Evaluation of cell activation promoted by tantalum and tantalum oxide coatings deposited by reactive DC magnetron sputtering

Surface & Coatings Technology 330 (2017) 260–269

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Evaluation of cell activation promoted by tantalum and tantalum oxide coatings deposited by reactive DC magnetron sputtering

MARK

Hugo Moreiraa,⁎, Augusto Costa-Barbosab, Sandra Mariana Marquesa, Paula Sampaiob, Sandra Carvalhoa,c,⁎ a b c

GRF-CFUM, Physics Department, University of Minho, 4800-058 Guimarães, Portugal CBMA, Department of Biology, University of Minho, 4710-335 Braga, Portugal SEG-CEMUC, Mechanical Engineering Department, University of Coimbra, 3030-788 Coimbra, Portugal

A R T I C L E I N F O

A B S T R A C T

Keywords: Cell activation Dental implant Inflammatory response Macrophage cell Tantalum Tantalum oxide

Previous studies have shown both Ta and TaeO to be bioactive, rapidly forming a strongly-bonded surfaceadherent layer of bone-like hydroxyapatite (HAp) when immersed in simulated body fluid (SBF). Consequently, Ta and TaeO coatings are promising for the surface-modification of Ti or stainless steel endodontic endosseous implants, being conducive to a reduction in the risk of developing post-operatory infection and/or peri-implantitis disease. That said, few studies have investigated the effect of Ta or TaeO coatings on such phenomena as cell activation, adhesion, and proliferation. To that effect, Ta, and TaeO films were deposited onto type 316L stainless steel (SS 316L) substrates by reactive DC magnetron sputtering, after which their biological response was evaluated following co-incubation with the murine macrophage-like cell line RAW 264.7. Cell morphology after adhesion was observed by SEM, whereas cell viability and proliferation were evaluated by the 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) colorimetric assay. Lastly, inflammatory response was assessed by quantification of the cytokines interleukin-6 (IL-6) and interleukin-10 (IL-10). In terms of phase composition, Ta showed a mixture of the α-Ta and β-Ta phases, whereas TaeO showed a nanocrystalline structure. Moreover, a decrease in average roughness (Ra) from 21 nm to 7 nm was observed between Ta and TaeO, accompanied by a decrease in water contact angle (θW) from 106° to 83°. In vitro studies showed that cells exhibited significantly better adhesion to Ta, in comparison with both TaeO and SS 316L. Furthermore, both Ta and TaeO were shown to be non-cytotoxic, with Ta outperforming TaeO in terms of relative cell viability, both at 24 h and 48 h. Lastly, both Ta and TaeO showed vastly inferior IL-6 and IL-10 levels to those obtained for cells treated with bacterial lipopolysaccharides (LPS)—prompting the conclusion that the coatings do not in any way induce an inflammatory response from macrophage cells.

1. Introduction The oral mucosa hosts a variety of microbial species ranging from bacteria to fungi [1], the proliferation of which is controlled by physiological barriers such as antimicrobial factors and the mucosa-associated lymphoid tissue (MALT)—a diffuse system of small concentrations of lymphoid tissue populated by plasma cells, macrophage cells, and lymphocytes [2]. However, the insertion of endodontic endosseous implants disrupts the continuity of these barriers, inducing a local inflammation with a huge potential for degenerating into severe infection due to the ingress and unhinged proliferation of potentially harmful

opportunistic microorganisms [3]. Monocyte and macrophage cells play a key role in the early stages of tissue healing after implant insertion, acting not only to regulate inflammation and dispel infection, but also to promote bone healing and osseointegration [4]. In short, macrophage attachment and activation to the implanted materials is crucial in determining the extent of acute/ chronic inflammation [5]. Nowadays, commercially pure Ti (CP Ti) and Ti alloys (e.g. Ti–6Al–4V) are the most extensively used materials in orthodontic reconstructive surgery, due to their excellent corrosion resistance, passivation capacity, and biocompatibility [6]. However, Ti implants are not

Abbreviations: HAp, hydroxyapatite; MTT, 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide; IL-6, Interleukin-6; IL-10, Interleukin-10; LPS, Lipopolysaccharides; SBF, simulated body fluid; SS 316L, type 316L stainless steel ⁎ Corresponding authors at: University of Minho, Physics Department, Campus of Azurém, 4800-058 Guimarães, Portugal. E-mail addresses: [email protected] (H. Moreira), sandra.carvalho@fisica.uminho.pt (S. Carvalho). http://dx.doi.org/10.1016/j.surfcoat.2017.10.019 Received 16 May 2017; Received in revised form 19 September 2017; Accepted 6 October 2017 Available online 07 October 2017 0257-8972/ © 2017 Elsevier B.V. All rights reserved.

