Development of β Type Ti23Mo-45S5 Bioglass Nanocomposites for Dental Applications

Titanium β-type alloys attract attention as biomaterials for dental applications. The aim of this work was the synthesis of nanostructured β type Ti23Mo-x wt % 45S5 Bioglass (x = 0, 3 and 10) composites by mechanical alloying and powder metallurgy methods and their characterization. The crystallization of the amorphous material upon annealing led to the formation of a nanostructured β type Ti23Mo alloy with a grain size of approximately 40 nm. With the increase of the 45S5 Bioglass contents in Ti23Mo, nanocomposite increase of the α-phase is noticeable. The electrochemical treatment in phosphoric acid electrolyte resulted in a porous surface, followed by bioactive ceramic Ca-P deposition. Corrosion resistance potentiodynamic testing in Ringer solution at 37 °C showed a positive effect of porosity and Ca-P deposition on nanostructured Ti23Mo 3 wt % 45S5 Bioglass nanocomposite. The contact angles of glycerol on the nanostructured Ti23Mo alloy were determined and show visible decrease for bulk Ti23Mo 3 wt % 45S5 Bioglass and etched Ti23Mo 3 wt % 45S5 Bioglass nanocomposites. In vitro tests culture of normal human osteoblast cells showed very good cell proliferation, colonization, and multilayering. The present study demonstrated that porous Ti23Mo 3 wt % 45S5 Bioglass nanocomposite is a promising biomaterial for bone tissue engineering.


Introduction
Since 1965, titanium has found an increasing application as an implant material in medicine. Commercial purity titanium has high corrosion resistance and outstanding biocompatibility [1][2][3]. One reason for these advantages may be a protective oxide layer, which forms spontaneously on the implant surface [4]. Titanium exists in two allotropic forms. At low temperatures, it has a closed packed hexagonal crystal structure (hcp), which is commonly known as α, whereas above 882˘2˝C, it has a body centered cubic structure (bcc) termed β. The alloying elements such as Al, O, N, etc. tend to stabilize the α phase while elements V, Mo, Nb, Fe, Cr, etc. stabilize the β phase [1,3].
Titanium and its alloys also attract a lot of attention in dental applications [5][6][7]. Pure titanium and Ti-6Al-4V alloy are the main materials in the dental field as well as in the surgical one. Implant sensitivity to titanium alloys is very seldom, despite components such as vanadium which are described to be cytotoxic [7]. By the elimination of toxic elements, it is possible to prepare Ti-type alloys with excellent biocompatibility. Ti-6Al-7Nb, which has been developed for surgical implants, is also attractive for dental applications [8]. Recently, Ti-40Zr, Ti-5Al-13Ta and Ti-43.1Zr-10.2Al-3.6V investigated. Yet to the authors' knowledge, there have been no papers regarding the addition of 45S5 Bioglass into β Ti23Mo nanocomposite to have appeared until now.

Sample Preparation
Mechanical alloying (MA) was performed using SPEX 8000 Mixer Mill (SPEX SamplePrep, Metuchen, NJ, USA). Argon was a protective atmosphere. Round bottom stainless vials were used. Elemental powders (titanium, CAS number 7440-32-6 (<45 µm, 99.9%, Alfa Aesar), molybdenum, CAS number 7439-98-7 (44 µm, 99.6%, Sigma Aldrich) and 45S5 Bioglass (45% SiO 2 , 24.5% Na 2 O, 24.5% CaO, 6% P 2 O 5 , 53 µm; all in wt % from Mo-Sci Health Care L.L.C.)) were weighted, blended and poured into vials in a glove box (Labmaster 130) filled with automatically controlled argon atmosphere (O 2 < 2 ppm and H 2 O < 1 ppm). The weight ratio of hard steel balls (10 mm diameter) to powder weight ratio equaled 15:1. The MA process lasted 30 h in all cases. In order to prevent severe cold welding during high-energy milling, the ball milling was stopped every 2 h to dissipate the heat and to reduce an excessive rise in temperature. In the next step, the produced powders with size distribution of the particles from 40-150 µm were placed into the matrix and uniaxially pressed at a pressure of 500 MPa. Finally, the green compacts were heated over 1 h to 800˝C and kept at this temperature for 30 min for particle sintering. After that, the sinters were slowly cooled down to room temperature (RT) together with the furnace. The sintering was done at 10´2 Pa vacuum in an alumina tube (
The experiments were carried out on Ti23Mo alloy and two composite materials. For brevity, in this work, materials are denoted as follows: • Ti23Mo nanocrystalline alloy is labeled as Ti23Mo.

