Mechanical and biological properties of graded porous tantalum coatings deposited on titanium alloy implants by vacuum plasma spraying

Surface & Coatings Technology 372 (2019) 399–409

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Mechanical and biological properties of graded porous tantalum coatings deposited on titanium alloy implants by vacuum plasma spraying

T

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Tsung-Yuan Kuoa, Wei-Han China, Chi-Sheng Chienb,c, , Yueh-Hung Hsieha a

Department of Mechanical Engineering, Southern Taiwan University of Science and Technology, Tainan 710, Taiwan, ROC Department of Orthopaedics, Chimei Foundation Hospital, Tainan 710, Taiwan, ROC c Department of Electrical Engineering, Southern Taiwan University of Science and Technology, Tainan 710, Taiwan, ROC b

A R T I C LE I N FO

A B S T R A C T

Keywords: Vacuum plasma spraying Tantalum Titanium Alkali and heat treatment Bioactivity Biocompatibility

Tantalum (Ta) has excellent mechanical properties, biocompatibility and chemical stability, and is more osteoconductive than titanium or cobalt-chromium alloys. However, it has a relatively high cost and is not easily fabricated. Accordingly, in the present study, a vacuum plasma spraying technique was used to deposit Ta powder in three successive layers on Ti6Al4V substrates. In the deposition process, the melting rate of the Ta powder was controlled by regulating the spray power and powder particle size in such a way as to form a coating with a porosity which increased from the inner layer to the outer layer. The Ta-coatings were processed by alkali treatment (AT), AT and heat treatment (AHT), and AHT with ultrasonic cleaning (AHT-UC), respectively. The average bond strength of the as-sprayed coating was approximately 54.5 ± 2.3 MPa. Moreover, the average porosity and roughness Ra of the coating surface were around 13% and 22.2 μm, respectively. The average porosity of the coating (with a thickness of 350–380 μm) increased from 0.6% in the inner layer to 7.6% in the outer layer. The corresponding hardness and elastic modulus (E) values decreased from 240 ± 20 to 167 ± 62 HV0.1 and 148 ± 5 to 123 ± 4 GPa, respectively. Bone-like apatite inducement tests performed in simulated body fluid (SBF) showed a particularly rapid growth of apatite on the AHT-UC Ta-coating due to the formation of amorphous sodium tantalate during AT. Furthermore, the results obtained from initial in vitro biocompatibility tests showed that osteoblast-like osteosarcoma MG-63 cells had a significantly better attachment and spreading on the Ta-coating and AHT-UC coating following culturing for 3 h than on the untreated Ti6Al4V sample. In other words, both coatings provided a better initial biocompatibility than the native Ti6Al4V substrate.

1. Introduction Metallic biomaterials are widely used for biological implants due to their optimal combination of mechanical properties, corrosion resistance in biological environments, and biocompatibility [1–4]. However, most common implant materials, such as Ti and its alloys and CoCr alloys, are much stiffer than human bone. The resulting mismatch between the stiffness of the implant and that of the surrounding bone may result in a stress shielding effect, which prompts bone resorption and an eventual loosening of the implant [4,5]. It has been shown that this problem can be mitigated by utilizing biomaterials with a greater porosity [6–8]. However, enhancing the bio-affinity of porous implant materials with the surrounding soft bone tissue remains an important concern. Tantalum (Ta) has excellent mechanical properties [9],

biocompatibility [10], chemical stability, and corrosion resistance [11]. Moreover, it is more osteoconductive than Ti or CoCr alloys and has better bioactivity [12–15]. As a result, it has been used for plates and suture wires in orthopedic, craniofacial, and dentistry applications [16]. However, the extremely high melting temperature (3017 °C) of Ta, together with its high affinity for oxygen, poses a significant challenge in fabricating all-tantalum implant structures using traditional processing methods [17]. As a result, Ta has traditionally found only limited use in medical practice. To address this problem, many studies have deposited Ta on the surface of common biomaterials such as Ti or CoCr alloys in order to confer the excellent properties of Ta to these materials while simultaneously reducing the processing and material cost [12,14,15,18,19]. In addition, several methods have been proposed for the direct manufacture of solid and porous Ta parts, including Laser Engineered Net Shaping (LENS) [15,17], Spark Plasma Sintering (SPS)

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Corresponding author at: Department of Orthopaedics, Chimei Foundation Hospital, Tainan 710, Taiwan, ROC. E-mail addresses: [email protected] (T.-Y. Kuo), [email protected] (W.-H. Chin), [email protected], [email protected] (C.-S. Chien), [email protected] (Y.-H. Hsieh). https://doi.org/10.1016/j.surfcoat.2019.05.003 Received 1 February 2019; Received in revised form 18 April 2019; Accepted 2 May 2019 Available online 03 May 2019 0257-8972/ © 2019 Published by Elsevier B.V.