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2. Materials and methods

entirely without fault: an approximate 14% of inserted implants go on to develop peri-implantitis [7]. Following insertion, Ti implants show an inconvenient propensity for becoming encapsulated in fibrous tissue [8]. At first, this results in a poor bond between implant and bone, and, over time, in implant loosening as a result of poor osseointegration [8]. Implant-associated infection—which results from the reduced immune resistance of the host following surgical trauma, as well as the implant itself acting as a foreign body and thus increasing the risk of infection by prompting the entry and proliferation of microorganisms—is another leading cause for Ti implant failure [8]. Particularly troublesome, however, is the case of peri-implantitis disease, in which the infection of the implant-adjacent bone tissue leads to a receding of the surrounding bone, the subsequent decrease of the biomechanical anchorage of the implant, and, ultimately, implant failure [9,10]. While the mechanical properties of the materials and the loading conditions in the host significantly influence material selection [6], cell and tissue interactions with the implant surface depend dominantly on surface characteristics, with rough, textured, and porous surfaces stimulating cell adhesion, differentiation, and the formation of extracellular matrices [4,11]. In an experiment conducted with surface oxide films formed on Ti plates, for example, it was shown that surface topography, roughness, and energy, as well as the concentration of surface-attached OH groups, significantly influence the initial behaviour of osteoblasts [11]. Thus, one of the most promising approaches in the way of improving implant in-service behaviour has been the surfacemodification of pre-existing Ti-based implants as a demand for bioactive surfaces that enhance the implant healing process and promote biomineralisation. Ta, in particular, has proven to be a promising alternative to Ti, being widely documented as bioactive [12,13] and having been shown to rapidly form a strongly-bonded surface-adherent layer of bone-like hydroxyapatite (HAp) when immersed in simulated body fluid (SBF) due to its high surface energy in comparison with CP Ti [14]. While the high density and cost of Ta implants limit their bulk use [13], a suitable compromise is found in the deposition of Ta thin films onto standard Ti or stainless steel implants. Indeed, studies have shown Ta coatings to improve the in vitro biocompatibility of CoeCr [12] and TieNi alloys [12,15], for example. On the other hand, the development of implants with oxidised surfaces constitutes another promising approach. In in-service conditions, osteoblasts interact with the oxidised surface, and, due to the oxide layer's ability to bind with Ca, form a diffusion zone, thus promoting a stronger bond between bone and implant [16]. TaeO layers were shown to improve the cytocompatibility of Ti, for example, with in vitro tests showing that the coatings promote the proliferation, alkaline phosphatase (ALP) activity, mineralisation, and osteogenic gene expressions of osteoblasts [13]. Of particular interest, in a study conducted with TaeO coatings by Almeida Alves et al., it was shown that TaeO shows higher HAp formation rates than both CP Ti and Ta when immersed in SBF, which could translate to better bioactivity and osseointegration [14]. Moreover, it was shown that the higher the O content of the coatings, the higher the Ca/P ratio of the bone-like HAp layer formed on their surface [14]. However, while the bioactivity of Ta and TaeO coatings in view of SBF has been successfully verified in vitro [14], few studies have investigated the effect of these coatings on phenomena such as cell activation, adhesion, and proliferation. Bioactivity notwithstanding, these coatings would be inapplicable were they to elicit an inflammatory response in the host. As such, the aim of this paper is to produce, characterise, and study the effect of Ta and TaeO coatings on the activation, adhesion, proliferation, and secretion of paracrine factors of macrophage-like cells.

2.1. Production of coatings Thin films were sputter-deposited from a high-purity Ta target (99.95% Ta) (200 × 100 mm2) onto SS 316L (20 × 20 mm2) and ptype (B-doped) Si (100) by DC magneton sputtering. In order to determine optimal deposition conditions, hysteresis curves were constructed for target current density (JTa) of 10 mA/cm2 and 5 mA/cm2, under constant Ar flow of 60 sccm and bias voltage of − 75 V. The target (Ta cathode) voltage and working pressure were measured for increments of 5% (0.75 sccm) in the O2 flow, with time step of 1 min between readings. Ahead of depositions, substrates were ultrasonically cleaned in distilled water, ethanol, and acetone, for 10 min each, with a Sonica 2400MH S3 ultrasonic cleaner (Soltec, Italy). Substrates were then sputter-etched under Ar flow of 80 sccm and JTa of 0.5 mA/cm2 for 15 min. Throughout etchings, a pulsed DC (PDC) was applied to the substrate-holder, with pulse width of 1536 ns, frequency of 200 kHz, and intensity of 250 mA. Ta depositions were carried out under JTa of 10 mA/cm2 for 2 h, whereas TaeO depositions were carried out under JTa of 5 mA/cm2 and O2 flow of 13 sccm for 4 h. All depositions were carried out under Ar flow of 60 sccm and bias of − 75 V. For the TaeO depositions, a Ta interlayer (∼ 200 nm) was deposited onto the substrates under JTa of 10 mA/cm2 for 10 min in order to improve film-substrate adhesion. For all proceedings, the substrate-holder's distance to the target was kept at 70 mm, its rotation speed at 7 rpm, and its temperature at around 200 °C. Lastly, the sputtering chamber base pressure and the working pressure never exceeded 8 × 10− 4 Pa and 8 × 10− 1 Pa, respectively. 2.2. Characterisation of coatings The phase composition of the coatings was determined by X-ray diffraction (XRD) with a D8 Discover diffractometer (Bruker, Germany), operating at 40 kV and 40 mA with Cu Kα radiation (λKαI = 1.540562 Å and λKαII = 1.544390 Å [17]). Tests were carried out with grazing angle of 1°, step size of 0.04°, time step of 1 s, and 2θ range of 10–80°. The morphology of the coatings was observed by scanning electron microscopy (SEM) with a Nova NanoSEM 200 microscope (FEI, USA), operating at 5 kV in secondary electron (SE) mode. Chemical composition was determined by energy-dispersive X-ray spectrometry (EDS) with a Pegasus X4M spectrometer (EDAX, USA), operating at 20 kV. The topography of the coatings was observed by atomic-force microscopy (AFM) with a NanoScope III AFM apparatus (Digital Instruments, USA), operating in tapping mode. AFM micrographs were taken over scanning areas of 5 × 5 μm2. Film roughness was obtained through the roughness subroutine of the AFM apparatus. Analyses were performed on coatings deposited onto Si, due to its lower surface roughness in comparison with SS 316L, which results in more homogenous coatings—thus improving the reproducibility of the tests. 2.2.1. Analysis of wettability The wettability of the coatings was evaluated by the sessile drop test. Static contact angles were measured at room temperature (RT) with a OCA20 optical contact angle measuring system (DataPhysics, Germany), in view of Milli-Q ultrapure water, α-bromonaphthalene, and glycerol. Probe liquids were dosed with a Hamilton 500 μL syringe, with dosing volume of 2 μL and dosing rate of 1 μL/s. Contact angles were measured for thin films deposited onto SS 316L, with uncoated SS 316L substrates having been used as commercial controls. Measurements were taken after the probe liquid droplet reached equilibrium, after approximately 1 min. 261

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DMSO:ethanol (1:1). Subsequently, 100 μL/well of the MTT formazan solution was plated onto 96-well tissue culture plates (Becton Dickinson, USA) and absorbance measured at 570 nm and RT on a Synergy HT microplate reader (BioTek, USA). Untreated cells were used as a control of viability, whereas cells incubated with 3 mL of DMSO:DMEM (1:4) for 5 min were used as a control of cytotoxicity. Lastly, cell viability was determined according to Eq. 1.

