- Email: [email protected]
Plasma electrolytic oxidation of a Ti–15Mo alloy in silicate solutions
Materials Letters 100 (2013) 252–256
Contents lists available at SciVerse ScienceDirect
Materials Letters journal homepage: www.elsevier.com/locate/matlet
Plasma electrolytic oxidation of a Ti–15Mo alloy in silicate solutions Dorota Babilas a, Katarzyna Służalska b, Agnieszka Krząkała a, Artur Maciej a, Robert P. Socha c, Grzegorz Dercz d, Grzegorz Tylko b, Joanna Michalska e, Anna M. Osyczka b, Wojciech Simka a,n a
Faculty of Chemistry, Silesian University of Technology, Gliwice, Poland Faculty of Biology and Earth Sciences, Jagiellonian University, Krakow, Poland c Institute of Catalysis and Surface Chemistry, PAS, Krakow, Poland d Institute of Materials Science, University of Silesia, Katowice, Poland e Faculty of Materials Science and Metallurgy, Silesian University of Technology, Katowice, Poland b
art ic l e i nf o
a b s t r a c t
Article history: Received 29 November 2012 Accepted 11 March 2013 Available online 20 March 2013
The influence of the parameters used in the plasma electrolytic oxidation of Ti–15Mo alloy on the microstructure and chemical composition of the obtained oxide layer was determined, and the bioactivity of the modified alloy was measured. The obtained oxide layers are composed of crystalline titanium oxides and incorporated silicon species. & 2013 Elsevier B.V. All rights reserved.
Keywords: Vanadium-free alloy Ti–15Mo Plasma electrolytic oxidation Bioactivity
1. Introduction Ti-6Al-4V is one of the most popular alloys used for hip and knee joint manufacturing [1]. Despite the high corrosion resistance and good mechanical properties, its use is limited due to its chemical composition, which includes aluminium [2]. This element, in the form of Al2O3, affects biological processes. Excessive amounts of Al2O3 in human organisms cause muscle pain, bone softening, increased susceptibility to broken bones, and concentration and memory aberrations. Higher concentrations of vanadium in the body cause damage to the respiratory, nervous, and digestive systems [2]. Recently, attention has been given to singlephase alloys of titanium, β-phase titanium alloys, which are more biocompatible [3,4]. The addition of molybdenum, niobium, or zirconium improves the mechanical properties of the alloys. The Ti–15Mo alloy, in comparison with α þβ phase alloys (Ti–6Al–4V and Ti–6Al–7Nb), exhibits much better corrosion and wear resistance and improved mechanical properties [3]. Additionally, the high electrochemical stability [5] and high biocompatibility of the Ti–15Mo alloy make it a promising material for orthopaedic implant production. The lifetime of titanium implants depends on the quality of integration with the bone [6]. To improve this property, the surface layer of the implant should be modified [7]. To improve the bioactivity of the surface layer, bioactive elements such as silicon are introduced to the coatings by plasma electrolytic oxidation
n
Corresponding author. E-mail address: [email protected] (W. Simka).
0167-577X/$ - see front matter & 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.matlet.2013.03.047
(PEO) [6,8,9]. Silicon is responsible for the growth and proper development of bones and connective tissues. In this work, we present for the first time the results of Ti–15Mo alloy modification by plasma electrolytic oxidation in a bath containing potassium silicate.
2. Materials and methods The composition of the Ti–15Mo (BIMO Metals, Wrocław, Poland) alloy used in this investigation was as follows: 14.73– 14.98 wt% Mo, 0.016 wt% N, 0.06 wt% Fe, 0.008 wt% C, 0.001 wt% H, 0.15 wt% O, and the balance Ti. The samples were anodised in a bath that contained potassium hydroxide (KOH—5 g dm−3) and potassium silicate (K2SiO3). The sample labels and treatment conditions are presented in Table 1. The methods of sample pretreatment and anodisation were described in our previous work [10]. The morphology and chemical composition of the anodic layer formed on the surface were examined using a scanning electron microscope (SEM, Hitachi S-3400N, accelerating voltage¼ 25 kV) and an energy-dispersive X-ray spectrometer (EDX, Thermo Noran). The X-ray photoelectron spectroscopy (XPS) measurements were performed in an ultrahigh vacuum (3 10–10 mbar) system equipped with a hemispherical analyser (SES R 4000, Gammadata Scienta) [11]. The X-ray diffraction experiments were carried out using an X'Pert Philips PW 3040/60 diffractometer. The GIXD diffraction patterns were registered in the 2θ range from 101 to 1001 with a 0.051 step for the incident angle α of 0.151 [10]. The roughness (Ra parameter) of the samples was measured using a Mitutoyo Surftest SJ-301 profilometer [10].