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alkali-heat-treatment (AHT) also provides an effective means of fabricating biomedical coatings through chemical modification [30–32]. In the present study, the VPS technique is used to coat Ta powder in three layers on Ti6Al4V substrates. The melting rate of the Ta powder is controlled by adjusting the spray power and powder particle size used in each layer so as to form a coating with a graded porosity which increases gradually from the inside to the outside. The porous nature of the deposited coating mitigates the difference in the mechanical properties of the coating and substrate, respectively, and hence results in an improved biomedical performance. The biological properties of the Tacoating are enhanced by means of AT, AHT, and AHT and ultrasonic cleaning (AHT-UC), respectively. The surface characteristics, bond strength, micro-hardness, elastic modulus and biological performance of the various Ta-coatings are examined and compared with those of an untreated Ta coating and a native Ti6Al4V substrate, respectively. 2. Materials and methods Fig. 1. SEM morphology of Ta powder.

2.1. Preparation of specimens Table 1 Chemical composition (wt%) of Ti6Al4V alloy. Al

V

O

Fe

C

N

H

Ti

6.10

4.24

0.152

0.16

0.017

0.008

0.0006

Balance

The VPS coatings were produced using Ta powder with a purity of 99.95% and a particle size of 5–120 μm (Infralloy™ S7301, Inframat® Advanced Materials™ Ltd., Manchester CT, USA). The particle size followed a granulometric distribution with D5 equal to approximately 5 μm, D50 around 50 μm, and D90 about 120 μm. Fig. 1 shows a typical Scanning Electron Microscope (SEM) image of the Ta powder. The coatings were deposited on commercially available Ti6Al4V plates (50 × 50 × 3 mm3) and rods (ϕ25.4 × 25.4 mm). The chemical composition of the Ti6Al4V alloy is shown in Table 1. The Ta coatings were deposited in three layers using a Metco A3000 VPS machine (Sulzer Metco, Switzerland). Prior to deposition, the substrates were preheated at 650 °C to enhance the interface fusion effect. In addition, the plates and rods were sandblasted with Al2O3 particles to increase the surface roughness to approximately Ra = 4.4 μm. To produce the desired porosity gradient, the inner, middle and outer coating layers were deposited using spray powers of 49 kW, 35 kW and 35 kW, respectively. Moreover, the particle diameters of the Ta powders used in the three layers were chosen as 5–53 μm (100%); 5–53 μm (83.5%) and 53–120 μm 16.5%); and 5–53 μm (50%) and 53–120 μm (50%), respectively. The remaining deposition parameters were set in accordance with [24] and a process of trial-and-error, respectively (see Table 2). Following the VPS process, the coated samples were cleaned ultrasonically with pure acetone and distilled (DI) water and then surface treated using AT, AHT or AHT-UC (see Section 2.3).

[20] and Selective Laser Melting (SLM) [21–23]. However, plasma spray methods provide a more cost-effective, straightforward and reliable approach. Of the various plasma spraying techniques available, Vacuum Plasma Spraying (VPS) is particularly suitable for the deposition of porous, high quality coatings of reactive and high melting point metals such as Ta [19,24]. Furthermore, VPS allows for the rapid deposition of near-net shapes with complex geometries on various substrates and produces coatings with fewer oxides and an improved adhesion strength. Notably, however, the literature contains very few studies on the deposition of Ta by VPS [19,24]. When Ti metal, Ti alloy and Ta metal implants are introduced into human bone, the implant and bone are frequently separated by fibrous tissue [25]. Consequently, the bonding strength between them is impaired and the service life of the implant correspondingly reduced. This problem is caused by the fact that these metal implants are bio-inert. To address this problem, various surface treatment methods have been proposed. Broadly speaking, these methods can be classified into two main approaches. In the first approach, the surface of the implant is coated with a bioceramic material such as calcium phosphate (e.g. hydroxyapatite, HA) by pulsed laser deposition [26] or bioglass ceramic material by plasma spraying [27]. In the second approach, a bioactive surface is induced directly via chemical reaction to produce, for example, an oxide layer by micro-arc oxidation [28] or an oxide layer together with nanotubular oxides via electrochemical-assisted deposition [29]. Various studies have shown that alkali-treatment (AT) or

2.2. Measurement of surface roughness, porosity, hardness and elastic modulus The average surface roughness (Ra) of the coating surfaces was determined using a 3D laser scanning confocal microscope (VK-X200, KEYENCE Ltd., Japan) based on a minimum of three separate

Table 2 Variable and constant VPS parameters. Spraying parameters Variable parameters

Constant parameters

Parameter

Pass number

Particle size

Plasma arc Power kW (Ampere/Voltage)

Parameter

Value

Inner layer Middle layer Outer layer *F: fine *I: intermediate *C: coarse

5 5 2

F I C

49 (700/70) 35 (500/70) 35 (500/70)

Primary plasma gas/Ar (L/min) Secondary plasma gas/H2 (L/min) Carried gas/Ar (L/min) Spray distance (mm) Powder feed rate (g/min) Chamber pressure (mbar)

50 12 3 250 40 150

400

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Fig. 3. Micro-hardness measurements of Ti6Al4V substrate, substrate-coating interface, and graded porous Ta coating (inner, middle and outer layers).