2.3. In vitro studies In vitro studies were performed using the murine macrophage-like cell line RAW 264.7. Cells were cultured in Dulbecco's modified Eagle's medium (DMEM) (Lonza, Switzerland) supplemented with 2 mM of Lglutamine (Lonza, Switzerland), 1 mM of Na pyruvate (Merck, USA), 10 mM of 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES) (Merck, USA), and 10% (v/v) of heat-inactivated foetal bovine serum (FBS) (Lonza, Switzerland) at 37 °C under a humidified atmosphere of 5% CO2. After confluent growth, cells were washed, recovered, and resuspended in complete DMEM to the final concentration of 1 × 105 cells/mL. For all experiments, Ta and TaeO films deposited onto SS 316L, as well as uncoated SS 316L substrates, were plated onto 6-well tissue culture plates (TPP, Switzerland) and immersed in 70% ethanol for 2 h. Following sterilisation, samples were washed with sterile water and air dried, after which the cellular suspension was plated at 3 × 105 cells/ well and incubated to adhere overnight at 37 °C under a humidified atmosphere of 5% CO2.

Viability (%) Experimental value (average ) − Control of cytotoxicity (average) = Control of viability (average) − Control of cytotoxicity (average ) × 100

(1)

2.3.3. Quantification of cytokine production The enzyme-linked immunosorbent assay (ELISA) was employed in order to quantify cytokines in the cell culture supernatant, previously collected in 2.3.2. The pro-inflammatory cytokine interleukin-6 (IL-6) and anti-inflammatory cytokine interleukin-10 (IL-10) were quantified using the Mouse IL-6 and Mouse IL-10 ELISA kits (Invitrogen, USA), respectively, following the manufacturer's instructions, and absorbance measured at 450 nm and RT on a SpectraMax Plus 384 microplate reader (Molecular Devices, USA). Lipopolysaccharides (LPS) bind with the CD14/TLR4/MD-2 receptor complex in cells such as monocytes and macrophages, promoting the secretion of pro-inflammatory cytokines [20]. As such, cells treated with LPS were used as a positive control of inflammation, whereas untreated cells were used as a negative control of inflammation.

2.3.1. Observation of cell morphology Following 24 h and 48 h of co-incubation, the cell culture medium was discarded and cells washed with phosphate-buffered saline (PBS). Cells were then fixed with 2.5% glutaraldehyde (Alfa Aesar, USA) at 4 °C for 24 h, dehydrated through a series of 3 × 15 min incubations in ethanol solutions with increasing concentration (30%, 50%, 70%, 80%, 90%, and 100%), and air dried for 20 min. Lastly, cells were sputtercoated with a ∼ 15 nm-thick layer of AuePd alloy and observed by SEM with a JSM-6010 LV microscope (JEOL, Japan), operating at 5 kV in SE mode.

2.4. Statistical analysis Where applicable, results are reported as Mean ± Standard Deviation (SD). For wettability analysis, statistical analyses were performed using an ordinary one-way analysis of variance (ANOVA) test, with Tukey's multiple comparisons test having been used for intergroup comparisons. For roughness analysis and in vitro studies, statistical analyses were performed using ordinary two-way ANOVA tests, with Sidak's multiple comparisons tests having been used for intergroup comparisons. Differences among means were deemed significant for p ≤ 0.05.

2.3.2. Assessment of cell viability The MTT assay is a colorimetric assay used for the assessment of cell metabolic activity, viability, and proliferation [18]. The water-soluble tetrazolium salt 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) yields a yellowish aqueous solution, which is converted into a water-insoluble purple formazan product upon cleavage of its tetrazolium ring by dehydrogenases within metabolically active cells [18,19]. This lipid-soluble formazan product may then be extracted with organic solvents (e.g. dimethyl sulfoxide (DMSO)) and estimated by spectrophotometry—with the MTT production being inversely correlated with cell death [18]. Following 24 h and 48 h of co-incubation, 1 mL/well of the cell culture supernatant was collected and frozen at −20 °C for further analysis, whereas the remaining culture medium was discarded and replaced with 2.5 mL/well of complete DMEM and 250 μL/well of 5 mg/mL thiazolyl blue tetrazolium bromide (MTT) (Sigma, USA). Following 2 h of incubation, the supernatant was discarded and the formed formazan product solubilised with 2.5 mL/well of

3. Results and discussion 3.1. Hysteresis behaviour In order to determine optimal deposition conditions, hysteresis curves were constructed for JTa of 10 mA/cm2 and 5 mA/cm2 (Fig. 1). For JTa of 10 mA/cm2 (Fig. 1A), the target voltage increased from 376 V to 537 V when the O2 flow was increased from 3 sccm to 15 sccm. For JTa of 5 mA/cm2 (Fig. 1B), the target voltage increased Fig. 1. Hysteresis curves for JTa of A) 10 mA/cm2 and B) 5 mA/cm2, respectively. ΦAr and T denote the Ar flow and temperature, respectively.

