Deposition of Ultrathin Nano-Hydroxyapatite Films on Laser Micro-Textured Titanium Surfaces to Prepare a Multiscale Surface Topography for Improved Surface Wettability/Energy

materials Article Deposition of Ultrathin Nano-Hydroxyapatite Films on Laser Micro-Textured Titanium Surfaces to Prepare a Multiscale Surface Topography for Improved Surface Wettability/Energy Maria Surmeneva 1, Polina Nikityuk 1, Michael Hans 2and Roman Surmenev 1,* 1Department of Experimental Physics, National Research Tomsk Polytechnic University, Lenin Avenue 30, Tomsk 634029, Russia; [email protected] (M.S.); [email protected] (P.N.) 2Functional Materials, Materials Science Department, Saarland University, Saarbrücken 66123, Germany; [email protected] *Correspondence: [email protected] or [email protected]; Tel.: +7-903-953-09-69 Academic Editor: Mark T. Whittaker Received: 22 July 2016; Accepted: 14 October 2016; Published: 25 October 2016 Abstract: The primary aim of this study was to analyse the correlation between topographical features and chemical composition with the changes in wettability and the surface free energy of microstructured titanium (Ti) surfaces. Periodic microscale structures on the surface of Ti substrates were fabricated via direct laser interference patterning (DLIP). Radio-frequency magnetron sputter deposition of ultrathin nanostructured hydroxyapatite (HA) films was used to form an additional nanoscale grain morphology on the microscale-structured Ti surfaces to generate multiscale surface structures. The surface characteristics were evaluated using atomic force microscopy and contact angle and surface free energy measurements. The structure and phase composition of the HA films were investigated using X-ray diffraction. The HA-coated periodic microscale structured Ti substrates exhibited a significantly lower water contact angle and a larger surface free energy compared with the uncoated Ti substrates. Control over the wettability and surface free energy was achieved using Ti substrates structured via the DLIP technique followed by the deposition of a nanostructured HA coating, which resulted in the changes in surface chemistry and the formation of multiscale surface topography on the nanoand microscale. Keywords: biocompatible coating; hydroxyapatite; multiscale topography; rf magnetron sputtering; surface patterning 1. Introduction Currently, a large number of devices and implants are used in medicine [ 1 ]. Biomaterials in the form of implants (e.g., ligaments, vascular grafts, heart valves, intraocular lenses, and dental implants) and medical devices (e.g., pacemakers, biosensors, and artificial hearts) are extensively used to replace and/or restore the function of disturbed or deteriorated tissues or organs, and thus improve the quality of life and longevity of human beings [ 2 , 3 ]. The field of biomaterials has shown rapid growth in keeping up with the demands of an aging population [2,3]. Implants should not only be mechanically resistant, but should also be able to rapidly heal the host organism. When implanted into living tissue, all materials initiate a host response, and this represents the first steps of tissue repair [ 4 , 5 ]. Modern implant design is directed towards making use of this immune response to improve implant integration while preventing the perpetuation of the immune response, which leads to chronic inflammation, foreign body reactions, and thus loss of the intended function [ 4 , 5 ]. In addition to determining some of the deformation and strength characteristics of Materials 2016,9, 862; doi:10.3390/ma9110862 www.mdpi.com/journal/materials Materials 2016,9, 862 2 of 15 implants (such as the durability, elasticity, and shape stability), it is necessary to minimize the trauma caused by their use and to view the implantation with respect to healing a wound in order to achieve a fast and full postoperative rehabilitation of patients [ 5 ]. As a result, it is necessary to fulfil frequently inconsistent requirements with respect to the physical and mechanical properties of the materials used to manufacture the specified products [ 5 ]. The physicochemical properties of implants strongly depend on the method used to form the surface of the implants. Numerous techniques have been used