 Ti51Zr5Cu41Sn3 MGCCs were 9
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Fig. 9. The morphology of adherent cells on difference sample surfaces after 1, 3 and 5 day for different coatings.
Appendix A. Supplementary material
(4) Although Ti51Zr5Cu41Sn3 coating displayed good biocompatibility in vitro, high levels of Cu ions caused a low rate of cell growth. Therefore, it is necessary to conduct long-term implantation to further study the biological properties. Hopefully, the excellent properties of Ni-free MGCCs make them promising candidates for orthopaedic implants and dental applications.
Supplementary data to this article can be found online at https:// doi.org/10.1016/j.apsusc.2019.145104. References [1] L. Kunčická, R. Kocich, T.C. Lowe, Advances in metals and alloys for joint replacement, Prog. Mater Sci. 88 (2017) 232–280. [2] M. Geetha, A.K. Singh, R. Asokamani, A.K. Gogia, Ti based biomaterials, the ultimate choice for orthopaedic implants – a review, Prog. Mater Sci. 54 (2009) 397–425. [3] N.J. Hallab, S.A.T. Stafford, J.J.J. Tibor, Glant, Lymphocyte responses in patients with total hip arthroplasty, J. Orthop. Res. 23 (2005) 384–391. [4] T.H. Lin, Y. Tamaki, J. Pajarinen, H.A. Waters, D.K. Woo, Z. Yao, S.B. Goodman, Chronic inflammation in biomaterial-induced periprosthetic osteolysis: NF-kappaB as a therapeutic target, Acta Biomater 10 (2014) 1–10. [5] X. Zheng, M. Huang, C. Ding, Bond strength of plasma-sprayed hydroxyapatite Ti composite coatings, Biomaterials 21 (21) (2000) 841–849. [6] A.R. Rafieerad, A.R. Bushroa, B. Nasiri-Tabrizi, J. Vadivelu, W.J. Basirun, Silver/ silver oxide nanorod arrays from physical vapor deposition and subsequent anodization processes, Surf. Coat. Technol. 302 (2016) 275–283. [7] U. Sen, Kinetics of titanium nitride coatings deposited by thermo-reactive deposition technique, Vacuum 75 (2004) 339–345. [8] N. Tsuji, S. Tanaka, T. Takasugi, Evaluation of surface-modified Ti–6Al–4V alloy by combination of plasma-carburizing and deep-rolling, Mater. Sci. Eng., A 488 (2008) 139–145. [9] Y.S. Tian, C.Z. Chen, S.T. Li, Q.H. Huo, Research progress on laser surface modification of titanium alloys, Appl. Surf. Sci. 242 (2005) 177–184. [10] J.J. Oak, G.W. Hwang, Y.H. Park, H. Kimura, S.Y. Yoon, A. Inoue, Characterization of surface properties, osteoblast cell culture in vitro and processing with flowviscosity of Ni-free Ti-based bulk metallic glass for biomaterials, J. Biomech. Sci. Eng. 4 (2009) 384–391. [11] P. Meagher, E.D. O'Cearbhaill, J.H. Byrne, D.J. Browne, Bulk metallic glasses for implantable medical devices and surgical tools, Adv. Mater. 28 (2016) 5755–5762. [12] S. Pang, Y. Liu, H. Li, L. Sun, Y. Li, T. Zhang, New Ti-based Ti–Cu–Zr–Fe–Sn–Si–Ag bulk metallic glass for biomedical applications, J. Alloy. Compd. 625 (2015) 323–327. [13] Y.B. Wang, H.F. Li, Y. Cheng, Y.F. Zheng, L.Q. Ruan, In vitro and in vivo studies on Ti-based bulk metallic glass as potential dental implant material, Mater. Sci. Eng. C Mater. Biol. Appl. 33 (2013) 3489–3497. [14] Y.F. Zheng, H.F. Li, Y.B. Wang, X.N. Gu, Research progress on the preliminary study of bulk metallic glasses as biomaterials, Mater. Rep. 24 (2010) 1–5. [15] J.C. Wataha, C.T. Ylardm, Z. Sun, Effect of cell line on in vivo metal ion
CRediT authorship contribution statement Hong Wu: Conceptualization, Funding acquisition and Writing review & editing. Luxin Liang: Data curation and Writing - original draft. Xiaodong Lan: Methodology. Yong Yin: Methodology. Min Song: Data curation. Ruidi Li: Data curation. Yong Liu: Formal analysis. Haiou Yang: Formal analysis. Ling Liu: Writing - review & editing. Anhui Cai: Writing - review & editing. Qingxiang Li: Validation. Weidong Huang: Validation. Declaration of Competing Interest The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. Acknowledgment This work was supported by the National Key Research and Development Program (Grant No. 2016YFB1100103), National Natural Science Foundation of China (Grant No. 51771233), Key Research and Development Program of Hunan Province (Grant No. 2016JC2003), China Postdoctoral Science Foundation (Grant No. 2018M633164), Fundamental Research Funds for the Central Universities of Central South University (Grant No. 2018ZZTS127). 10
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