Sample Preparation
Mechanical alloying (MA) was performed using SPEX 8000 Mixer Mill (SPEX ® SamplePrep, Metuchen, NJ, USA). Argon was a protective atmosphere. Round bottom stainless vials were used. Elemental powders (titanium, CAS number 7440-32-6 (<45 μm, 99.9%, Alfa Aesar), molybdenum, CAS number 7439-98-7 (44 μm, 99.6%, Sigma Aldrich) and 45S5 Bioglass (45% SiO2, 24.5% Na2O, 24.5% CaO, 6% P2O5, 53 μm; all in wt % from Mo-Sci Health Care L.L.C.)) were weighted, blended and poured into vials in a glove box (Labmaster 130) filled with automatically controlled argon atmosphere (O2 < 2 ppm and H2O < 1 ppm). The weight ratio of hard steel balls (10 mm diameter) to powder weight ratio equaled 15:1. The MA process lasted 30 h in all cases. In order to prevent severe cold welding during high-energy milling, the ball milling was stopped every 2 h to dissipate the heat and to reduce an excessive rise in temperature. In the next step, the produced powders with size distribution of the particles from 40-150 μm were placed into the matrix and uniaxially pressed at a pressure of 500 MPa. Finally, the green compacts were heated over 1 h to 800 °C and kept at this temperature for 30 min for particle sintering. After that, the sinters were slowly cooled down to room temperature (RT) together with the furnace. The sintering was done at 10 −2 Pa vacuum in an alumina tube (McDanel Adv. Ceramic Technologies, Beaver Falls, PA, USA). The final bulk sinters have diameter and height of 8 and 4 mm, respectively ( Figure 1). The Solartron 1285 potentiostat (Solatron analytical, Farnborough, UK) was applied in the electrochemical etching stage of the sinters, followed at 10 V vs. open-circuit-potential (OCP) for 60 min in 1 M H3PO4 + 2% HF electrolyte. In the following stage, calcium phosphate was deposited on the Experimental setup Ti23Mo-x wt % 45S5 Bioglass nanocomposite synthesis and electrochemical treatment procedure.
The Solartron 1285 potentiostat (Solatron analytical, Farnborough, UK) was applied in the electrochemical etching stage of the sinters, followed at 10 V vs. open-circuit-potential (OCP) for 60 min in 1 M H 3 PO 4 + 2% HF electrolyte. In the following stage, calcium phosphate was deposited on the surface. We applied cathodic deposition at´5 V vs. OCP in the electrolyte containing 0.042 M Ca(NO 3 ) 2 + 0.025 M (NH 4 ) 2 HPO 4 + 0.1 M HCl for 1 h.

Material Characterization
The crystallographic structure of the samples during different processing stages was investigated at room temperature using a Panalytical Empyrean X-ray diffraction (XRD) with CuKα 1 (λ = 1.54056 Å) radiation (Panalytical, Empyrean model, Almelo, The Netherlands). The conditions of XRD measurements were: voltage 45 kV, anode current 40 mA, 2θ range 20˝-80˝, time per step 12.54 s/step, step size 0.0167˝. The average crystallite sizes d were estimated by Scherrer method: β = 0.9 λ/d cosθ, where β is the full-width at half maximum intensity of a Bragg reflection excluding instrumental broadening, θ the Bragg angle, and λ the wavelength of the X-ray radiation. Scanning electron microscope (SEM, VEGA 5135 Tescan, Brno, Czech Republic) with energy dispersive spectrometer (EDS, PTG Prison Avalon, Princeton Gamma Tech., Princeton, NY, USA) was used to characterize the chemical composition of the prepared samples. EDS was calibrated using a typical Cu calibration procedure. Additionally, the Ti (99.9% from Alfa Aesar), Mo (99.6% from Sigma Aldrich) and SiO 2 , Na 2 O, CaO, P 2 O 5 (all from Sigma Aldrich) specimens were used as reference.
The porosity of the porous materials was calculated by the formula p = (1´ρ/ρ th )ˆ100%, where ρ and ρ th are the density of the porous material and its corresponding theoretical density calculated for the rule of the mixtures, respectively. The density of the sintered samples was determined by Archimedes method. The Vickers microhardness of the bulk samples was measured using a microhardness tester by applying a load of 300 g for 10 s on the polished surfaces of the samples. For each sample, 10 separate indents were created on the investigated surface.
The surface morphology of the samples was investigated using an Optical Profiler NT 1100 (Veeco WykoR , Mannheim, Germany). Using the 3D optical profiler, the following standard roughness parameters were estimated: arithmetic mean roughness (µm)-R a , maximum height of the profile (µm)-R t , ten point mean roughness (µm)-R z and hybrid 3D parameters: root mean square (˝)-S dq , surface area ratio gradient (%)-S dr .
The Solartron 1285 potentiostat (Solatron analytical) was applied for corrosion measurements. The corrosion resistance of different samples was measured in Ringer's solution (simulated body fluid with composition: NaCl: 9 g/L, KCl: 0.42 g/L, CaCl 2 : 0.48 g/L, NaHCO 3 : 0.2 g/L) applying potentiodynamic mode with scan rate 0.5 mV/s at temperature of 37˘1˝C, controlled by thermostat. The corrosion test was run in EG & G K0047 corrosion cell. The counter electrode consisted of two graphite rods, and a platinum electrode (SCE, Hydromet, Gliwice, Poland) was used as the reference electrode. The surface area exposed to the electrolyte was 0.5 mm 2 . Polarization curves were obtained for each specimen in the potential range for´2 to 3 V. The corrosion potentials (E corr ) and corrosion current densities (I corr ) were estimated from the Tafel extrapolations of the corrosion curves, using CorrView software (Scribner Associates Inc., Southern Pines, NC, USA).
For the surface wettability measurements, contact angles Θ were determined by computer software analysis Musial (Elektronika Jadrowa, Krakow, Poland) from geometrical shape of droplets recorded by the optical system with a digital camera. Droplets placement were realized by a special micropipette with a constant volume of test liquid (glycerol (99.5% purity); 2 µdm 3 ) on a clear, polished and rinsed with alcohol surface, by it raise until the bottom touches the specimen. If the drop is large enough, the adhesion to the surface pulls it off from the tip. All the data were obtained under equilibrium conditions, procedure was repeated three times for all samples. For each recorded picture (i.e., for each liquid drop), the geometrical shape analysis was repeated 10 times: the extreme values were rejected, and the arithmetic mean value was calculated for the accepted findings.