to enhance the surface compatibility of tissue implants. Most of these methods involve multiple preparation procedures that incorporate coating and/or patterning steps. Direct laser interference patterning (DLIP) requires a single processing step and can be applied to a wide range of materials. The principle behind this method is based on the unique intensity pattern generated by interfering laser beams, which show periodicities in the (sub-)micron range [6–8]. The compatibility of the medical implant surface with biological tissues is achieved through the use of a biocompatible coating, which can be a layer of calcium phosphate (CaP). Biologically-relevant CaP belongs to the orthophosphate group, and naturally occurs in several biological structures, including teeth and bone [ 9 ]. Bone consists of an inorganic component of biological apatite as well as an organic component, which primarily consists of collagen and water. Currently, hydroxyapatite (HA), Ca 10 (PO 4 ) 6 (OH) 2 , is an optimum material for clinical practice, with a similar structure to that of the mineral component of bone tissue [ 9 ]. The basic technological methods that are used to prepare biocompatible coatings include plasma spraying, ion-beam deposition, radio-frequency (rf) magnetron sputtering, and electrochemical deposition. Rf magnetron sputtering is a prospective technique for the fabrication of implant coatings, because it can be used to form HA films that have low roughness and exhibit good adhesion to Ti substrate [ 10 , 11 ]. The plasma parameters of the rf magnetron sputtering process affect the physicochemical and mechanical properties of the CaP films [ 12 ]. The deposition parameters can be adjusted to produce single-phase HA films generated at a high deposition rate and high thermal stability. Moreover, the chemical composition of the precursor material used to deposit the HA coating is preserved [ 13 ]. The biocompatibility of HA has been thoroughly investigated and established, and it has been shown to promote the proliferation and differentiation of mesenchymal stem cells and adhesion of human keratinocyte cell lines. In addition, the improved adhesion, proliferation, and differentiation, the increased alkaline phosphatase activity of primary human osteoblast cells, and the normal cell growth of human embryonic kidney cell lines in the presence of HA have been shown experimentally [14]. The living tissue and the artificial implant interact at the molecular level, and the size of the biological structures ranges from a few nanometres to tens of micrometres. The scientific experience gained to date has demonstrated that the successful interaction between the biotissue and the surface of the implant frequently depends on physico-chemical material properties, such as the chemical composition, microstructure, roughness, wetting angle, and free surface energy (FSE) [ 15 ]. Thus, the aim of this study was to achieve the surface structuring of Ti using the DLIP technique followed by HA coating deposition through rf magnetron sputtering with pure HA. Analyses of the correlation between variations in the topographical multiscale Ti surface feature, including Ti coated with HA, and changes in the surface wettability and energy are reported. 2. Materials and Methods 2.1. Sample Preparation Technically pure Ti was used as a substrate. A high-powered pulsed Nd:YAG laser (Quanta-Ray PRO210, Spectra Physics, Santa Clara, CA, USA) was used for laser interference patterning. The repetition rate and pulse duration of the laser were 10 Hz and 10 ns, respectively. The fundamental wavelength of the Nd:YAG laser system was 1064 nm, and shorter wavelengths were obtained through second-harmonic generation. Samples with a surface area of 20 × 20 mm2 were irradiated at 355 nm with multiple adjacent 1 × 1–2 × 2 mm 2 spots at a fluency of 3.05 ± 0.15 J · cm 2 . Line-like structures Materials 2016,9, 862 3 of 15 with periodicities of approximately 4.5 µ m and 8.4 µ m were obtained using a two-beam laser setup. For cross-like structure types with the same periodicity, samples were structured once, rotated 90 ◦ , and then structured a second time. 2.2. Coating Deposition A commercially available apparatus with a rf magnetron source (13.56 MHz) was used to