D. Babilas et al. / Materials Letters 100 (2013) 252–256
After anodic passivation, the samples of Ti–15Mo alloy were submitted for preliminary biocompatibility investigations with human bone marrow-derived mesenchymal stem cells (hMSCs). Cells were grown on alloy samples and evaluated for their viability (MTS test), alkaline phosphatase (ALP) activity, extracellular matrix and total collagen production. The details regarding these biocompatibility investigations can be found in our previous paper [10].
3. Results and discussion Fig. 1 shows the morphology of the surface of Ti–15Mo alloy samples anodised in different conditions. At voltages of 100 and 200 V and independent of the K2SiO3 concentration in the bath, Table 1 The sample labels, process conditions, and its influence on the roughness factor Ra, and Si/Ti atomic ratio; current density – 0.1 A cm−1; time – 5 min. Sample
K2SiO3 (mol dm−3)
Voltage (V)
Ra (μm)
Si/Ti
TM-ANO-0.5-100 TM-ANO-0.5-200 TM-ANO-0.5-350
0.5
100 200 350
0.36 2.53 2.44
0.01 0.01 11.88
TM-ANO-1.0-100 TM-ANO-1.0-200 TM-ANO-1.0-350
1.0
100 200 350
0.33 0.44 3.19
0.01 0.01 64.14
TM-ANO-0.5-100
253
the surfaces of the obtained samples were practically the same and were unchanged in comparison to the surfaces of Ti–15Mo alloy samples that were only etched. This indicates that a very thin oxide layer formed, which is typical for the oxidation of titanium alloys at low voltages [11]. The roughness factor of those samples was comparable and varied from 0.33 to 0.44 μm, which was close to that of the etched samples (0.58 μm). The atomic ratio of Si/Ti calculated from the EDX analysis was 0.01 (Table 1). An increase of the voltage to 350 V resulted in the formation of a more porous oxide layer containing a significant amount of silicon. The obtained oxide layer was typical for the PEO of metals in silicate solutions [11]. With an increase of the K2SiO3 concentration in the bath, the amount of silicon in the obtained oxide layer increased. The atomic ratio of Si/Ti in the TM-ANO-350-0.5 and TM-ANO-350-1.0 samples was 11.88 and 64.14 (Table 1), respectively. Additionally, the roughness of the Ti–15Mo alloy surface significantly increased up to 2.44 and 3.19 μm for the TM-ANO-350-0.5 and TM-ANO-350-1.0 samples (Table 1), respectively. The PEO layer formed at 350 V was composed of crystalline titanium oxides – anatase and rutile (Fig. 2) – and incorporated silicon species. Analysis of the O 1s core excitation (Fig. 2) showed four components, where the most intensive one (B) could be assigned to oxygen in silica or silicate compounds [12]. The A component could be related to oxygen bonded to a metal ion (Ti and/or Mo) and/or hydroxyl groups. The high BE values of the C and D peaks suggested that they could be assigned to organic species adsorbed from air onto the surface. The Si 2p spectrum
TM-ANO-1.0-100
10 µm
TM-ANO-0.5-200
10 µm
TM-ANO-1.0-200
10 µm
10 µm
TM-ANO-0.5-350
TM-ANO-1.0-350
10 µm Fig. 1. The SEM images of the Ti–15Mo alloy after anodising.
10 µm
254
D. Babilas et al. / Materials Letters 100 (2013) 252–256
titanium rutile anatase
O 1s
Si 2p
Fig. 2. The TL-XRD pattern and deconvoluted spectra of O 1s, Si 2p, for the Ti–15Mo alloy anodised at 350 V in 1.0 mol dm−3 K2SiO3 (TM-ANO-1.0-350 sample).