In vitro Cytocompatibility
The in vitro cytocompatibility tests were performed under static conditions. The discs of Ti23Mo 3BG nanocomposite after electrochemical etching and additional Ca-P deposition were sterilized by autoclaving at 120˝C for 15 min and were separately located at the bottom of 24-well microplates. Normal Human Osteoblast (NHOst) cells from Cambrex (CC-2538, Walkersville, MD, USA) were cultured onto each disc at a concentration of 5000 cells/well in 1 mL of culture medium under static conditions. The cells were cultured at 37˝C in a 5% CO 2 incubator for 1 and 5 days. The medium was replaced every day. Then, the cells were fixed with a 25% glutaraldehyde solution for 10 min and stained with a 10% Giemsa's staining solution for 10 min. The specimens were sputter-coated with gold and examined using scanning electron microscope (SEM).  (Figure 2d). The high plastic deformations of the powders results in high density of dislocation lines and subsequently subgrain formation, which finally leads to amorphisation [24,25]. The low temperature sintering results in crystallization and nanostructure formation. microplates. Normal Human Osteoblast (NHOst) cells from Cambrex (CC-2538, Walkersville, MD, USA) were cultured onto each disc at a concentration of 5000 cells/well in 1 mL of culture medium under static conditions. The cells were cultured at 37 °C in a 5% CO2 incubator for 1 and 5 days. The medium was replaced every day. Then, the cells were fixed with a 25% glutaraldehyde solution for 10 min and stained with a 10% Giemsa's staining solution for 10 min. The specimens were sputter-coated with gold and examined using scanning electron microscope (SEM).  (Figure 2d). The high plastic deformations of the powders results in high density of dislocation lines and subsequently subgrain formation, which finally leads to amorphisation [24,25]. The low temperature sintering results in crystallization and nanostructure formation.  (Table 1). According to the Scherrer method of XRD profiles, the mean crystallite size of heat treated Ti23Mo alloy was 8 nm. Ti23Mo 3BG nanocomposite is mostly single phase cubic β-type phase material (reference code 01-077-3482) with some impurities as α-Ti, MoTi (reference code 01-071-9821) and SiC (reference code 01-075-834; Figure 2e). With the increase of the 45S5 Bioglass contents in Ti23Mo nanocomposite increase of α-phase is noticeable.

Structure Properties
The smooth and porous Ti23Mo alloy surface is presented in Figure 3a. The relative density of the bulk Ti23Mo 3BG nanocomposite was measured to be 91%. The results of EDS analysis of the surface of sintered Ti23Mo 3BG nanocomposite are shown in Figure 3d. It can be confirmed that synthesized nanocomposite mainly consists of Ti-Mo matrix with elements of O, Na, Si, Ca and P.  (Table 1). According to the Scherrer method of XRD profiles, the mean crystallite size of heat treated Ti23Mo alloy was 8 nm. Ti23Mo 3BG nanocomposite is mostly single phase cubic β-type phase material (reference code 01-077-3482) with some impurities as α-Ti, MoTi (reference code 01-071-9821) and SiC (reference code 01-075-834; Figure 2e). With the increase of the 45S5 Bioglass contents in Ti23Mo nanocomposite increase of α-phase is noticeable.
The smooth and porous Ti23Mo alloy surface is presented in Figure 3a. The relative density of the bulk Ti23Mo 3BG nanocomposite was measured to be 91%. The results of EDS analysis of the surface of sintered Ti23Mo 3BG nanocomposite are shown in Figure 3d. It can be confirmed that synthesized nanocomposite mainly consists of Ti-Mo matrix with elements of O, Na, Si, Ca and P. Table 1. Structural parameters (a, c, V), crystallite size (d), theoretical density (ρ th ) and microhardness (HV 0.3 ) for bulk Ti23Mo-x wt % Bioglass nanocomposites.