deposit the nanostructured HA coating [ 12 , 13 ]. The HA coating was deposited at an operating pressure of 0.4 Pa (the vacuum chamber was evacuated to 10 −4 Pa) at a target–substrate distance of 40 mm, with argon as the working gas, and with an rf generator power of 500 W. The HA coating was deposited for 8 h onto a substrate mounted on a grounded substrate holder, which resulted in a coating thickness of 650 ± 50 nm. A target of pure HA—synthesized by the mechanochemical method—was prepared according to the previously described procedures [ 16 , 17 ]. The target for rf magnetron sputtering (220 mm diameter, 10 mm thick) was prepared via ceramic technology—i.e., the powder was pressed at a pressure of 70 MPa and then annealed at 1100 ◦C for 1 h in air. 2.3. Atomic Force Microscopy (AFM) Measurements The quantitative analysis of the surface morphology of the uncoated and HA-coated substrates was performed with a Solver P47-PRO (NT-MDT, Moscow, Russia) atomic force microscope (AFM) using triangular golden silicon probes (NT-NDT) with a typical spring constant of 28 N ·m−1 and a resonance frequency of 420 kHz. All of the images were collected in contact AFM mode in air at a typical frequency of 1.5 Hz with an image resolution of 256 points per line. Squares of different sizes (5 × 5 µm2 and 35 × 35 µm2 ) were scanned, and the Nova SPM software (NT-MDT, Moscow, Russia) was used to analyse the surface roughness. Three different 3D parameters were used to characterize the surface roughness: (S a ), which is the arithmetic mean of the absolute values of the surface departure from the mean plane in the samples area; the root mean square roughness (S q ), which is an index used to represent the standard deviation of the surface heights; and S dr , which is the developed interfacial area ratio. 2.4. X-ray Diffraction (XRD) An X-ray diffractometer (Shimadzu XRD-6000, Tokyo, Japan) was used to identify the crystalline structure of the HA-coated Ti that were previously treated to prepare parallel and crossed grooves with different periodicities. The typical irradiation conditions were 40 kV and 30 mA using Cu-K α radiation (1.5405 Å); the 2 θ scan ranged from 10 ◦ to 60 ◦ with a step size of 0.02 ◦ at a speed of 2 deg/min and a grazing angle of 3 ◦ . The average crystallite size was determined using Scherrer’s equation from the broadening of the diffraction peaks; this determination was performed with the Powder Cell 2.4 software (FIMRT, Berlin, Germany), and the instrumental broadening was considered. An instrumental broadening of 0.1 ◦ in 2 θ was determined by the full width at half maximum (FWHM) of a silicon powder. 2.5. Contact Angle and Surface Free Energy Measurements Contact angle analyses were performed using an optical contact angle apparatus (OCA15 Plus Data Physics Instruments GmbH, Filderstadt, Germany) along with the SCA20 software (Data Physics Instruments GmbH, Filderstadt, Germany). The contact angle (CA) of water in air was measured using a sessile drop method. A minimum of 10 droplets (2 µ L, 5 µ L ·s−1 ) of water and three droplets of diiodomethane or ethylene glycol were seeded on the surface of each sample. The surface free energy was calculated using the Owens–Wendt–Rabel–Kaelble (ORWK) method. Three different media (water, diiodomethane, and ethylene glycol) were used for the calculations, and all the measurements were performed according to the study [ 18 ]. The phenomenon of contact angle hysteresis was also observed in the optical contact angle apparatus using a sessile liquid (water) droplet. After the liquid has advanced over a previously unwetted surface (i.e., when the solid/liquid contact area increases), Materials 2016,9, 862 4 of 15 the maximum contact angle at the three-phase contact line is referred to as the advancing angle ( θa ). A minimum contact angle is measured at the contact line when the liquid is retracted over a previously wetted surface (i.e., when the contact area shrinks); this is referred to as a receding angle ( θr ). Contact angle hysteresis is defined as the difference between the (maximum) advancing and (minimum) receding angles: ∆θhyst =θa−θr[19]. 