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from 365 V to 630 V when the O2 flow was increased from 0 sccm to 12 sccm. It is known that the sputter rate of metals drops drastically when compounds form on the targets [21]. Moreover, deposition rates decrease in excess of 50% as a consequence of the lower sputter yield of the formed compounds in comparison with the original target materials [21]. The hysteresis effect, i.e. the abrupt transition of the sputtering rate and reactive gas partial pressure with increasing/decreasing gas flow, was described by Berg et al., and allows the identification of distinct metallic and compound deposition modes, separated by a steep transition zone [22,23]. At low reactive gas flows (metallic mode), the reactive gas flow is insufficient in order to react with the target and the target voltage is approximately constant. As the reactive gas flow increases, the reaction between reactive gas and target takes place, causing an abrupt change in the target voltage. Finally, at high reactive gas flows (compound mode), the reaction between reactive gas and target concludes and the target voltage becomes constant once more. This phenomenon is known as target poisoning, and is accountable for the target voltage and working pressure behaviour observed in Fig. 1. Indeed, as target poisoning takes place, larger amounts of O are incorporated into the Ta target matrix, prompting the formation of TaeO compounds on the target, which, due to their ceramic nature, effect an increase in the target voltage. On the other hand, it is also known that deposition rates depend both on the sputtering pressure and the applied current, being generally proportional to the square of the applied current density [21]. It stands to reason, then, that higher applied current densities correlate to higher deposition rates. For JTa of 10 mA/cm2 (Fig. 1), even relatively high O2 flows (15 sccm) were insufficient to reach the compound mode. In order to tackle this hindrance, JTa was halved to 5 mA/cm2 (Fig. 1B) and the compound mode reached for O2 flow of 12 sccm. Deposition conditions were adjusted accordingly, with Ta coatings having been deposited under JTa of 10 mA/cm2 and TaeO coatings under JTa of 5 mA/cm2 and O2 flow of 13 sccm, as previously described in 2.1. Additionally, in order to counter the theoretically predicted decrease in deposition rate, deposition times were doubled for the TaeO coatings, from 2 h to 4 h. Lastly, the discussed hysteresis behaviour is in good accordance with experimental results obtained for thin films deposited by reactive magnetron sputtering by Wang et al. for VO [24], Baroch et al. for TiOx [25], and Cristea et al. for TaNxOy [26].

Fig. 2. EDS spectra of the Ta and TaeO coatings. The Ta MαI, Ta LαI, Ta LβI, Ta LβII, and O KαI emission lines are highlighted in the figure.

Table 1 Chemical formula, composition, thickness, and deposition rate of the coatings. Sample group

Ta TaeO

Chemical formula

Ta0.95O0.05 Ta0.34O0.66

Chemical composition (at.%) Ta

O

95.0 33.7

5.0 66.3

Thickness (μm)

Deposition rate (μm/h)

4.6 6.7

2.3 1.7

et al. under similar deposition conditions [14]. 3.3. Phase composition In order to properly comprehend the influence of the O content on the developed structures, XRD tests were carried out on Ta and TaeO films deposited onto Si (Fig. 3). For the Ta coating, three diffraction peaks were identified, at 2θ of 36.76° (I), 38.24° (II), and 69.29° (III), respectively. While peaks I and II show a good fit for tetragonal β-Ta (ICDD card no 00–025-1280), peaks

3.2. Chemical composition and deposition rate The chemical composition of the coatings was determined by EDS for thin films deposited onto Si (Fig. 2). The Ta coating showed Ta content of 95 at.% and a residual O content of 5 at.%, whereas the TaeO coating showed Ta content of 34 at.% and O content of 66 at.%. Moreover, for the TaeO coating, the obtained Ta/O ratio (∼ 0.51) seems to suggest a TaO2-type stoichiometry. Film thickness was determined by SEM for the same films. The obtained cross-section micrographs (Fig. 4) showed that Ta had film thickness of 4.6 μm, whereas TaeO had thickness of 6.7 μm. The chemical composition, thickness, and deposition rate of the coatings are summarised in Table 1. As expected, the increase in the O2 flow from 0 sccm to 13 sccm resulted in an increase in the O content of the coatings, from 5 at.% to 66 at.%. In contrast, the deposition rate dropped by about 26% between Ta and TaeO. As detailed in 3.1, this decrease in deposition rate can be attributed to the decrease in the sputtering rate of the target as a result of both target poisoning due to O2 flow increase and the decrease in JTa from 10 mA/cm2 to 5 mA/cm2. Results for both chemical composition and deposition rate are in good accordance with experimental results obtained by Almeida Alves

Fig. 3. XRD spectra of the Ta and TaeO coatings. The 2θ range of 20–40° is highlighted for TaeO, revealing a broad diffraction peak.

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II and III show a good fit for body-centred cubic (BCC) α-Ta (ICDD card no 00–004-0788), prompting the conclusion that the Ta coatings are composed of a mixture of both the α-Ta and β-Ta phases. Deposition parameters such as temperature and bias voltage play a key role in the development of specific crystalline structures [27–29]. On the one hand, the application of a negative bias voltage to the substrates during deposition has a prominent impact on the structural properties of the films, due in part to the modification of atomic arrangement by selective removal of weakly-adsorbed contaminants [28]. On the other hand, the energy of the bombarding ions during deposition strongly influences the crystallographic structure of the films, with rearrangements of atoms and atomic displacements on the growing film resulting in a change in the crystal lattice type of the coating [28]. While the literature shows that the formation of α-Ta—the preferred phase in protective coatings subject to loads and wear—typically occurs for substrate temperatures above 700 °C [30], the occurrence of α-Ta is possible at relatively lower temperatures when a negative bias is applied [29], as was the case in this paper. However, as was earlier established, the Ta coatings are composed of a mixture of both Ta phases, prompting the conclusion that the applied bias was insufficient in order to obtain α-Ta exclusively. For the TaeO coating, no diffraction peaks were identified. However, closer analysis of the diffraction pattern revealed a broad diffraction peak in the 2θ range of 20–40° (highlighted in Fig. 3) and peak maxima at 2θ of 28.24° and 53.52°, thus suggesting a nanocrystalline structure. Moreover, the broad peak of the pattern coincides with the location of several peaks for oxide phases such as tetragonal TaO2 (ICDD card no 00–019-1297) and orthorhombic Ta2O5 (ICDD card no 00–025-0922), implying that either one or a mixture of these phases may be present. The nanocrystalline structure of the TaeO coatings is in good accordance with structures developed in O-containing thin films deposited by reactive magnetron sputtering, such as WeO [31]. Yang et al. obtained similar diffraction patterns for TaeO films deposited by reactive unbalanced magnetron sputtering [32], whereas Cristea et al. experienced a decrease in crystallinity as a function of increasing reactive gas flow for TaNxOy films, reporting a similar broad peak in the 2θ range of 24–37° [26]. Guimarães et al. showed for Ti–Si–C–ON coatings that at high O2 flows the incorporation of O in the coatings promotes the precipitation of oxide phases in the grain boundaries, thus inhibiting grain boundary migration and grain growth, and, subsequently, decreasing grain size—consequently leading to the formation of an amorphous structure [33]. Lastly, Chang et al. obtained the β-Ta phase for Ta films deposited by magnetron sputtering under temperature of 110 °C and bias of − 75 V, whereas an amorphous structure was obtained for as-deposited Ta2O5, having reported a similar broad peak in the 2θ range of 20–40° [12]. Of particular interest, however, Ta2O5 films were shown to recrystallize according to the crystalline structure of β-Ta2O5 after rapid thermal annealing at 700 °C [12].