Sample
Phase  Electrochemical anodic oxidation results in pore formation on the surface of Ti23Mo 3BG nanocomposite. The porosity of bulk Ti23Mo 3BG nanocomposite is about 9%. In the etched surface large macropores with size up to 20 μm (Figure 4a) and smaller mesopores with size smaller than 0.5 μm are formed (not shown). The macropores are created at the places, where previously remnant pores existed as a result of the sintering, while the mesopores on the surface are the result of the large volume grain boundary etching. The prepared rough surface states a good base for tissue growth and their strong fixing or for intermediate bioceramic (Ca-P) layer formation.
Next the Ca-P deposition (Figure 4b) leads to intermediate bioactive layer formation. The electrolyte with Ca/P ratio equal to 1.67 was used. The obtained layers are porous and rough, which result in increased surface area and enlarged area for contacts with tissue. The Ca-P is deposited in pits and pores, growing into them, therefore fixating the layer in the metallic background. EDS analysis of the deposited Ca-P layer (Figure 4c) shows that Ca content is 35.00 at % and P content is Next the Ca-P deposition (Figure 4b) leads to intermediate bioactive layer formation. The electrolyte with Ca/P ratio equal to 1.67 was used. The obtained layers are porous and rough, which result in increased surface area and enlarged area for contacts with tissue. The Ca-P is deposited in pits and pores, growing into them, therefore fixating the layer in the metallic background. EDS analysis of the deposited Ca-P layer (Figure 4c) shows that Ca content is 35.00 at % and P content is 21.69 at %, which means that the Ca/P atomic ratio is equal to 1.61, which corresponds almost to the value of hydroxyapatite (reference code 00-024-0033). On the spectrum, the calcium, phosphorus, and oxygen peaks dominate.  Figure 5c. The sintered Ti23Mo 3BG nanocomposite has β-type structure with some trace of α-phase. On the XRD spectrum of Ti23Mo alloy the impurity is marked (Figure 5a). For etched Ti23Mo 3BG nanocomposite SiC phase is absent, which only confirms its surface character. In the anodized surface, besides pores, α-Ti phase is formed. XRD spectrum did not reveal strong peaks belonging to titanium oxides. It is likely during the electrochemical etching process, that thin amorphous titanium oxides were formed. After Ca-P deposition, the surface layer is composed of the apatite and small trace of MoTi phase.
The corrosion resistance of the different surfaces of the Ti23Mo 3BG composite in Ringer solution was investigated (Table 2, Figure 6). Using CorrView software (Solatron analytical,), from the Tafel extrapolations of the recorded potentiodynamic corrosion curves, corrosion current density (Icor) and corrosion potential (Ecor) were determined. The addition of 45S5 Bioglass to Ti23Mo alloy had also a positive effect on the corrosion resistance in Ringer's solution. Ti23Mo composite with 3 wt % of 45S5 Bioglass has better corrosion resistance (Icor = 64.44 μA/cm 2 , Ecor = −0.79 V) than nanostructured Ti23Mo alloy (Icor = 70.98 μA/cm 2 ). The best corrosion resistance is shown by Ti23Mo 3BG nanocomposite after electrochemical etching and Ca-P deposition (Icor = 0.0168 μA/cm 2 , Ecor = −0.44 V). After the electrochemical etching the surface is much rougher in comparison to the polished one. The oxides prepared in the anodic oxidation and Ca-P phase deposition from the electrolyte on the surface of Ti23Mo 3BG nanocomposite showed improved corrosion resistance. Additionally, the corrosion potential (Ecor) is shifted to the nobler direction (Ecor = −0.44 V).  Figure 5c. The sintered Ti23Mo 3BG nanocomposite has β-type structure with some trace of α-phase. On the XRD spectrum of Ti23Mo alloy the impurity is marked (Figure 5a). For etched Ti23Mo 3BG nanocomposite SiC phase is absent, which only confirms its surface character. In the anodized surface, besides pores, α-Ti phase is formed. XRD spectrum did not reveal strong peaks belonging to titanium oxides. It is likely during the electrochemical etching process, that thin amorphous titanium oxides were formed. After Ca-P deposition, the surface layer is composed of the apatite and small trace of MoTi phase.
The corrosion resistance of the different surfaces of the Ti23Mo 3BG composite in Ringer solution was investigated (Table 2, Figure 6). Using CorrView software (Solatron analytical), from the Tafel extrapolations of the recorded potentiodynamic corrosion curves, corrosion current density (I cor ) and corrosion potential (E cor ) were determined. The addition of 45S5 Bioglass to Ti23Mo alloy had also a positive effect on the corrosion resistance in Ringer's solution. Ti23Mo composite with 3 wt % of 45S5 Bioglass has better corrosion resistance (I cor = 64.44 µA/cm 2 , E cor =´0.79 V) than nanostructured Ti23Mo alloy (I cor = 70.98 µA/cm 2 ). The best corrosion resistance is shown by Ti23Mo 3BG nanocomposite after electrochemical etching and Ca-P deposition (I cor = 0.0168 µA/cm 2 , E cor =´0.44 V). After the electrochemical etching the surface is much rougher in comparison to the polished one. The oxides prepared in the anodic oxidation and Ca-P phase deposition from the electrolyte on the surface of Ti23Mo 3BG nanocomposite showed improved corrosion resistance. Additionally, the corrosion potential (E cor ) is shifted to the nobler direction (E cor =´0.44 V).