3. Results and Discussion 3.1. Surface Roughness and Morphology To study the changes in the morphology, AFM analysis was carried out on the uncoated and HA-coated patterned Ti substrates. The results presented in Figure 1reveal that deposition of the HA coating resulted in changes in microscale surface roughness of the initial structured Ti substrate. In general, after HA coating deposition, the roughness parameters S a and S q decreased. For the grooves with a periodicity of 8.4 µm , the average surface roughness after deposition of the HA coating was lower than that of the structured Ti substrate. This phenomenon may be related to the growth mechanism of the HA coating. During the growth, the HA coating tends to fill the grooves, which results in a surface smoothening effect. Materials 2016, 9, 862 4 of 16 is retracted over a previously wetted surface (i.e., when the contact area shrinks); this is referred to as a receding angle (θ). Contact angle hysteresis is defined as the difference between the (maximum) advancing and (minimum) receding angles: Δθ =θ −θ [19]. 3. Results and Discussion 3.1. Surface Roughness and Morphology To study the changes in the morphology, AFM analysis was carried out on the uncoated and HA-coated patterned Ti substrates. The results presented in Figure 1 reveal that deposition of the HA coating resulted in changes in microscale surface roughness of the initial structured Ti substrate. In general, after HA coating deposition, the roughness parameters Sa and Sq decreased. For the grooves with a periodicity of 8.4 μm, the average surface roughness after deposition of the HA coating was lower than that of the structured Ti substrate. This phenomenon may be related to the growth mechanism of the HA coating. During the growth, the HA coating tends to fill the grooves, which results in a surface smoothening effect. (a) (b) Figure 1. The roughness parameters of different Ti surface patterns with a 35 × 35 μm scan area: ■—without hydroxyapatite (HA) film, and ○—with HA film. (a) Sa roughness parameter; (b) Sq roughness parameter Figures 2–5 show typical AFM images of the HA coating prepared using rf magnetron sputtering on the laser micro-textured Ti surfaces. The microstructure of the surfaces is easily observed using 35 × 35 μmscan areas. For the 5 × 5 µm2 scan area, clear grains of the HA coating with definite boundaries can be observed (Figures 2d and 3d). The deposited coating was homogenous and revealed a regular grain-like morphology, which is typical for a thin film deposited by rf magnetron sputtering [20,21]. The surface had a grain-like morphology with a grain size from 0.36 to 0.73 µm in the 5 × 5 μm scan area (Figures 2d and 3d). Therefore, the surface of the treated Ti substrates exhibited a multiscale structure. The surface topography on the microscale, sub-microscale, and nanoscale changed after the HA coating deposition, and this was observed in the AFM profile. Scans obtained for the 35 × 35 µm2 scan area showed that changes in the surface topography occurred after the HA coating deposition. During deposition, the film presumably fills the grooves of the Ti substrate. The AFM profile analysis of the HA-coated Ti shows a surface decorated with nanoscale grains (Figures 2d and 3d). The average grain height was approximately 30 nm, which resulted in an increase in the nanoscale roughness of the surface. The most significant changes of the surface topography occurred in surface structured Ti with parallel grooves. For a 5 × 5 µm2 scan area, the HA coating resulted in significant shrinkage in the as-prepared grooves, and the structure of the grooves was modified (Figure 2b,d). As shown in Figure 2b, separate grooves with parallel periodicity can be characterized with a full width at half-maximum (FWHM) of ~2.2 µm and an average groove depth of ~500 nm. After the HA Figure 1. The roughness parameters of different Ti surface patterns with a 35 × 35 µm2 scan area:  —without hydroxyapatite (HA) film, and # —with HA film. ( a )S a roughness parameter; (b)Sqroughness parameter. Figures 2–5show typical AFM images of the HA coating prepared using rf magnetron sputtering on the laser micro-textured Ti surfaces. The microstructure of the surfaces is easily observed using 35 × 35 µm2 scan areas. For the 5 × 5 µ m 2 scan area, clear grains of the HA coating with definite boundaries can be observed (Figures 2d and 