amorphous morphologies [34], suggesting that the addition of O promotes the densification of the coatings. The roughness of the coatings was determined through the roughness subroutine of the AFM apparatus (Table 2). Ra denotes the average roughness, whereas Rq denotes the root mean square (RMS) roughness, as described by Raposo et al. [35]. As seen in Table 2, a decrease in roughness in excess of 50% was observed between Ta and TaeO, from Ra of 21 nm to 7 nm, and from Rq of 26 nm to 9 nm. Differences among means between Ta and TaeO were found to be significant, both for Ra (p ≤ 0.01) and Rq (p ≤ 0.001). AFM results are in good accordance with the literature, with similar morphological evolutions having been reported as a function of increasing O content by Vaz et al. for ZrNxOy films deposited by reactive magnetron sputtering [36]. 3.5. Wettability Wettability, defined as the degree to which a liquid will maintain contact with a solid surface, is a crucial surface property in biomaterials, having been shown to be linked to such phenomena as protein adsorption/desorption, cell adhesion, and phagocytosis [37]. The wettability of the coatings was evaluated through the quantification of their contact angles, surface energy, and surface free energy (Table 3). Contact angles were measured for thin films deposited onto SS 316L, with uncoated SS 316L substrates having been used as commercial controls. θW, θB, and θG denote contact angles in view of Milli-Q ultrapure water, α-bromonaphthalene, and glycerol, respectively. γLW, γ+, and γ− denote the dispersive apolar Lifshitz–van der Waals (LW), positive polar, and negative polar surface energy components, respectively, as described by the Van Oss–Chaudhury–Good (VCG) theory of wettability [38]. Lastly, ΔG denotes the total surface free energy. Additionally, the contact angles and wetting behaviour in view of Milli-Q water for all sample groups are displayed in Fig. 5. As seen in Table 3, in view of Milli-Q water, the Ta coatings (106°) showed much higher θW in comparison with both TaeO (83°) and SS 316L (76°), revealing the most hydrophobic character. Regarding ΔG, similar trends were observed, with Ta showing the lowest ΔG (− 85 mJ/m2), followed by TaeO (−63 mJ/m2) and SS 316L (− 36 mJ/m2). Differences among means were found to be significant between all sample groups (p ≤ 0.0001). According to Vogler, a surface is deemed hydrophilic for θW ≤ 65° [39], prompting the conclusion that all sample groups are hydrophobic. Furthermore, the decrease in θW with the increase in O content observed between Ta and TaeO was explained by Sharma and Paul as due to the increasing oxidation of the coatings, with oxide layer-coated surfaces having been reported as showing improved wettability and fibrinogen adsorption in comparison with bare Ta [40]. Yang et al. showed that the wettability of TaeO coatings is crucially dependant on the O2/Ar ratio, with maximum (45 mJ/m2) and minimum (40 mJ/m2) surface energies having been obtained for O2/Ar ratios of 0.5 and 1.2, respectively [32]. Furthermore, the authors also linked the low surface energy and dispersive energy component of the coatings as being advantageous to good blood compatibility and antithrombotic performance [32]. Chang et al. obtained similar θW for β-Ta (97°) and as-deposited amorphous Ta2O5 (86°) [12]. Interestingly, annealed β-Ta2O5 showed a colossal decrease in hydrophobicity, showing θW of 6° [12]. This decrease in θW was attributed to the severing of bonds between the surface and O following annealing, resulting in a more chemically-stable hydrophilic surface [12]. Moreover, while the higher contact angles of amorphous Ta2O5 were linked to its better antibacterial behaviour, its cell viability in view of human fibroblast cells was relatively poor, as opposed to β-Ta2O5 [12]. Results suggest that surface chemistry is crucial to the description of the wettability. For both Ta and TaeO, the polar surface energy components γ+ and γ− were close to zero, indicating that the surfaces are approximately apolar. A slight increase was observed in γ− between Ta

3.4. Morphology, topography, and roughness The morphology and topography of the coatings were observed by SEM and AFM for thin films deposited onto Si (Fig. 4). For the Ta coating, the SEM surface micrograph shows structures typically observed in coatings grown with columnar morphologies. For the same coating, the SEM cross-section micrograph corroborates these findings, revealing a columnar morphology with column width of a few nanometers. This type of microstructure consists of a network of lowdensity material surrounding bundles of higher-density rod-like columns, and is characteristic of coatings grown with limited ad-atom mobility [34]. For the TaeO coating, the SEM surface micrograph shows a smoother surface, typically observed in compact coatings grown with 264

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Fig. 4. SEM and AFM micrographs of the Ta and TaeO coatings. SEM micrographs were taken in SE mode. AFM micrographs were taken in tapping mode over scanning areas of 5 × 5 μm2.