Surface Properties
Surface roughness is a considerably important property for the attachment of cells to the implant. It has been previously reported that not only micro-but also nano-topography can support the proliferation of different types of cells [12]. Optical profiler 3D topography and profiler surface scans and X-profiles of the bulk Ti23Mo 3BG nanocomposite, at the different processing stages, are presented in Figures 7 and 8, respectively.
At different processing stages changes occur in the surface porosity (Figure 7). Primarily the polished surface exhibited only few pores (Figure 7a), which resemble the remnants of the sintering process. Significant differences were observed after etching and additional Ca-P deposition (Figure 7b).

Surface Properties
Surface roughness is a considerably important property for the attachment of cells to the implant. It has been previously reported that not only micro-but also nano-topography can support the proliferation of different types of cells [12]. Optical profiler 3D topography and profiler surface scans and X-profiles of the bulk Ti23Mo 3BG nanocomposite, at the different processing stages, are presented in Figures 7 and 8, respectively.
At different processing stages changes occur in the surface porosity (Figure 7). Primarily the polished surface exhibited only few pores (Figure 7a), which resemble the remnants of the sintering process. Significant differences were observed after etching and additional Ca-P deposition (Figure 7b).

Surface Properties
Surface roughness is a considerably important property for the attachment of cells to the implant. It has been previously reported that not only micro-but also nano-topography can support the proliferation of different types of cells [12]. Optical profiler 3D topography and profiler surface scans and X-profiles of the bulk Ti23Mo 3BG nanocomposite, at the different processing stages, are presented in Figures 7 and 8 respectively. pore size for the cell attachment, differentiation, and ingrowth of osteoblasts and vascularization has been reported to be approximately 200-500 μm [26].
In Figure 8, the respective X-line profiles of the studied materials are shown. Large pores are formed in the surface during the etching process. This surface is able to support cell growth and proliferation, which provides a surface with large hillocks and a very non-uniform surface profile after culturing of the osteoblasts. The profile is an effect of the multilayered nature of growing cells.    In Figure 8, the respective X-line profiles of the studied materials are shown. Large pores are formed in the surface during the etching process. This surface is able to support cell growth and proliferation, which provides a surface with large hillocks and a very non-uniform surface profile after culturing of the osteoblasts. The profile is an effect of the multilayered nature of growing cells.   At different processing stages changes occur in the surface porosity (Figure 7). Primarily the polished surface exhibited only few pores (Figure 7a), which resemble the remnants of the sintering process. Significant differences were observed after etching and additional Ca-P deposition (Figure 7b). The increase of surface together with surface composition modification plays a key role for the living cell attachment and proliferation.
The bulk Ti23Mo 3BG nanocomposite had R a , R t and R z values of approximately 0.29, 11.29 and 8.24 µm, respectively (Table 3). In contrast, all of the roughness parameters increased for the etched and Ca-P deposited Ti23Mo 3BG nanocomposite surface (Figure 7b). This sample surface had an average surface roughness with R a , R t and R z values in the range of 35.80-276.55 µm. The optimal pore size for the cell attachment, differentiation, and ingrowth of osteoblasts and vascularization has been reported to be approximately 200-500 µm [26]. In Figure 8, the respective X-line profiles of the studied materials are shown. Large pores are formed in the surface during the etching process. This surface is able to support cell growth and proliferation, which provides a surface with large hillocks and a very non-uniform surface profile after culturing of the osteoblasts. The profile is an effect of the multilayered nature of growing cells. Table 3. 2D (R a , R t , R z ) and hybrid (S dq , S dr ) parameters for the Ti23Mo 3BG nanocomposite on different processing routes; parameters taken from surface area of 1.08 mm 2 . The hybrid parameters (S dq , S dr ,) reflect slope gradients and exhibit comparable behavior to the roughness parameters. The S dq value that affects the wetting of the surface by fluids significantly increases for the etched Ti23Mo 3BG sample. An increase in the surface slope (S dq ) may improve the fixing force of the cells, which ultimately provides a more stable connection between the implant and bone. Finally, S dr , which expresses the increase of the interfacial surface area relative to the area of the projected (flat) x, y plane, may improve the area for the attachment of cells. In our case, there is a considerable increase of S dr for the etched and Ca-P deposited Ti23Mo 3BG surface (255.66%) compared to the polished surface (1.22%).