2.3. Bond strength testing The adhesion strength of the coatings was evaluated by means of ASTM C633-79 tensile tests [33]. Briefly, Ta coatings were sprayed on the end of Ti6Al4V rod test specimens and the coated surfaces were then bonded to uncoated rods (grit-blasted) of the same material and geometry using a special adhesive-bonding glue (C21H25ClO5, 3M Exopy Adhesive 2214 Regular). Based on extensive preliminary tests, the optimal curing condition for the glue was found to be a temperature of 150 °C for 4 h; resulting in an average glue bond strength of 55.0 ± 4.7 MPa. The glued couples underwent tensile tests at a constant cross-head speed of 0.02 mm/s until failure by a universal testing machine AI-7000 LA 10 (Gotech, Taiwan). The bond strength was then evaluated as the average value obtained over five separate tests. After the tests, the failure modes of the specimens were examined by SEM observations. 2.4. Bioactive surface modification In the alkali surface treatment process (AT), the specimens were immersed in 1 M NaOH aqueous solution at 60 °C for 24 h and were then washed gently with acetone and DI water, and dried at 40 °C for 24 h in an oven with an air atmosphere. For the combined alkali and heat treatment process (AHT), AT specimens were heated to 300 °C at a rate of 5 °C min−1 in an electric furnace, maintained at this temperature for 1 h, and then cooled to room temperature in the furnace. Finally, for the AHT-UC process, AHT specimens were cleaned ultrasonically in acetone solution for 5 min and then dried in an oven at 40 °C for 24 h. 2.5. Bone-like apatite inducement tests Bone-like apatite inducement performance of the coatings was evaluated by soaking the specimens in a standard simulated body fluid (SBF) solution [34] for 3–14 days at a temperature of 37 ± 0.1 °C. Prior to immersion in the SBF, the specimens were sterilized at 120 °C for 20 min. Following the immersion tests, the surface morphology, chemical composition and phases of the coatings were identified using SEM, Energy Dispersive X-ray Spectrometry (EDS) and X-ray Diffraction (XRD).

Fig. 2. Graded porous Ta-coating: (a) surface morphology (x100); (b) 3D confocal laser microscope image; (c) cross-section microstructure.

measurement points. The porosity of the surface and each cross-sectional layer of the final coated sample was evaluated using commercial Image Pro Plus 5.0 software. The hardness was determined using a micro-Vickers hardness tester under a maximum indentation load of 100 g. Finally, the elastic modulus (E) was measured using a commercial nano-indentation system (MTS-G200) with five tests per sample.

2.6. Initial in vitro cell tests An initial assessment of the in vitro biocompatibility of the coatings was made by examining the attachment and spread of osteoblast-like osteosarcoma MG-63 cells on the specimen surface. Briefly, MG-63 cells were digested by trypsin to form cell suspensions, and the suspensions were then seeded with a cell density of 5 × 103 cells/cm2 on specimens placed in 12-well plates. The plates were cultured with Dulbecco's Modified Eagle Medium (DMEM) supplement containing 10% fetal 401

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Fig. 4. EDS line scans showing Ta, Ti, Al and V contents in coating layer and substrate region of as-sprayed Ta-coating sample.