(Fig. 2) was deconvoluted into three components. The most intensive one (B), at a BE of 103.7 eV, was typical for crystalline silica, whereas the component C could be ascribed to silicon in hydroxysilicate structures. Component A exhibited a very low intensity, which indicated silicate species. In the case of the Ti 2p spectrum, three doublet components were fitted to the spectrum. The most intensive component (A) was assigned to metallic titanium in an alloy, and the component (B), at a BE of 455.3 eV, was assigned to a Ti–O bond but ascribed to the formal Ti þ oxidation state. The presence of component C at a BE of 460.0 eV indicated the formation of the titanium silicate species [13,14], for which the BE were observed in the range 459.3–460.1 eV. The metallic component in the Ti 2p spectrum suggested that the oxide layer thickness approached 12 nm or that the layer was thicker than 12 nm but partially cracked. The Mo 3d spectrum showed a very low intensity, which prevented the estimation of the electronic state of molybdenum. As presented in Fig. 3a, bone marrow-derived hMSCs were viable on the examined Ti–15Mo alloy samples after 10 and 21 days of culture. Anodisation in the bath containing K2SiO3 decreased the viability of hMSCs at day 10 of the culture (TMANO-0.5-350). However, at longer culture times, the viability of hMSCs increased in these samples. This suggests that this initial decrease in the viable cell number was temporary and that the remaining cells continued to proliferate. The cells that remained viable on TM-ANO-0.5-350 alloys also displayed increased ALP activity (Fig. 3b), although this increase was not significant compared with other studied growth surfaces. Collagen production by the hMSCs increased on samples anodised in either a 0.5 or 1.0 M solution of K2SiO3 (Fig. 3c). However, the level of mineralisation of the extracellular matrix produced by the hMSCs was slightly decreased for ANO-350-1.0 samples compared with the
two other studied surfaces (Fig. 3d). In Fig. 3, SEM images of the hMSCs on a representative etched sample and on TM-ANO-0.5350 and TM-ANO-1.0-350 samples are presented. The cells on all studied surfaces exhibited typical fibroblastic morphology, though, as mentioned above, there were differences in their number and distribution on the studied surfaces. The cells cultured on the TMANO-0.5-350 sample were relatively large, whereas cells on the TM-ANO-350-0.1 sample appeared smaller and less spread out. Martins Junior et al. reported the lowered viability of MC3TC cells on a Ti–15Mo alloy [15]. However, they did not observe changes in cell morphology or cytotoxic effects [15]. Our results further support these previous reports. The applied modification depends on the introduction of silicates into the surface layer, which influences the viability of the hMSCs in a negative way. However, the morphology of the cells is not affected, and the cells exhibit a typical fibroblastic shape. Munir et al. noted that covering the metallic materials with silicon particles favours apatite crystallisation, which later improves the osteogenic cell response of HOB [16]. According to the obtained results, anodising the Ti–15Mo alloy in a 0.5 M K2SiO3 solution can lead to increased bioactivity of the alloy. This may lead to increased ALP activity of the hMSCs, favouring generation of the collagen-rich extracellular matrix by these cells and extracellular matrix calcification.
4. Conclusions In this work, the influence of PEO parameters on the morphology and chemical and phase composition of a Ti–15Mo alloy surface is presented. It was found that the increase in the applied voltage caused the formation of a silicon-enriched porous layer.
cell viability % of change compared to control = 100%
D. Babilas et al. / Materials Letters 100 (2013) 252–256
255
120
TM sample 100 80 60 40 20 0 TM
TM-ANO-0.5-350
TM-ANO-1.0-350
ALP per cell % of change compared to control = 100%
160 140 120 100 80
TM-ANO-0.5-350 sample
60 40 20 0 TM
TM-ANO-0.5-350
TM-ANO-1.0-350
Collagen production per cell % of change compared to control= 100%
600
500
400
300
200
100
TM-ANO-1.0-350 sample
0 TM
TM-ANO-0.5-350
TM-ANO-1.0-350
TM
TM-ANO-0.5-350
TM-ANO-1.0-350
Mineralization per cell % of change conpared to control = 100%
120
100
80
60
40
20
0
at day 10 culture
at day 20 culture
Fig. 3. Cell viability of human MSC after 10- and 21-day cultures; ALP activity of human MSC at day 10 cultures; collagen production, and extracellular matrix calcification of human MSC cultures after 21-day cultures on control and modified Ti–15Mo alloy samples.
This modification also led to an increase of bioactivity of the Ti–15Mo alloy. References [1] Kuphasuk C, Oshida Y, Andres CJ, Hovijitra ST, Barco MT, Brown DT. J Prost Dent 2001;85:195–202.