Processing Route
As shown in Figure 9 surface wettability assay recorded a lower glycerol contact angle of Ti23Mo 3BG bulk nanocomposite (41.41˝˘0.94˝) than that of the nanocrystalline Ti23Mo alloy (85.64˝˘0.72˝). Etched nanocomposite exhibited the enhanced surface hydrophilicity (26.11˝˘1.62˝). Additional Ca-P deposition after electrochemical etching showed, that contact angle significantly increases (72.41˝˘0.56˝), which means that wettability decreases ( Table 2). The change of two or more surface characteristics at the same time, such as surface roughness and chemistry, whether deliberate or not, complicates the evaluation of the roles of the parameters on the wetting behavior and the biological performance [27].  Table 3. 2D (Ra, Rt, Rz) and hybrid (Sdq, Sdr) parameters for the Ti23Mo 3BG nanocomposite on different processing routes; parameters taken from surface area of 1.08 mm 2 .

Processing Route
Ra ( The hybrid parameters (Sdq, Sdr,) reflect slope gradients and exhibit comparable behavior to the roughness parameters. The Sdq value that affects the wetting of the surface by fluids significantly increases for the etched Ti23Mo 3BG sample. An increase in the surface slope (Sdq) may improve the fixing force of the cells, which ultimately provides a more stable connection between the implant and bone. Finally, Sdr, which expresses the increase of the interfacial surface area relative to the area of the projected (flat) x, y plane, may improve the area for the attachment of cells. In our case, there is a considerable increase of Sdr for the etched and Ca-P deposited Ti23Mo 3BG surface (255.66%) compared to the polished surface (1.22%).
As shown in Figure 9 surface wettability assay recorded a lower glycerol contact angle of Ti23Mo 3BG bulk nanocomposite (41.41° ± 0.94°) than that of the nanocrystalline Ti23Mo alloy (85.64° ± 0.72°). Etched nanocomposite exhibited the enhanced surface hydrophilicity (26.11° ± 1.62°). Additional Ca-P deposition after electrochemical etching showed, that contact angle significantly increases (72.41° ± 0.56°), which means that wettability decreases ( Table 2). The change of two or more surface characteristics at the same time, such as surface roughness and chemistry, whether deliberate or not, complicates the evaluation of the roles of the parameters on the wetting behavior and the biological performance [27].

In vitro Cytocompatibility
The culture was prepared on the Ti23Mo 3BG nanocomposite sample before electrochemical etching (Figure 10a-c), after electrochemical etching only (Figure 10d-f) and after etching with Ca-P deposition (Figure 10g-i). The collected data reveals a significant difference in the morphological characteristics of the cells on the porous and polished materials even after the first day of cell culturing ( Figure 10). The surface irregularities (protrusions, hillocks and pores) improve osteoblast adhesion, thus cells attach on the micro-and nano-surface irregularities. It has been proven that cell attachment and proliferation depend on the surface topography and roughness [28]. The osteoblasts that grew on the Ti23Mo 3BG nanocomposite sample before electrochemical etching exhibited adhesion to the material surface after one day and covered most of the surface after five days (Figure 10a-c). In the case of the electrochemically etched sample, after 24 h of incubation, cells

In vitro Cytocompatibility
The culture was prepared on the Ti23Mo 3BG nanocomposite sample before electrochemical etching (Figure 10a-c), after electrochemical etching only (Figure 10d-f) and after etching with Ca-P deposition (Figure 10g-i). The collected data reveals a significant difference in the morphological characteristics of the cells on the porous and polished materials even after the first day of cell culturing ( Figure 10). The surface irregularities (protrusions, hillocks and pores) improve osteoblast adhesion, thus cells attach on the micro-and nano-surface irregularities. It has been proven that cell attachment and proliferation depend on the surface topography and roughness [28]. The osteoblasts that grew on the Ti23Mo 3BG nanocomposite sample before electrochemical etching exhibited adhesion to the material surface after one day and covered most of the surface after five days (Figure 10a-c). In the case of the electrochemically etched sample, after 24 h of incubation, cells penetrated into the structure of the surface instead of forming filapodia (Figure 10d). After culturing the cells for five days, the cells on the etched Ti23Mo 3BG nanocomposite were observed to be well spread, both on the surface of the porous scaffold and inside the pores (Figure 10e,f).  On the other hand, in the case of the etched and Ca-P deposited Ti23Mo 3BG sample, after the first day of incubation, cells showed good adhesion to the surface of the studied sample in the form of filapodia (Figure 10g). A monolayer was formed on the surface of the sample after five days of incubation (Figure 10h,i). The covering of the implant surface with Ca-P layer is the final stage of the surface treatment.