3. Results and discussion 3.1. Surface morphology, surface roughness and cross-sectional microstructure of as-sprayed graded porous Ta-coating Fig. 2 shows the surface morphology and cross-sectional microstructure of the as-sprayed Ta-coating. As shown in Fig. 2(a), the coating has a rough and irregular surface with a small amount of uniformly distributed micropores (e.g., < 10 μm) and partially/completely melted particles. An inspection of Fig. 2(b) shows that the coating has an average surface roughness of Ra = 22.2 μm. By contrast, the original sandblasted Ti6Al4V substrate had a surface roughness of just 4.4 μm. In other words, the spray treatment yields a significant increase in the surface roughness. The resulting three-dimensional topography increases the total surface area of the implant and hence improves the tensile strength of the bone-implant interface [35]. In general, surface roughness can be divided into three levels depending on the scale of the features, i.e., nano- (1–100 nm), micro- (1–10 μm) and macro-sized (10 μm - mm)topologies [36]. Nano roughness plays an important role in the adsorption of proteins and the adhesion of osteoblastic cells, thereby enhancing the rate of osseointegration [37]. Meanwhile, micro roughness maximizes the interlocking effect between the mineralized bone and the implant surface [38]. The surface roughness of the present VPS Ta-coatings belongs to the macroscale level. Many studies have shown that macro level topographies result in both an improved early fixation and a better long-term mechanical stability of prostheses compared to smooth surfaces [35,36]. However, a higher surface roughness increases the risk of peri-implantitis due to greater bacterial adhesion and may also result in higher ionic leakage [39]. An inspection of Fig. 2(a) shows that the upper surface of the Tacoating has a porosity of approximately 13% and a pore size generally < 5 μm. Previous studies have shown that pore sizes < 15–50 μm lead to fibrovascular ingrowth, pore sizes of 50–150 μm promotes osteoid formation, pore size > 150 μm encourage bone ingrowth [40,41]. Such a porous structure in current study promotes fibrovascular ingrowth. As shown in Fig. 2(c), the VPS coating has a total thickness of around 350–380 μm. Moreover, the effectiveness of the VPS process in creating a porosity gradient within the coating is clearly observed. In particular, the porosity increases from 0.6 ± 0.2% near the substrate (i.e., the inner layer) to 7.6 ± 0.8% near the upper surface of the

Fig. 5. Young's modulus of Ti6Al4V substrate and graded porous Ta coating (inner, middle and outer layers).

bovine serum (FBS) in a humidified incubator with 5% CO2 at 37 °C for 3 h, 24 h and 48 h, respectively. Following the culturing process, the DMEM solution and non-adherent cells were removed from the specimen surface by washing three times with phosphate-buffered saline (PBS). The remaining cells were fixed using PBS containing 2.5 vol% glutaraldehyde, serially dehydrated using ethanol with an incrementally increasing concentration (30, 50, 70, 95 and 100%), and then critical-point-dried in hexamethyldisilazane (HMDS) for 10 min. The morphologies of the cells were then observed by SEM.

2.7. Surface characterization The microstructures of the various samples were observed using SEM (JEOL JSM-6390LV, JEOL Ltd., Tokyo, Japan). The chemical compositions were analyzed by EDS. Moreover, the various phases were identified using continuous mode XRD with Cu Kα radiation (Rigaku D/ Max Ш.V, Rigaku Ltd., Tokyo, Japan) with a 2θ scanning range of 20°~60° and a scanning rate of 2°min−1.

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Fig. 6. Failure fracture analysis results for Ta-coating: (a) schematic representation of fracture mode, (b) macro-fractrographs, (c) micro-fractrograph of substrate rod side and associated EDS spectra. Note spectrum 1 shows mainly Ta composition and spectrum 2 shows mainly epoxy composition. (CO: Cohesive failure, i.e., failure between similar molecules; AD: Adhesive failure, i.e., failure between dissimilar molecules.)

interface region (268 ± 35 HV0.1) lies between that of the Ti6Al4V substrate and the coating. Fig. 4 presents the EDS line scan results for the individual chemical elements of the coating and substrate near the interface. The results indicate that no chemical element inter-diffusion takes place between the coating and the substrate. Therefore, the intermediate hardness of the interface region is the result most likely of buffering between the high hardness region (i.e., the Ta-coating) and the low hardness region (i.e., the Ti6Al4V substrate). Fig. 5 shows the Young's modulus (E) distribution of the Ti6Al4V substrate and graded porous Ta-coating. The Young's modulus of the coating decreases gradually from 148 GPa in the inner layer to 123 GPa in the outer layer and is around 20–34% lower than that of bulk Ta (186 GPa). It is noted that the Young's modulus of the outer layer (123 GPa) is only slightly higher than that of HA ceramics (80–110 GPa) and is significantly lower than that of AW glass ceramics (218 GPa) [42]. Shirazi et al. [43] and Hedia and Mahmoud [44] examined the application of radially-distributed functionally-graded materials (FGMs) to implants. The finite element analysis (FEA) results showed that the E value of the considered material decreased from the cervical region of the implant to the apical region. The resulting distribution of the stress reduced the stress concentration in the cortical bone and implant compared with that in conventional titanium and stainless steel implants, and therefore reduced the stress shielding effect. Notably, the gradient can be distributed through the whole implant volume [45] or only the coating [7,44]. In the present study, the E value decreased from the inner layer of the graded porous Ta-coating to the outer layer,

coating (i.e., the outer layer). It is additionally seen in Fig. 2(c) that the Ta-coating has a dense microstructure free of microcracks. The absence of microcracks can be attributed to the preheating process performed prior to spraying and the close match between the coefficients of thermal expansion (CTE) of the Ta material (6.6 × 10−6 K−1) and Ti6Al4V substrate (8.8 × 10−6 K−1), respectively. The preheating process reduces the thermal gradient between the coating and the substrate, coupled with a low CTE mismatch, both of which minimizes the residual stress produced during the cooling. The interface between the Ta-coating and the Ti6Al4V substrate is similarly free of defects. Furthermore, no obvious surface discontinuities are observed. The dense crack-free coating and continuous coating-substrate interface result in a high bonding strength, as discussed later in Section 3.3.