3d). The deposited coating was homogenous and revealed a regular grain-like morphology, which is typical for a thin film deposited by rf magnetron sputtering [ 20 , 21 ]. The surface had a grain-like morphology with a grain size from 0.36 to 0.73 µ m in the 5 × 5 µm2 scan area (Figures 2d and 3d). Therefore, the surface of the treated Ti substrates exhibited a multiscale structure. The surface topography on the microscale, sub-microscale, and nanoscale changed after the HA coating deposition, and this was observed in the AFM profile. Scans obtained for the 35 × 35 µ m 2 scan area showed that changes in the surface topography occurred after the HA coating deposition. During deposition, the film presumably fills the grooves of the Ti substrate. The AFM profile analysis of the HA-coated Ti shows a surface decorated with nanoscale grains (Figures 2d and 3d). The average grain height was approximately 30 nm, which resulted in an increase in the nanoscale roughness of the surface. The most significant changes of the surface topography occurred in surface structured Materials 2016,9, 862 5 of 15 Ti with parallel grooves. For a 5 × 5 µ m 2 scan area, the HA coating resulted in significant shrinkage in the as-prepared grooves, and the structure of the grooves was modified (Figure 2b,d). As shown in Figure 2b, separate grooves with parallel periodicity can be characterized with a full width at half-maximum (FWHM) of ~2.2 µ m and an average groove depth of ~500 nm. After the HA coating deposition, the FWHM decreased to ~1.8 µ m and the average groove depth reduced to ~200 nm (Figure 2d). Materials 2016, 9, 862 5 of 16 coating deposition, the FWHM decreased to ~1.8 µm and the average groove depth reduced to ~200 nm (Figure 2d). Figure 2. Atomic force microscopy (AFM) images of the surface topographies of Ti samples with parallel grooves and a periodicity of approximately 4.5 µm. (a) 35 × 35 μm  ; (b) 5 × 5 μm  without a HA film; (c) 35 × 35 μm  ; and (d) 5 × 5 μm  with a HA film. Figure 2. Atomic force microscopy (AFM) images of the surface topographies of Ti samples with parallel grooves and a periodicity of approximately 4.5 µ m. ( a ) 35 × 35 µm2 ; ( b ) 5 × 5 µm2 without a HA film; (c) 35 ×35 µm2; and (d) 5 ×5µm2with a HA film. Materials 2016,9, 862 6 of 15 Materials 2016, 9, 862 6 of 16 Figure 3. AFM images of the topographies of Ti samples with a crossed periodicity of approximately 4.5 µm. (a) 35 × 35 μm  ; (b) 5 × 5 μm  without a HA film; (c) 35 × 35 μm  ; and (d) 5 × 5 μm  with a HA film. Figure 3. AFM images of the topographies of Ti samples with a crossed periodicity of approximately 4.5 µ m. ( a ) 35 × 35 µm2 ; ( b ) 5 × 5 µm2 without a HA film; ( c ) 35 × 35 µm2 ; and ( d ) 5 × 5 µm2 with a HA film. Materials 2016,9, 862 7 of 15 Materials 2016, 9, 862 7 of 16 Changes in the surface topography can be observed for the crossed patterns on the Ti surface. Moreover, after deposition of the HA coating, the cross-pattern topography could no longer be clearly observed, as the grooves were filled with the coating material. For the Ti samples with parallel grooves with a periodicity of 8.4 µm, a lower amount of the coating material filled the grooves (Figure 4a,b). For the Ti samples with crossed grooves with a periodicity of approximately 8.4 µm, the changes in the surface topography were more pronounced than in the grooves with a periodicity of approximately 4.5 µm (Figure 5a,b). The crossed patterns on the Ti surface were still observed after the deposition process. Figure 4. AFM images of the topographies of Ti samples with parallel grooves with a periodicity of approximately 8.4 µm. (a) 35 × 35 μm; (b) 35 × 35 μm with a HA film. Figure 5. AFM images of the topographies of Ti samples with crossed grooves with a periodicity of approximately 8.4 µm. (a) 35 × 35 μm; (b) 35 × 35 μm with a HA film. Topographical features on material surfaces are important to cell and tissue response to biomaterials [22–24]. Micro-rough surfaces may stimulate greater bone contact with the material by promoting the production of local osteogenic factors and the expression of differentiation markers. However, the role of nanometer roughness has not been clearly defined [22–24]. Previously published results showed that the topography of