Discussion
Nanotechnology is used for improving biomaterials or creating new biomaterials used for specific dental applications [29]. Nanostructured biomaterials can be personalized by engineering their structure, shape, size, and surface properties in order to be applied in precise anatomical sites. Based on the in vitro results, nanostructured metals undoubtedly have the potential to yield a faster and more stable integration of dental implants [29][30][31]. The in vivo effects of several nanoscale surface modification approaches and a potential applicability of these techniques to already commercialized dental implants are currently under investigation [28,29].
Current research focuses on improving the mechanical performance and biocompatibility of metal-based systems through changes in alloy composition, microstructure, and surface treatments [12,15,22,26,32,33]. In the case of titanium, a lot of attention is being paid to enhance the strength characteristics of commercial purity grades in order to avoid potential biotoxicity of alloying elements, especially in dental implants [34,35]. Improvement of the physicochemical and mechanical performance of Ti-based implant materials can be achieved through microstructure control, the top-down approach known as mechanical alloying technique [13][14][15][16][17]22,36,37]. Recent studies showed clearly that nanostructuring of titanium can considerably improve not only the mechanical properties, but also the biocompatibility [33]. Nanostructured materials can exhibit enhanced mechanical, biological, and chemical properties compared with their conventional counterparts [30,31,37].
While these biomaterials have been successful in encouraging bone ingrowth both in vivo and in clinical trials, the range of materials and microstructures available is still rather limited. To optimize On the other hand, in the case of the etched and Ca-P deposited Ti23Mo 3BG sample, after the first day of incubation, cells showed good adhesion to the surface of the studied sample in the form of filapodia (Figure 10g). A monolayer was formed on the surface of the sample after five days of incubation (Figure 10h,i). The covering of the implant surface with Ca-P layer is the final stage of the surface treatment.