3.2. Cross-sectional hardness and Young's modulus distribution of coating Fig. 3 shows the micro-hardness values of the Ti6Al4V substrate, substrate-coating interface, and graded porous Ta-coating (inner, middle and outer layers), respectively. It is seen that the hardness decreases smoothly from the substrate to the outer layer. The reduction in the Ta-coating hardness, i.e., from 240 ± 20 HV0.1 in the inner layer to 167 ± 62 HV0.1 in the outer layer, reflects the increasing porosity and decreasing melting degree of the coating from the inside to the outside. The outer layer of the coating contains more closed pores than the inner layers. In practice, it is impossible to avoid these pores during the measurement process. Consequently, the measurement error in the outer layer is higher than that in the inner layers. The hardness of the 403

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Fig. 7. Microstructure and EDS analysis of: (a) original Ta-coating, (b) AT Ta-coating, (c) AHT Ta-coating, (d) AHT-UC Ta-coating.

in the glue layer or outer coating layer of the specimen (i.e., between the Ta-coating and the uncoated counterpart bar). Fig. 6(a) presents a schematic representation of the test specimen before and after tensile testing. The fractographs in Fig. 6(b) and (c) show that the fracture originated mainly in the glue layer (i.e., glue cohesive (CO) failure), rather than the outer coating layer (i.e., coating cohesive failure) [46–48]. In addition, on the uncoated counterpart side, there are only a few fracture locations at the glue-substrate interface (i.e., glue adhesive (AD) failure). In other words, although the failure mode of the coating is technically speaking a mixed failure mode, the failure is dominated by cohesive failure. Anderson and Levine [49] attached a fiber‑titanium pad to a solid Ti6Al4V substrate by diffusion bonding, and obtained tensile strengths

and is thus similar to the reducing E value observed for the radiallydistributed FGMs. Consequently, the graded porous Ta-coating implant is also expected to result in a lower stress shielding effect. In other words, the graded porous Ta-coating implant reduces the stress shielding effect around the bone-implant interface compared to that in a solid or dense Ta implant [7]. In other words, the VPS Ta-coating provides an effective means of reducing the stress shielding effect in biometal implant applications. 3.3. Bond strength measurements and fracture surface investigation Fig. 6 presents the failure fracture analysis results for the Ta-coating specimen in the bond strength tests. In each test, the specimen fractured 404

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[51] prepared HA/Ta composite coatings with different Ta contents (20% and 60%) using the VPS method, and showed that the coating containing 60% Ta (40% HA) achieved a bonding strength of 37.2 MPa, i.e., around 1.9 times higher than that of the pure HA coating. According to the ASTM C633 test results obtained in the present study, the as-sprayed VPS Ta-coating had an average bonding strength of 54.5 ± 2.3 MPa. However, as described above, for most of the samples, the fracture zone was located in the glue layer rather than the outer coating layer. Therefore, it can be inferred that the cohesion of the three-layer coating and adhesion of the inner layer to the substrate exceed 54.5 MPa [49]. 3.4. Microstructure, EDS and XRD analysis of Ta-coating before and after surface treatments Fig. 7 shows the microstructure and EDS analysis results for the Tacoating before and after AT, AHT and AHT-UC treatment. As shown in Fig. 7(a), the VPS Ta-coating (without treatment) consists predominantly of Ta, with only a small quantity of O. The low O content indicates that the surface of the coating underwent only very slight oxidation during the spray process. Following alkali treatment (AT), the coating surface contains a large number of irregular block-shaped compounds, as shown in Fig. 7(b). The EDS results show that, besides Ta, the surface has a significant Na and O content. The Na stems from the aqueous NaOH solution used in the AT process. Moreover, the Ta, Na and O elements indicate the presence of sodium tantalate

Fig. 8. XRD patterns of: (a) original Ta-coating, (b) AT Ta-coating, (c) AHT Tacoating, (d) AHT-UC Ta-coating.

ranging from 29.3 to 65.6 MPa, depending on the particular heat treatment and sample preparation applied. Yao et al. [50] fabricated insitu titanium composite coatings with bond strengths of up to 55.4 MPa using a reactive atmospheric plasma spraying technique. Huang et al.