[2] Donato TAG, de Almeida LH, Nogueira RA, Neimeyer TC, Grandini CR, Caram R, et al. Mater Sci Eng C 2009;29:1365–9. [3] Kumar S, Narayan TSNS. J Dent 2008;36:500–7. [4] Oliveira NTC, Alexio G, Caram R, Guastaldi AC. Mater Sci Eng A 2007;452– 453:727–31. [5] Oliveira NTC, Guastaldi AC. Acta Biomater 2009;5:399–405. [6] Hu H, Liu X, Ding C. Surf Coat Technol 2010;204:3265–71. [7] Liu X, Chu PK, Ding C. Mater Sci Eng R 2004;47:49–121.
256
D. Babilas et al. / Materials Letters 100 (2013) 252–256
[8] Krupa D, Badszkiewicz J, Zdunek J, Smolik J, Słomka Z, Sobczak JW. Surf Coat Technol 2010;205:1743–9. [9] Zhang W, Du K, Yan C, Wang F. Appl Surf Sci 2008;254:5216–23. [10] Krząkała A, Służalska K, Widziołek M, Szade J, Winiarski A, Dercz G, et al. Electrochim Acta 2012. http://dx.doi.org/10.1016/j.electacta.2012.07.075. [11] Simka W, Sowa M, Socha RP, Maciej A, Michalska J. Electrochim Acta 2012. http://dx.doi.org/10.1016/j.electacta.2012.10.130. [12] NIST X-ray photoelectron spectroscopy database; 〈http://srdata.nist.gov/xps/〉.
[13] Blasco T, Camblor MA, Corma A, Perez-Pariente JJ. Am Chem Soc 1993;115:11806–13. [14] Dementjev AP, Ivanova OP, Vasilyev LA, Naumkin AV, Nemirovsky DM, Shalaev DYJ. Vac Sci Technol A 1994;12:423–7. [15] Martins Júnior JRS, Nogueira RA, Araújo RO, Donato TAG, Arana-Chavez VE, A. P.R.A. Claro, et al. Mater Res 2011;14:107–12. [16] Munir G, Koller G, Di Silvio L, Edirisinghe MJ, Bonfield W, Huang JJR. Soc Interface 2011;8:678–88.
Contents lists available at SciVerse ScienceDirect
Materials Letters journal homepage: www.elsevier.com/locate/matlet
Plasma electrolytic oxidation of a Ti–15Mo alloy in silicate solutions Dorota Babilas a, Katarzyna Służalska b, Agnieszka Krząkała a, Artur Maciej a, Robert P. Socha c, Grzegorz Dercz d, Grzegorz Tylko b, Joanna Michalska e, Anna M. Osyczka b, Wojciech Simka a,n a
Faculty of Chemistry, Silesian University of Technology, Gliwice, Poland Faculty of Biology and Earth Sciences, Jagiellonian University, Krakow, Poland c Institute of Catalysis and Surface Chemistry, PAS, Krakow, Poland d Institute of Materials Science, University of Silesia, Katowice, Poland e Faculty of Materials Science and Metallurgy, Silesian University of Technology, Katowice, Poland b
art ic l e i nf o
a b s t r a c t
Article history: Received 29 November 2012 Accepted 11 March 2013 Available online 20 March 2013
The influence of the parameters used in the plasma electrolytic oxidation of Ti–15Mo alloy on the microstructure and chemical composition of the obtained oxide layer was determined, and the bioactivity of the modified alloy was measured. The obtained oxide layers are composed of crystalline titanium oxides and incorporated silicon species. & 2013 Elsevier B.V. All rights reserved.
Keywords: Vanadium-free alloy Ti–15Mo Plasma electrolytic oxidation Bioactivity
1. Introduction Ti-6Al-4V is one of the most popular alloys used for hip and knee joint manufacturing [1]. Despite the high corrosion resistance and good mechanical properties, its use is limited due to its chemical composition, which includes aluminium [2]. This element, in the form of Al2O3, affects biological processes. Excessive amounts of Al2O3 in human organisms cause muscle pain, bone softening, increased susceptibility to broken bones, and concentration and memory aberrations. Higher concentrations of vanadium in the body cause damage to the respiratory, nervous, and digestive systems [2]. Recently, attention has been given to singlephase alloys of titanium, β-phase titanium alloys, which are more biocompatible [3,4]. The addition of molybdenum, niobium, or zirconium improves the mechanical properties of the alloys. The Ti–15Mo alloy, in comparison with α þβ phase alloys (Ti–6Al–4V and Ti–6Al–7Nb), exhibits much better corrosion and wear resistance and improved mechanical properties [3]. Additionally, the high electrochemical stability [5] and high biocompatibility of the Ti–15Mo alloy make it a promising material for orthopaedic implant production. The lifetime of titanium implants depends on the quality of integration with the bone [6]. To improve this property, the surface layer of the implant should be modified [7]. To improve the bioactivity of the surface layer, bioactive elements such as silicon are introduced to the coatings by plasma electrolytic oxidation
n
Corresponding author. E-mail address: [email protected] (W. Simka).