micro-rough Ti surfaces can significantly influence the attachment and growth of cells [22]. Multidirectional groove designs have been shown to contribute to improved MC3T3-E1 cell adhesion in different directions [24]. Larger Figure 4. AFM images of the topographies of Ti samples with parallel grooves with a periodicity of approximately 8.4 µm. (a) 35 ×35 µm2; (b) 35 ×35 µm2with a HA film. Materials 2016, 9, 862 7 of 16 Changes in the surface topography can be observed for the crossed patterns on the Ti surface. Moreover, after deposition of the HA coating, the cross-pattern topography could no longer be clearly observed, as the grooves were filled with the coating material. For the Ti samples with parallel grooves with a periodicity of 8.4 µm, a lower amount of the coating material filled the grooves (Figure 4a,b). For the Ti samples with crossed grooves with a periodicity of approximately 8.4 µm, the changes in the surface topography were more pronounced than in the grooves with a periodicity of approximately 4.5 µm (Figure 5a,b). The crossed patterns on the Ti surface were still observed after the deposition process. Figure 4. AFM images of the topographies of Ti samples with parallel grooves with a periodicity of approximately 8.4 µm. (a) 35 × 35 μm; (b) 35 × 35 μm with a HA film. Figure 5. AFM images of the topographies of Ti samples with crossed grooves with a periodicity of approximately 8.4 µm. (a) 35 × 35 μm; (b) 35 × 35 μm with a HA film. Topographical features on material surfaces are important to cell and tissue response to biomaterials [22–24]. Micro-rough surfaces may stimulate greater bone contact with the material by promoting the production of local osteogenic factors and the expression of differentiation markers. However, the role of nanometer roughness has not been clearly defined [22–24]. Previously published results showed that the topography of micro-rough Ti surfaces can significantly influence the attachment and growth of cells [22]. Multidirectional groove designs have been shown to contribute to improved MC3T3-E1 cell adhesion in different directions [24]. Larger Figure 5. AFM images of the topographies of Ti samples with crossed grooves with a periodicity of approximately 8.4 µm. (a) 35 ×35 µm2; (b) 35 ×35 µm2with a HA film. Changes in the surface topography can be observed for the crossed patterns on the Ti surface. Moreover, after deposition of the HA coating, the cross-pattern topography could no longer be clearly observed, as the grooves were filled with the coating material. For the Ti samples with parallel grooves with a periodicity of 8.4 µm, a lower amount of the coating material filled the grooves (Figure 4a,b). For the Ti samples with crossed grooves with a periodicity of approximately 8.4 µ m, the changes in the surface topography were more pronounced than in the grooves with a periodicity of approximately 4.5 µ m (Figure 5a,b). The crossed patterns on the Ti surface were still observed after the deposition process. Topographical features on material surfaces are important to cell and tissue response to biomaterials [ 22 – 24 ]. Micro-rough surfaces may stimulate greater bone contact with the material by promoting the production of local osteogenic factors and the expression of differentiation markers. However, the role of nanometer roughness has not been clearly defined [22–24]. Materials 2016,9, 862 8 of 15 Previously published results showed that the topography of micro-rough Ti surfaces can significantly influence the attachment and growth of cells [ 22 ]. Multidirectional groove designs have been shown to contribute to improved MC3T3-E1 cell adhesion in different directions [ 24 ]. Larger microgrooves (greater than ~1 µm in depth) have been proposed to increase plaque uptake in dental applications. Ulerich et al. investigated the effect of multiscale laser texturing of a Ti6Al4V substrate on the adhesion of osteoblasts [ 25 ]. The surface chemical composition can influence cell attachment and reaction to a metal by modifying the adsorption of proteins or by activating different cellular pathways of nearby cells [ 25 ]. Unique laser-induced structures can modify the morphology and local chemistry of the surface, which makes it more favourable for cells to grow in certain patterns or to grow at a particular density, depending on the features at various size scales [26]. 