Fig. 9. SEM micrographs of specimens soaked in SBF for various periods: (a) Ti6Al4V substrate, (b) Ta-coating, (c) AHT-UC Ta-coating (x1000). 405

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Fig. 10. EDS spectra of AHT-UC Ta-coatings soaked in SBF for various periods: (a) 3 days, (b) 7 days, (c) 14 days.

the hydrogel layer is dehydrated and densified to form a stable amorphous or crystalline alkali sodium tantalate layer [53]. Finally, for the AHT-UC specimen (Fig. 7(d)), some of the densified layer is removed together with the irregular block-shaped compounds. At low magnification (×100), the surface of the AHT-UC specimen resembles that of the untreated Ta-coating (Fig. 7(a)). However, under higher magnification (×1000, Fig. 9), the coating is found to have a flatter, denser and more porous morphology. Notably, even though the samples have slightly different morphologies, there is no significant difference in their bioactivity or initial biocompatibility (see Sections 3.5 and 3.6).

(NaxTayOz). For the AHT sample, the irregular block-shaped compounds are smaller and smoother, as shown in Fig. 7(c). However, the chemical composition is essentially unchanged from that of the AT sample. Following ultrasonic cleaning (AHT-UC), the block-shaped compounds are completely removed, as shown in Fig. 7(d). In other words, the bonding force between the compounds and the coating is very weak. Nonetheless, the EDS results show the presence of a small quantity of Na and O, which indicates that the surface still contains a certain amount of sodium tantalate after the cleaning process. Fig. 8 shows the XRD patterns of the VPS Ta-coating and AT, AHT and AHT-UC coatings, respectively. The pattern for the untreated Tacoating contains only Ta peaks (α-Ta, JCPDS ref. card 04–0788), as shown in curve (a). Although a small amount of O is detected in the EDS analysis results (Fig. 7(a)), the XRD pattern shows that the O content is insufficient to produce Ta-O oxide or is very small amount of oxide. In other words, the XRD results confirm that the coating was properly protected from oxidization during the VPS process. For the AT coating, the spectrum contains not only Ta peaks, but also peaks corresponding to sodium tantalate (Na3TaO4, Na8Ta6O19·4H2O, Na14Ta12O37·31H2O) and Ta2O5 (Fig. 8, curve (b)). Following heat treatment, these compounds transform to Ta-O oxides (Ta2O5, Ta4O, Ta0.703O1.65) with sodium tantalate (Na5Ta11O30) as a secondary compound (Fig. 8, curve (c)). For the sample processed by both heat treatment and ultrasonic cleaning (AHT-UC), the XRD pattern contains only Ta and Ta4O peaks (Fig. 8, curve (d)). As discussed above, the cleaning process results in the removal of the block-shaped compounds formed during the AT process. Hence, the absence of Na5Ta11O30, Ta2O5 and Ta0.703O1.65 peaks in the XRD pattern confirms that these compounds are composed mainly of crystalline sodium tantalite and Ta-oxides. Referring back to Fig. 7(d), it is seen that the surface of the AHT-UC sample contains a small amount of sodium tantalate despite the removal of the irregular blocks during ultrasonic cleaning. Miyazaki et al. [32,52] reported that continuous scanning XRD cannot detect amorphous sodium tantalate on AT or AHT Ta metal surfaces treated with 0.2–0.5 M NaOH. In the present study, the continuous scanning mode XRD analysis also failed to detect amorphous sodium tantalate on the AHT-UC coating (Fig. 8, curve (d)). However, the EDS results presented in Fig. 7(d) suggest that amorphous sodium tantalate does in fact exist on the sample surface. In other words, the AT treatment with 1.0 M NaOH produces both crystalline (XRD detectable) and amorphous (XRD non-detectable) sodium tantalate phase, but the crystalline phase is removed during ultrasonic cleaning; leaving only the amorphous sodium tantalate phase in the surface layer of the coating. Notably, the presence of amorphous sodium tantalate is beneficial in improving the bioactivity of the Ta coating during immersion in simulated body fluid [32,52], as discussed below in Section 3.5. In general, the various surface treatments induce different morphological changes in the Ta coating. For example, for the Ta-coating treated with AT (Fig. 7(b)), the surface contains not only irregular block-shaped compounds, but also a densified layer. The EDS and XRD analysis results, together with the findings of [32,52], suggest that this layer is sodium tantalate hydrogel. For the AHT specimen (Fig. 7(c)),