0167-577X/$ - see front matter & 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.matlet.2013.03.047
(PEO) [6,8,9]. Silicon is responsible for the growth and proper development of bones and connective tissues. In this work, we present for the first time the results of Ti–15Mo alloy modification by plasma electrolytic oxidation in a bath containing potassium silicate.
2. Materials and methods The composition of the Ti–15Mo (BIMO Metals, Wrocław, Poland) alloy used in this investigation was as follows: 14.73– 14.98 wt% Mo, 0.016 wt% N, 0.06 wt% Fe, 0.008 wt% C, 0.001 wt% H, 0.15 wt% O, and the balance Ti. The samples were anodised in a bath that contained potassium hydroxide (KOH—5 g dm−3) and potassium silicate (K2SiO3). The sample labels and treatment conditions are presented in Table 1. The methods of sample pretreatment and anodisation were described in our previous work [10]. The morphology and chemical composition of the anodic layer formed on the surface were examined using a scanning electron microscope (SEM, Hitachi S-3400N, accelerating voltage¼ 25 kV) and an energy-dispersive X-ray spectrometer (EDX, Thermo Noran). The X-ray photoelectron spectroscopy (XPS) measurements were performed in an ultrahigh vacuum (3 10–10 mbar) system equipped with a hemispherical analyser (SES R 4000, Gammadata Scienta) [11]. The X-ray diffraction experiments were carried out using an X'Pert Philips PW 3040/60 diffractometer. The GIXD diffraction patterns were registered in the 2θ range from 101 to 1001 with a 0.051 step for the incident angle α of 0.151 [10]. The roughness (Ra parameter) of the samples was measured using a Mitutoyo Surftest SJ-301 profilometer [10].
D. Babilas et al. / Materials Letters 100 (2013) 252–256
After anodic passivation, the samples of Ti–15Mo alloy were submitted for preliminary biocompatibility investigations with human bone marrow-derived mesenchymal stem cells (hMSCs). Cells were grown on alloy samples and evaluated for their viability (MTS test), alkaline phosphatase (ALP) activity, extracellular matrix and total collagen production. The details regarding these biocompatibility investigations can be found in our previous paper [10].
3. Results and discussion Fig. 1 shows the morphology of the surface of Ti–15Mo alloy samples anodised in different conditions. At voltages of 100 and 200 V and independent of the K2SiO3 concentration in the bath, Table 1 The sample labels, process conditions, and its influence on the roughness factor Ra, and Si/Ti atomic ratio; current density – 0.1 A cm−1; time – 5 min. Sample
K2SiO3 (mol dm−3)
Voltage (V)
Ra (μm)
Si/Ti
TM-ANO-0.5-100 TM-ANO-0.5-200 TM-ANO-0.5-350
0.5
100 200 350
0.36 2.53 2.44
0.01 0.01 11.88
TM-ANO-1.0-100 TM-ANO-1.0-200 TM-ANO-1.0-350
1.0
100 200 350
0.33 0.44 3.19
0.01 0.01 64.14
TM-ANO-0.5-100
253
the surfaces of the obtained samples were practically the same and were unchanged in comparison to the surfaces of Ti–15Mo alloy samples that were only etched. This indicates that a very thin oxide layer formed, which is typical for the oxidation of titanium alloys at low voltages [11]. The roughness factor of those samples was comparable and varied from 0.33 to 0.44 μm, which was close to that of the etched samples (0.58 μm). The atomic ratio of Si/Ti calculated from the EDX analysis was 0.01 (Table 1). An increase of the voltage to 350 V resulted in the formation of a more porous oxide layer containing a significant amount of silicon. The obtained oxide layer was typical for the PEO of metals in silicate solutions [11]. With an increase of the K2SiO3 concentration in the bath, the amount of silicon in the obtained oxide layer increased. The atomic ratio of Si/Ti in the TM-ANO-350-0.5 and TM-ANO-350-1.0 samples was 11.88 and 64.14 (Table 1), respectively. Additionally, the roughness of the Ti–15Mo alloy surface significantly increased up to 2.44 and 3.19 μm for the TM-ANO-350-0.5 and