appearance were found. Moreover, cytoplasmic extensions were shown to form characteristic ring-like shapes in between adjacent cells. After 48 h, cell clusters coalesced, spreading across larger film surfaces. Interestingly, cells seemed to concentrate around the contours of the earlier formed ring-like structures, suggesting this to be their preferred mechanism of proliferation. In terms of spread appearance, larger amounts of fully spread cells were found, as well as more pronounced cytoplasmic extensions. For the TaeO coatings, cells were more or less evenly dispersed along the surface of the films after 24 h, with no evidence of clusters or preferential loci for cell proliferation. However, cells were more rounded in shape, with fewer fully spread cells. Cytoplasmic extensions were also thinner and weaker, forming characteristic dumbbell-like shapes linking adjacent cells. After 48 h, cells showed uniform proliferation along the surface of the films. Moreover, the dumbbell-like structures began to coalesce, giving rise to the ring-like structures found for Ta. For the SS 316L substrates, similar trends were observed in terms of cell proliferation. However, cell adhesion was poor, with cells exhibiting predominantly rounded morphology, both at 24 h and 48 h. Moreover, cells were more warped and uneven, and cell networks more erratic and unorganised. While at 24 h TaeO and SS 316L showed more densely-packed surfaces in comparison with Ta, cell proliferation seemed to be better for Ta than for the remaining sample groups. Moreover, cells showed better spread appearance for Ta in comparison with both TaeO and SS 316L. Thus, SEM results indicate that Ta shows better macrophage adhesion in comparison with both TaeO and SS 316L, which suggests that macrophage adhesion is improved the more hydrophobic the surface: the more hydrophobic Ta surfaces showed better adhesion, as opposed to the more moderately hydrophobic TaeO and SS 316L surfaces.

Table 2 Average and RMS roughness of the coatings. Sample group

Roughness ± SD (nm)

Ta TaeO

Ra

Rq

20.9 ± 2.8 6.5 ± 2.5

26.1 ± 3.5 9.1 ± 4.8

and TaeO, from 1 mJ/m2 to 5 mJ/m2. Accordingly, Almeida Alves et al. reported an increase in the negative polar surface energy component γ− as a function of O content increase [14]. Lastly, it is known that surface wettability is a crucial physiochemical materials property, being linked to the regulation of adsorption/desorption of cell adhesion-mediating proteins such as fibronectin and vitronectin [14]. Moreover, it has been shown that optimal cell adhesion occurs for moderately hydrophilic and positively charged substrates, due to the adsorption of cell adhesion-mediating proteins in an advantageous geometrical conformation, whereas highly hydrophilic surfaces prevent the adsorption of proteins, and, for highly hydrophobic surfaces, proteins are adsorbed in rigid and denatured forms, thus hindering cell adhesion [41]. 3.6. Cell morphology Cell morphology after adhesion was observed by SEM following 24 h and 48 h of co-incubation for thin films deposited onto SS 316L (Fig. 6). For the Ta coatings, cells bundled into clusters with up to a few dozen micrometres in width after 24 h. Upon closer inspection, cell clusters seemed to form along the grain boundaries of the films, which constitute preferential loci for cell proliferation. While a majority of cells exhibited rounded morphology, adherent cells with better spread Table 3 Contact angles, surface energies, and surface free energies for Ta, TaeO, and SS 316L. Sample group

Ta TaeO SS 316L

ΔG (mJ/m2)

Surface energy (mJ/m2)

Contact angle ± SD (°) θW

θB

θG

γLW

γ+

γ−

106.0 ± 2.9 82.6 ± 2.9 75.7 ± 3.0

29.6 ± 3.1 22.3 ± 2.8 44.9 ± 3.0

100.3 ± 2.9 72.2 ± 1.9 68.5 ± 3.0

38.81 41.14 32.39

0.00 0.00 0.34

1.16 4.91 10.11

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Fig. 5. A) contact angles and B) wetting behaviour in view of Milli-Q ultrapure water for Ta, TaeO, and SS 316L. Differences among means were found to be significant between all sample groups (p ≤ 0.0001).

3.7. Cell viability and proliferation Cell viability was determined by the MTT colorimetric assay following 24 h and 48 h of co-incubation (Fig. 7). As seen in Fig. 7, the cell viability of the coatings was higher than 100%, both at 24 h and 48 h. After 24 h, Ta (145%) showed only slightly higher cell viability in comparison with TaeO (137%). After 48 h, however, Ta showed a cell viability increase of about 21% (175%), whereas TaeO showed a 15% decrease (116%). Differences among means between Ta and TaeO were found to be significant only at 48 h (p ≤ 0.05). Chang et al. reported similar trends regarding cell viability for Ta and Ta2O5 coatings co-incubated with human fibroblast cells, with asdeposited Ta2O5 having shown lower cell viability in comparison with Ta, whereas annealed β-Ta2O5 showed higher cell viability than both as-deposited Ta2O5 and Ta [12]. In summation, results indicate that the coatings are non-cytotoxic, with the high values obtained for metabolic cell viability suggesting that they are successful in promoting cell proliferation, as predicted in 3.6. It has been well established in the literature that cell proliferation is deeply influenced by the surrounding three-dimensional (3D) extracellular matrix (ECM) [42]. In a study performed with cell layers

Fig. 7. Cell viability following 24 h and 48 h of co-incubation. The cell viability of the coatings was determined according to Eq. 1 in relation to absorbance values obtained for SS 316L. Differences among means between Ta and TaeO were found to be significant only at 48 h (p ≤ 0.05).

Fig. 6. SEM micrographs of cell morphology after adhesion following 24 h and 48 h of co-incubation for Ta, TaeO, and SS 316L.