3.5. Bone-like apatite inducement tests of Ta-coatings Fig. 9 shows the SEM surface morphologies of the Ti6Al4V substrate, Ta-coating and AHT-UC Ta-coating before and after soaking in SBF for 0–14 days. Prior to immersion in the SBF (0 days), the Ti6Al4V substrate has a smooth surface with some minor abrasion marks (Fig. 9(a)). By contrast, the Ta-coating and AHT-UC Ta-coating have rough and irregular surfaces (Fig. 9(b) and (c), respectively). After soaking in SBF for 3–14 days, neither the Ti6Al4V substrate nor the Tacoating show any signs of particle (apatite) formation. However, for the AHT-UC Ta-coating, hemispherical particles are formed, which cover the entire porous surface after just 3 days of immersion. The particles have an average size of around 3–5 μm. After 7 days, the particle size increases to 10–20 μm. Furthermore, after soaking for 14 days, the particles continue to grow in size to approximately 15–25 μm. Before soaking in SBF, the AHT-UC Ta-coating consists mainly of Ta, O and Na (Fig. 7(d)). However, after soaking in SBF for 3 days, the coating consists predominantly of Ca, P and O, which are the main components of bone-like apatite (Fig. 10(a)). Notably, the Ta and Na elements are greatly reduced compared to the original AHT-UC coating (i.e., 0 days' SBF immersion), which suggests that the Na is released from the surface of the specimen into the SBF [52] and the apatite covers almost the entire coating surface. As the SBF immersion time increases to 7 and 14 days, respectively, the Ta element completely disappears (Fig. 10(b) and (c)). Furthermore, there is very little change in the Ca, P and O peaks. Consequently, it is inferred that the apatite layer is almost completely formed after just 3 days of SBF immersion, but increases in thickness as the immersion time increases. Fig. 11 shows the XRD patterns of the Ti6Al4V substrate, Ta-coating and AHT-UC Ta-coating soaked in SBF for 0–14 days. As shown in Fig. 11(a) and (b), no change in the peak positions or intensity occurs in the Ti6Al4V and Ta-coating patterns before and after SBF immersion, respectively. In other words, neither sample exhibits apatite peaks after soaking for 14 days in SBF. For the AHT-UC Ta-coating, the XRD pattern (Fig. 11(c)) shows that the original coating (i.e., prior to SBF immersion) consists mainly of Ta with a small amount of Ta4O. However, following immersion in SBF, a peak corresponding to apatite appears. The intensity of this peak increases with an increasing soaking time, which indicates that the apatite quantity increases as the immersion time increases. Previous studies have also shown that Ta metal [32,52] or Ta oxide 406

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gel [54] subjected to AT or AHT to form amorphous sodium tantalate on their surface exhibit rapid apatite formation during the early soaking period in SBF. In these studies, the rapid growth of apatite was attributed to a release of Na+ ions from the amorphous sodium tantalate and the subsequent formation of Ta-OH groups on the sample surface via an exchange process with the H3O+ ions in the SBF. These groups then further combined with the Ca2+ and phosphate ions in the SBF to form apatite, as described in [10,32,52–55]. The possible effects of Ta4O on the bioactivity of Ta metals are not discussed in the open literature and are not easily identifiable since Ta4O is a metastable phase. However, related studies on Ta metals have shown that the effect of Ta-O oxides on bioactivity is significantly less than that of amorphous sodium tantalate [32,52]. Thus, it can be inferred that the apatite formation observed in the present study is the result mainly of the presence of amorphous sodium tantalate on the AHT-UC coating surface. For Ta plate metal processed by AT (with 0.2–0.5 M NaOH aqueous solution at 60 °C for 24 h [32]) or AHT (with 0.5 M NaOH aqueous solution at 60 °C for 24 h followed by heat treatment at 300 °C for 1 h [52]), the induction period for the formation of complete apatite on the surface of the specimen is shortened from around 4 weeks to approximately 1 week due to the formation of amorphous sodium tantalate on the sample surface. In the present study, dense apatite was formed on the surface of the AHT-UC Ta-coating after just 3 days' immersion in SBF. In other words, the apatite formation time is much shorter than that reported in previous studies [32,52]. This can be attributed most likely to the effect of the VPS process in producing a porous and rough coating surface. 3.6. Initial investigation into in vitro biocompatibility of various specimens In general, the results presented above suggest that the AHT-UC Tacoating is not only beneficial in enhancing the ingrowth of cortical bone into the porous structure, but also in promoting chemical integration with the bone via apatite formation on the implant surface [10]. The biocompatibility of the AHT-UC Ta-coating was thus further examined by means of in vitro MG-63 cell tests. Fig. 12 presents SEM images showing the attachment, growth and spreading of the MG-63 cells on the bare Ti6Al4V substrate (control sample), the Ta-coating, and the AHT-UC Ta-coating after culturing for 3 h, 24 h and 72 h. After 3 h, no obvious spreading of the cells is observed on the Ti6Al4V substrate. However, a uniform attachment and spreading is observed for both the Ta-coating and the AHT-UC Ta-coating. After 24 h, filopodia and/or lamelliopodia appear on all the samples; indicating that the cells have begun to grow and spread. However, the cells on the Ta-coating and AHT-UC Ta-coating extend filopodia over a wider surface area than those on the bare Ti6Al4V sample. After 48 h, all of the samples show a good proliferation and spreading of the cells. In other words, no immediate cytotoxic effects are present for any of the samples. However, both VPS samples exhibit a greater cellular spreading rate than the Ti6Al4V sample. Thus, the results confirm that the surface chemistry properties of the VPS Ta-coating are beneficial in enhancing the biocompatibility of the Ti6Al4V substrate. Balla et al. [15] showed that the proliferation of hFOB cells on the surface of a laser-processed Ta-coating was around six times higher than that on a Ti control surface after 14 days of culturing. Similarly, Tang et al. [19] showed that a porous VPS Ta-coating not only enhanced the adhesion and proliferation of hBMSC cells compared to a porous VPS Ti-coating, but also stimulated mesenchymal stem cell osteogenetic differentiation in vitro. In both studies [15,19], the improved cell-material interaction of the Ta-surface was attributed to its high wettability and high surface energy. Many studies [56–59] have shown that the surface topography, surface chemistry and surface energy have a significant effect on the biologic response of osteogenic cells and hence on the ultimate clinical success of bioimplants. The Tacoatings fabricated in the present study have a surface pore structure (see Fig. 2(a)) and a rough surface morphology, and are therefore