3.2. Phase and Structure Characterization The typical XRD patterns of HA-coated patterned Ti substrates with crossed grooves and periodicities of 4.5 µ m and 8.4 µ m are presented in Figure 6. Only the reflexes attributed to the HA coating are observed, which confirms that the HA film is crystalline and phase-pure. The preferential (002) crystallographic orientation is observed, which is typical for an rf magnetron sputter-deposited thin film [ 27 ]. Moreover, the crystallographic orientations of the deposited HA coating can be controlled to either (002) or (300) by modulating the deposition parameters, the thickness of the HA film, and the movement of the substrate during the deposition process [ 12 ]. The crystallite size and microstrain determined for the HA coating were 34 ± 1 nm and 0.06%, respectively. In the pattern, only the peaks attributed to the hexagonal α -Ti phase were identified. The lattice parameters of the Ti substrate were determined to be b=a= 2.9458 and c= 4.677 Å. The lattice parameters of the HA coating prepared on the Ti surface were determined as b=a= 9.4125 and c= 6.9167 Å. The crystallite size and microstrain evaluated for the Ti substrate were 42 ± 2 nm and 0.05%, respectively. No significant effect of the Ti surface, which was patterned by the DLIP technique, was observed on the phase composition of either the Ti substrates or the HA-coated structured Ti substrates. Thus, according to the XRD results, no surface TiO x (x < 2) layer was formed through surface processing by the DLIP technique. It has also been reported elsewhere that DLIP processing of the Ti surface does not result in the formation of an oxide layer [ 7 , 28 ]. This is the primary difference between surface processing of Ti by DLIP and conventional heat treatment of Ti in dry air, where Ti is oxidized [29]. Materials 2016, 9, 862 8 of 16 microgrooves (greater than ~1 µm in depth) have been proposed to increase plaque uptake in dental applications. Ulerich et al. investigated the effect of multiscale laser texturing of a Ti6Al4V substrate on the adhesion of osteoblasts [25]. The surface chemical composition can influence cell attachment and reaction to a metal by modifying the adsorption of proteins or by activating different cellular pathways of nearby cells [25]. Unique laser-induced structures can modify the morphology and local chemistry of the surface, which makes it more favourable for cells to grow in certain patterns or to grow at a particular density, depending on the features at various size scales [26]. 3.2. Phase and Structure Characterization The typical XRD patterns of HA-coated patterned Ti substrates with crossed grooves and periodicities of 4.5 µm and 8.4 µm are presented in Figure 6. Only the reflexes attributed to the HA coating are observed, which confirms that the HA film is crystalline and phase-pure. The preferential (002) crystallographic orientation is observed, which is typical for an rf magnetron sputter-deposited thin film [27]. Moreover, the crystallographic orientations of the deposited HA coating can be controlled to either (002) or (300) by modulating the deposition parameters, the thickness of the HA film, and the movement of the substrate during the deposition process [12]. The crystallite size and microstrain determined for the HA coating were 34 ± 1 nm and 0.06%, respectively. In the pattern, only the peaks attributed to the hexagonal α-Ti phase were identified. The lattice parameters of the Ti substrate were determined to be b = a = 2.9458 and c = 4.677 Å. The lattice parameters of the HA coating prepared on the Ti surface were determined as b = a = 9.4125 and c = 6.9167 Å. The crystallite size and microstrain evaluated for the Ti substrate were 42 ± 2 nm and 0.05%, respectively. No significant effect of the Ti surface, which was patterned by the DLIP technique, was observed on the phase composition of either the Ti substrates or the HA-coated structured Ti substrates. Thus, according to the XRD results, no surface TiO x (x < 2) layer was formed through surface processing by the DLIP technique. It has also been reported elsewhere that DLIP processing of the Ti surface does not result in