Discussion
Nanotechnology is used for improving biomaterials or creating new biomaterials used for specific dental applications [29]. Nanostructured biomaterials can be personalized by engineering their structure, shape, size, and surface properties in order to be applied in precise anatomical sites. Based on the in vitro results, nanostructured metals undoubtedly have the potential to yield a faster and more stable integration of dental implants [29][30][31]. The in vivo effects of several nanoscale surface modification approaches and a potential applicability of these techniques to already commercialized dental implants are currently under investigation [28,29].
Current research focuses on improving the mechanical performance and biocompatibility of metal-based systems through changes in alloy composition, microstructure, and surface treatments [12,15,22,26,32,33]. In the case of titanium, a lot of attention is being paid to enhance the strength characteristics of commercial purity grades in order to avoid potential biotoxicity of alloying elements, especially in dental implants [34,35]. Improvement of the physicochemical and mechanical performance of Ti-based implant materials can be achieved through microstructure control, the top-down approach known as mechanical alloying technique [13][14][15][16][17]22,36,37]. Recent studies showed clearly that nanostructuring of titanium can considerably improve not only the mechanical properties, but also the biocompatibility [33]. Nanostructured materials can exhibit enhanced mechanical, biological, and chemical properties compared with their conventional counterparts [30,31,37].
While these biomaterials have been successful in encouraging bone ingrowth both in vivo and in clinical trials, the range of materials and microstructures available is still rather limited. To optimize Ti-based alloys for dental implant applications, several studies have focused on the design of biomaterials with optimized architecture to fulfill physico-chemical, mechanical as well as regeneration requirements. β-type titanium alloys are a class of interesting biomaterials that can exhibit a unique combination of physical, chemical and mechanical properties [38].
In this work, mechanical alloying was used to synthesize nanostructured Ti23Mo xBG nanocomposites. This technique enables alloying of elements that are difficult or impossible to combine by conventional melting methods (i.e., Ti and Bioglass). The nanocomposite surface improvement was achieved by electrochemical treatment. This process was composed of two stages: anodic oxidation and Ca-P deposition. The applied etching electrolyte was composed of a mixture of 1M H 3 PO 4 and 2% HF. Four stages of anodic oxidation were suggested [39]: (i) a compact oxide barrier layer formation; (ii) field-enhanced dissolution of the oxide layer; (iii) formation of porous structure in the slits and cracks (pore formation and dissolution of the oxide layer is possible as well) and (iv) the porous structure is consumed when the dissolution rate is larger than the pore formation rate.
It has been proven that the surface properties of the Ti-based implants determine successful osseointegration. The results published recently by Sul et al. showed that the morphology of different commercially available clinical titanium implants differed due to the surface modification techniques used during manufacture [40]. The bulk nanostructured Ti23Mo 3BG nanocomposite had R a , R t and R z values of around 0.29, 11.29 and 8.24 µm, respectively. The etched Ti23Mo 3BG composite had average surface roughnesses, with R a , R t and R z in the 6.42-42.47 µm range. After the Ca-P layer deposition a still relatively rough surface with Ca-P protrusions is well visible.
Implants, due to the corrosive environment of the tissue and body fluids, may undergo unexpected local corrosion attacks, leading to a release of the corrosion products to the tissue and its poisoning. The corrosion test results indicated that the nanostructured Ti23Mo alloy possesses a lower corrosion resistance and consequently a higher corrosion current density (I corr = 70.93 µA/cm 2 , E corr =´0.7925) in the Ringer's solutions. The bulk Ti23Mo 3BG nanocomposite has a better corrosion resistance (I corr = 0.0168 µA/cm 2 , E corr =´0.4423). For comparison, corrosion current density of microcrystalline Ti is 0.658 µA/cm 2 in Ringer's solution at 37˝C.
The surface wettability was confirmed to be a key parameter influencing the proliferation of osteoblasts [27,41]. Enhanced surface wettability would improve the cellular adhesion by optimizing the protein adsorption and providing a suitable surface presented to the filopodia. As a result, the electrochemically etched Ti23Mo 3BG nanocomposite, due to its good corrosion resistance and the surface wettability (contact angle Θ = 26.1˝), shows a high extent of cell proliferation.
In order to test the toxicity of materials, as well as interaction between the tested materials and the cells, in vitro cytocompatibility tests are conducted. The first results of increased osteoblast adhesion on the nanostructured metals was published by Webster and co-workers [12,26,30,31]. In this work, the in vitro cytocompatibility test was performed on the samples with different surfaces. The residual porosity after powder compaction played an important role in the cell adhesion. A significant difference in the morphological characteristics of the cells on the etched and polished materials was visible after one day of cell culturing. The cells tend to adhere with their entire surface to the porous sample penetrating the pores, whereas more spherical cells were seen with a smaller contact surface but more filapodia on the polished surface. After the first day, the cells strongly attached to the surface of the etched nanocomposite Ti23Mo 3BG sample, not producing filapodia or only producing them in a smaller amount compared to the cell culture on the smooth surface. After five days of cell culture, a smaller portion of the observed area was covered by cells on the etched Ti23Mo 3BG composite in comparison with the analogous bulk sample.
Bioactive glasses are biocompatible and exhibit a strong interfacial bonding with bone. Their bioactivity is attributed to the formation on their surface of a hydroxycarbonated apatite (HCA) layer. The rate of tissue bonding appears to depend on the rate of HCA formation, which follows a sequence of reactions between the implanted material and the surrounding tissues and physiologic fluids [42][43][44][45]. Cao and Hench proposed a three-step mechanism of HCA formation when the bioactive glass comes into contact with physiologic fluids: ion exchange, dissolution, and precipitation [42]. In the last case, precipitation of the calcium and phosphate ions released from the glass together with those from the solution, form a calcium-phosphate-rich layer (CaP) on the surface [44].
Modification at the nanoscale of the surface properties of dental implants can be achieved by various techniques, in order to create a more efficient implant integration with the bone [46]. Alterations in the cell shape and cytoskeleton, thus influencing specific gene expression, have been noticed after modifying the surface texture or roughness of the implants [46,47]. Moreover it has been demonstrated that anodized titanium surfaces enhances osteogenic activity in vitro [48,49]. On the other hand, oxidative nanopatterning confers titanium-based metals the ability to selectively guide cell behavior, facilitating the growth of cells having the osteoblastic potential [50]. Therefore these techniques are crucial for tissue regeneration, leading in the future to higher predictability in tissue healing around dental implants.

Conclusions
In this work, new kinds of biomedical Ti23Mo xBG nanocomposite were developed by the introduction of 45S5 Bioglass powders into the Ti23Mo matrix. Additionally the formation of a porous nanocrystalline Ti23Mo 3BG nanocomposite with an electrochemically modified surface was demonstrated. On the etched surface a bioactive HA layer was deposited. The following conclusions can be withdrawn: (1) etching improves porosity of the surface; (2) Ti23Mo 3BG nanocomposite after etching and Ca-P deposition had an average surface roughness in the 35.80-276.55 µm range; (3) the corrosion resistance significantly changes after the electrochemical treatment and Ca-P deposition; (4) noticeable difference in the morphology of the NHOst cells on porous and polished materials is observed in SEM, even after one day of cell culturing; (5) the rough, electrochemically biofunctionalized surface (porous with Ca-P layer) supports osteoblast cell growth and proliferation.
The key factors for the success of implant integration with the surrounding hard tissues are surface topography and chemical composition. Therefore the biofunctionalized nanocrystalline Ti23Mo 3BG composite may be an important step forward in the development of such a structure, which will support the process of osseointegration of the implant.