Fig. 11. XRD patterns of specimens soaked in SBF for 0–14 days: (a) Ti6Al4V substrate, (b) Ta-coating, (c) AHT-UC Ta-coating. 407

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(a) Ti6Al4V

(c) AHT-UC Ta-Coating

(b) Ta-Coating

3 hr

x2.0k

20 μm

x2.0k

20 μm

x2.0k

20 μm

x2.0k

20 μm

x2.0k

20 μm

x2.0k

20 μm

x2.0k

20 μm

x2.0k

20 μm

x2.0k

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48 hr

Fig. 12. SEM images of MG-63 cells cultured on various specimens for 3–48 h: (a) Ti6Al4V substrate, (b) Ta-coating, (c) AHT-UC Ta-coating (x2000).

beneficial in promoting fibrovascular ingrowth and implant fixation. Moreover, through an appropriate treatment (i.e., AHT or AHT-UC), the SBF bioactivity of the Ta-coating can be significantly enhanced; resulting in an improved long-term stability of the implant.

5. In MG-63 cell culture tests, the cells cultured on the Ta-coating and AHT-UC Ta-coating began to spread on the coating surface after 3 h. Both coatings showed a superior biocompatibility to that of the uncoated Ti6Al4V substrate. However, no obvious difference in the biocompatibility of the two coatings was observed. Thus, it can be inferred that neither the alkali treatment nor the heat treatment had any significant effect on the initial biocompatibility of the original VPS Ta-coating.

4. Conclusion 1. Three-layered Ta-coatings have been deposited on Ti6Al4V substrates using the VPS technique. It has been shown that through an appropriate selection of the powder particle size and spray parameters, a porosity gradient can be produced from the inner layer to the outer layer. Moreover, the adhesion of the Ta coating to the substrate can be improved by preheating the substrate to 650 °C prior to the spray process. 2. The hardness and elastic modulus of the Ta-coating reduced from 240 ± 20 (inner layer) to 167 ± 62 HV0.1 (outer layer) and 148 ± 5 (inner layer) to 123 ± 4 GPa (outer layer), respectively. In bond strength tests, the specimens failed mainly in the glue layer and outer coating layer. The average bonding strength was found to be 54.5 ± 2.3 MPa, which exceeds the tensile load experienced in most bioimplant applications. 3. Following alkali treatment (AT), the Ta-coating contained not only Ta, but also sodium tantalate (Na3TaO4, Na8Ta6O19·4H2O, Na14Ta12O37·31H2O) and Ta2O5. After heat treatment (AHT), the compounds were converted mainly to Ta-O oxides (Ta2O5, Ta4O, Ta0.703O1.65) with sodium tantalate (Na5Ta11O30) as a secondary compound. Following ultrasonic cleaning (AHT-UC), the coating consisted of only Ta, Ta4O and amorphous sodium tantalate. 4. After immersion in SBF for 3–14 days, the surfaces of the Ti6Al4V substrate and Ta-coating showed no signs of apatite formation. However, for the AHT-UC Ta-coating sample, apatite was formed after just 3 days and covered the entire Ta surface. The enhanced bioactivity can be attributed to the formation of amorphous sodium tantalate during the AT process.

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