the formation of an oxide layer [7,28]. This is the primary difference between surface processing of Ti by DLIP and conventional heat treatment of Ti in dry air, where Ti is oxidized [29]. Figure 6. A typical X-ray diffraction (XRD) pattern of HA-coated structured Ti with crossed grooves and a periodicity of 4.5 µm. Figure 6. A typical X-ray diffraction (XRD) pattern of HA-coated structured Ti with crossed grooves and a periodicity of 4.5 µm. Materials 2016,9, 862 9 of 15 3.3. Wettability and Surface Free Energy The evolution of the contact angle and water hysteresis for droplets on the laser-treated uncoated (  ) and HA-coated ( # ) Ti surfaces is presented in Figure 7. The contact angle for the crossed grooves with a periodicity of approximately 8.4 µ m is larger than that for the parallel grooves with a periodicity of approximately 8.4 µm for both the HA-coated and uncoated substrates. Materials 2016, 9, 862 9 of 16 3.3. Wettability and Surface Free Energy The evolution of the contact angle and water hysteresis for droplets on the laser-treated uncoated (■) and HA-coated (○) Ti surfaces is presented in Figure 7. The contact angle for the crossed grooves with a periodicity of approximately 8.4 µm is larger than that for the parallel grooves with a periodicity of approximately 8.4 µm for both the HA-coated and uncoated substrates. The uncoated Ti surface with parallel grooves and a periodicity of approximately 4.5 µm resulted in a high static water CA of 99° ± 2°, which is indicative of surface hydrophobicity. The average water CA for the Ti substrate after the deposition of the HA film was measured as 75° ± 4°, which is associated with surface hydrophilicity. In Figure 6, the water CA hysteresis was determined to be 96.4° ± 1.9° and 67.4° ± 3.4° for the uncoated and HA-coated Ti surfaces, respectively. The surface wettability of artificial materials is one of the most important factors that determines cell adhesion. Tamada et al. claimed that a surface with a water contact angle of 70° represents the most suitable surface for cell adhesion [30]. In our study, the water contact angle was close to 70° after the deposition of HA. Therefore, the coating can provide beneficial effects for cell adhesion compared with uncoated Ti surfaces. Surface nanotextures—which provide increased surface area and finer surface roughness—may result in improved mechanical interlocking between the tissue and implant [31]. 5101520 70 75 80 85 90 95 100 105 Patterns 4,5 parallel 4,5 crossed 8,4 parallel 8,4 crossed Contact angle, î (a) 5101520 55 60 65 70 75 80 85 90 95 100 105 Patterns 4,5 parallel 4,5 crossed 8,4 parallel 8,4 crossed Water Hysteresis, î (b) Figure 7. The contact angle (a) and water hysteresis; (b) measurements for different patterns on the surface of Ti: ■—without HA film, and ○—with HA film. Figure 7. The contact angle ( a ) and water hysteresis; ( b ) measurements for different patterns on the surface of Ti: —without HA film, and #—with HA film. The uncoated Ti surface with parallel grooves and a periodicity of approximately 4.5 µ m resulted in a high static water CA of 99 ◦± 2 ◦ , which is indicative of surface hydrophobicity. The average water CA for the Ti substrate after the deposition of the HA film was measured as 75 ◦± 4 ◦ , which is associated with surface hydrophilicity. In Figure 6, the water CA hysteresis was determined to be 96.4 ◦± 1.9 ◦ and 67.4 ◦± 3.4 ◦ for the uncoated and HA-coated Ti surfaces, respectively. The surface wettability of artificial materials is one of the most important factors that determines cell adhesion. Tamada et al. claimed that a surface with a water contact angle of 70 ◦ represents the most suitable surface for cell adhesion [ 30 ]. In our study, the water contact angle was close to 70 ◦ after the deposition of HA. Therefore, the coating can provide beneficial effects for cell adhesion compared with uncoated Ti surfaces. Surface nanotextures—which provide increased surface area and finer surface roughness—may result in improved mechanical interlocking between the tissue and implant [31]. The surface free energy σ calculations using the CA data indicated that HA-coated Ti surfaces with parallel grooves and a periodicity of approximately 8.4 µ m had a significantly larger surface