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(Fig. 8B) were quantified by the ELISA assay. LPS denotes LPS-treated cells (positive control of inflammation), whereas UC denotes untreated cells (negative control of inflammation). As seen in Fig. 8A, IL-6 levels were sensibly the same for Ta and TaeO, both at 24 h and 48 h (13 pg/mL), whereas SS 316L experienced a slight increase in IL-6 concentration from 24 h (13 pg/mL) to 48 h (15 pg/mL). Moreover, untreated cells showed only slightly lower IL-6 levels when compared with the previous sample groups, both at 24 h and 48 h (12 pg/mL). Differences among means were found to be significant between all sample groups and LPS-treated cells, both at 24 h and 48 h (p ≤ 0.0001). After 24 h, differences among means were found to be significant between Ta and untreated cells (p ≤ 0.01), and, to a lesser extent, between TaeO and untreated cells and SS 316L and untreated cells (p ≤ 0.05). After 48 h, differences among means were found to be significant between SS 316L and untreated cells (p ≤ 0.0001), TaeO and SS 316L (p ≤ 0.001), and Ta and SS 316L (p ≤ 0.01). As seen in Fig. 8B, Ta showed slightly higher IL-10 levels after 24 h (163 pg/mL) in comparison with SS 316L (158 pg/mL) and TaeO (156 pg/mL). After 48 h, a slight increase in IL-10 concentration was observed for all sample groups, and specially for TaeO (164 pg/mL). Unlike with IL-6, untreated cells showed slightly higher IL-10 levels when compared with the previous sample groups, both at 24 h (164 pg/ mL) and 48 h (166 pg/mL). Differences among means were found to be significant between all sample groups and LPS-treated cells, both at 24 h and 48 h (p ≤ 0.0001). In comparison with the IL-6 (165 pg/mL) and IL-10 (426 pg/mL) levels obtained for LPS-treated cells, it becomes clear that the coatings in no way induce an inflammatory response from macrophage cells. Lastly, results are in good accordance with the literature. Studies performed with murine macrophage-like J774.A1 cells co-incubated with nanostructured Ti surfaces surface-modified with Ca2 + and Sr2 + divalent cations, for example, showed similar trends in regards to the cytokines interleukin-1 beta (IL-1β), IL-10, and tumour necrosis factor alpha (TNFα) [46].

Table 4 A48/A24 ratio for Ta, TaeO, and SS 316L. Sample group

Ta TaeO SS 316L

A48/A24 ratio

Absorbance ± SD 24 h

48 h

2.33 ± 0.12 2.21 ± 0.17 1.65 ± 0.26

2.75 ± 0.05 2.00 ± 0.13 1.78 ± 0.56

1.18 0.91 1.08

embedded in dense 3D collagen matrices using the easy (one-step) 2,3bis-(2-methoxy-4-nitro-5-sulfophenyl)-2H-tetrazolium-5-carboxanilide (XTT) assay, Huyck et al. showed that the proliferation fold at a given time point x (tx) versus time zero (t0) can be easily calculated as the ratio between the net absorbances at tx versus t0 (Ax/A0), with Ax/A0 < 1 being indicative of cytotoxicity and Ax/A0 > 1 of cytopositive effect [42]. In the present paper, and as previously described in 2.3.2, cell viability was assessed by the MTT assay, with relative cell viability having been determined from absorbance measurements at 570 nm following 24 h and 48 h of co-incubation. Therefore, an estimate for cell proliferation can be easily obtained through the ratio between the average absorbances at 48 h and 24 h (A48/A24) (Table 4). As seen in Table 4, while TaeO showed negative cell proliferation (A48/A24 < 1), both Ta and SS 316L showed positive cell proliferation (A48/A24 > 1), thus indicating a cytopositive effect. That said, it must be underscored that even though TaeO showed slight cytotoxicity in terms of proliferation kinetics, its cell viability was well above 100% at both considered time points.

3.8. Inflammatory response Cytokines are small, non-structural proteins that play a key role in cell signalling, regulating host responses to infection, immune responses, inflammation, and trauma [43]. While some cytokines act to increase inflammation (pro-inflammatory cytokines), others act to reduce inflammation and promote healing (anti-inflammatory cytokines) [43]. IL-6, on the one hand, is an interleukin involved in inflammation and infection responses, and in the regulation of metabolic, regenerative, and neural processes, being secreted by T cells and macrophages in order to stimulate immune response during infection and after trauma [44]. IL-10, on the other hand, is a multifunctional cytokine that displays anti-inflammatory action—by inhibiting the activation and effector functions of T cells, monocytes, and macrophages—and whose principal routine function is to limit and ultimately terminate inflammatory responses, as well as regulate the growth and/ or differentiation of such cells as B cells, natural killer (NK) cells, cytotoxic/helper T cells, mast cells, and endothelial cells [45]. In order to evaluate the inflammatory response of the cells upon coincubation with the coatings, the cytokines IL-6 (Fig. 8A) and IL-10

4. Conclusion In order to evaluate such phenomena as cell activation, adhesion, and proliferation, Ta and TaeO films were deposited onto SS 316L substrates by reactive DC magnetron sputtering. Initially, the coatings were characterised in terms of their chemical composition, phase composition, morphology, topography, roughness, and wettability. In terms of phase composition, Ta showed a mixture of the α-Ta and β-Ta phases, whereas TaeO showed a nanocrystalline structure. In terms of morphology, Ta showed a columnar morphology, whereas TaeO showed a compact morphology. A decrease in roughness was observed between Ta and TaeO, from Ra of 21 nm to 7 nm. Likewise, a decrease in hydrophobicity was observed between Ta and TaeO, from θW of 106° to 83°. In vitro studies were performed with the murine macrophage-like Fig. 8. A) IL-6 and B) IL-10 concentrations following 24 h and 48 h of co-incubation, respectively. Both for IL-6 and IL-10, differences among means were found to be significant between all sample groups and LPS-treated cells, both at 24 h and 48 h (p ≤ 0.0001).

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cell line RAW 264.7. Cells exhibited better spread appearance, and thus better adhesion, when co-incubated with Ta, in comparison with both TaeO and SS 316L. In terms of cell viability, both Ta and TaeO were shown to be non-cytotoxic. However, Ta outperformed TaeO in terms of relative cell viability, both at 24 h and 48 h. Lastly, in terms of cytokine release, both Ta and TaeO showed the same levels of IL-6 after 24 h (13 pg/mL), whereas Ta showed slightly higher levels of IL-10 in comparison with TaeO after 24 h (163 pg/mL and 156 pg/mL, respectively). That said, in comparison with the levels of IL-6 and IL-10 secreted by cells treated with LPS, cytokine release was deemed not significant for either Ta or TaeO—prompting the conclusion that the coatings do not in any way induce an inflammatory response from macrophage cells. Acknowledgements This research was sponsored by FEDER funds, through the program COMPETE – Programa Operacional Factores de Competitividade; and by the Portuguese Foundation for Science and Technology (FCT), in the framework of the Strategic Funding UID/FIS/04650/2013, UID/EMS/ 00285/2013, and UID/BIA/04050/2013, as well as the ERA-SIINN/ 0004/2013 and PTDC/CTM-NAN/4242/2014 projects. References [1] M. Avila, D.M. Ojcius, Ö. Yilmaz, The oral microbiota: living with a permanent guest, DNA Cell Biol. 28 (2009) 405–411, http://dx.doi.org/10.1089/dna.2009. 0874. [2] M.F. Cesta, Normal structure, function, and histology of mucosa-associated lymphoid tissue, Toxicol. Pathol. 34 (2006) 599–608, http://dx.doi.org/10.1080/ 01926230600865531. [3] J. Raphel, M. Holodniy, S.B. Goodman, S.C. Heilshorn, Multifunctional coatings to simultaneously promote osseointegration and prevent infection of orthopaedic implants, Biomaterials 84 (2016) 301–314, http://dx.doi.org/10.1016/j. biomaterials.2016.01.016. [4] Q.-L. Ma, L.-Z. Zhao, R.-R. Liu, B.-Q. Jin, W. Song, Y. Wang, Y.-S. Zhang, L.-H. Chen, Y.-M. Zhang, Improved implant osseointegration of a nanostructured titanium surface via mediation of macrophage polarization, Biomaterials 35 (2014) 9853–9867, http://dx.doi.org/10.1016/j.biomaterials.2014.08.025. [5] E.F. Irwin, K. Saha, M. Rosenbluth, L.J. Gamble, D.G. Castner, K.E. Healy, Modulusdependent macrophage adhesion and behavior, J. Biomater. Sci. Polym. Ed. 19 (2008) 1363–1382, http://dx.doi.org/10.1163/156856208786052407. [6] F.A. Shah, M. Trobos, P. Thomsen, A. Palmquist, Commercially pure titanium (cpTi) versus titanium alloy (Ti6Al4V) materials as bone anchored implants — is one truly better than the other? Mater. Sci. Eng. C 62 (2016) 960–966, http://dx.doi. org/10.1016/j.msec.2016.01.032. [7] M. Sánchez-Siles, D. Muñoz-Cámara, N. Salazar-Sánchez, J.F. Ballester-Ferrandis, F. Camacho-Alonso, Incidence of peri-implantitis and oral quality of life in patients rehabilitated with implants with different neck designs: a 10-year retrospective study, J. Cranio-Maxillofacial Surg. 43 (2015) 2168–2174, http://dx.doi.org/10. 1016/j.jcms.2015.10.010. [8] G. Li, Q. Zhao, H. Yang, L. Cheng, Antibacterial and microstructure properties of titanium surfaces modified with ag-incorporated nanotube arrays, Mater. Res. 19 (2016) 735–740, http://dx.doi.org/10.1590/1980-5373-MR-2015-0534. [9] A.S. de S. Mello, P.L. dos Santos, A. Marquesi, T.P. Queiroz, R. Margonar, A.P. de Souza Faloni, Some aspects of bone remodeling around dental implants, Rev. Clínica Periodoncia Implantol. Y Rehabil. Oral. (2016) 1–9, http://dx.doi.org/10. 1016/j.piro.2015.12.001. [10] I. Atsuta, Y. Ayukawa, R. Kondo, W. Oshiro, Y. Matsuura, A. Furuhashi, Y. Tsukiyama, K. Koyano, Soft tissue sealing around dental implants based on histological interpretation, J. Prosthodont. Res. 60 (2016) 3–11, http://dx.doi.org/ 10.1016/j.jpor.2015.07.001. [11] B. Feng, J. Weng, B.C. Yang, S.X. Qu, X.D. Zhang, Characterization of surface oxide films on titanium and adhesion of osteoblast, Biomaterials 24 (2003) 4663–4670, http://dx.doi.org/10.1016/S0142-9612(03)00366-1. [12] Y.-Y. Chang, H.-L. Huang, H.-J. Chen, C.-H. Lai, C.-Y. Wen, Antibacterial properties and cytocompatibility of tantalum oxide coatings, Surf. Coat. Technol. 259 ( (2014) 193–198, http://dx.doi.org/10.1016/j.surfcoat.2014.03.061. [13] G. Xu, X. Shen, Y. Hu, P. Ma, K. Cai, Fabrication of tantalum oxide layers onto titanium substrates for improved corrosion resistance and cytocompatibility, Surf. Coat. Technol. 272 (2015) 58–65, http://dx.doi.org/10.1016/j.surfcoat.2015.04. 024. [14] C.F. Almeida Alves, A. Cavaleiro, S. Carvalho, Bioactivity response of Ta1-xOx coatings deposited by reactive DC magnetron sputtering, Mater. Sci. Eng. C 58 (2016) 110–118, http://dx.doi.org/10.1016/j.msec.2015.08.017. [15] K. McNamara, O. Kolaj-Robin, S. Belochapkine, F. Laffir, A.A. Gandhi, S.A.M. Tofail, Surface chemistry and cytotoxicity of reactively sputtered tantalum oxide films on NiTi plates, Thin Solid Films 589 (2015) 1–7, http://dx.doi.org/10.1016/j.tsf.2015.

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