TM-ANO-350-1.0 samples (Table 1), respectively. The PEO layer formed at 350 V was composed of crystalline titanium oxides – anatase and rutile (Fig. 2) – and incorporated silicon species. Analysis of the O 1s core excitation (Fig. 2) showed four components, where the most intensive one (B) could be assigned to oxygen in silica or silicate compounds [12]. The A component could be related to oxygen bonded to a metal ion (Ti and/or Mo) and/or hydroxyl groups. The high BE values of the C and D peaks suggested that they could be assigned to organic species adsorbed from air onto the surface. The Si 2p spectrum
TM-ANO-1.0-100
10 µm
TM-ANO-0.5-200
10 µm
TM-ANO-1.0-200
10 µm
10 µm
TM-ANO-0.5-350
TM-ANO-1.0-350
10 µm Fig. 1. The SEM images of the Ti–15Mo alloy after anodising.
10 µm
254
D. Babilas et al. / Materials Letters 100 (2013) 252–256
titanium rutile anatase
O 1s
Si 2p
Fig. 2. The TL-XRD pattern and deconvoluted spectra of O 1s, Si 2p, for the Ti–15Mo alloy anodised at 350 V in 1.0 mol dm−3 K2SiO3 (TM-ANO-1.0-350 sample).
(Fig. 2) was deconvoluted into three components. The most intensive one (B), at a BE of 103.7 eV, was typical for crystalline silica, whereas the component C could be ascribed to silicon in hydroxysilicate structures. Component A exhibited a very low intensity, which indicated silicate species. In the case of the Ti 2p spectrum, three doublet components were fitted to the spectrum. The most intensive component (A) was assigned to metallic titanium in an alloy, and the component (B), at a BE of 455.3 eV, was assigned to a Ti–O bond but ascribed to the formal Ti þ oxidation state. The presence of component C at a BE of 460.0 eV indicated the formation of the titanium silicate species [13,14], for which the BE were observed in the range 459.3–460.1 eV. The metallic component in the Ti 2p spectrum suggested that the oxide layer thickness approached 12 nm or that the layer was thicker than 12 nm but partially cracked. The Mo 3d spectrum showed a very low intensity, which prevented the estimation of the electronic state of molybdenum. As presented in Fig. 3a, bone marrow-derived hMSCs were viable on the examined Ti–15Mo alloy samples after 10 and 21 days of culture. Anodisation in the bath containing K2SiO3 decreased the viability of hMSCs at day 10 of the culture (TMANO-0.5-350). However, at longer culture times, the viability of hMSCs increased in these samples. This suggests that this initial decrease in the viable cell number was temporary and that the remaining cells continued to proliferate. The cells that remained viable on TM-ANO-0.5-350 alloys also displayed increased ALP activity (Fig. 3b), although this increase was not significant compared with other studied growth surfaces. Collagen production by the hMSCs increased on samples anodised in either a 0.5 or 1.0 M solution of K2SiO3 (Fig. 3c). However, the level of mineralisation of the extracellular matrix produced by the hMSCs was slightly decreased for ANO-350-1.0 samples compared with the
two other studied surfaces (Fig. 3d). In Fig. 3, SEM images of the hMSCs on a representative etched sample and on TM-ANO-0.5350 and TM-ANO-1.0-350 samples are presented. The cells on all studied surfaces exhibited typical fibroblastic morphology, though, as mentioned above, there were differences in their number and distribution on the studied surfaces. The cells cultured on the TMANO-0.5-350 sample were relatively large, whereas cells on the TM-ANO-350-0.1 sample appeared smaller and less spread out. Martins Junior et al. reported the lowered viability of MC3TC cells on a Ti–15Mo alloy [15]. However, they did not observe changes in cell morphology or cytotoxic effects [15]. Our results further support these previous reports. The applied modification depends on the introduction of silicates into the surface layer, which influences the viability of the hMSCs in a negative way. However, the morphology of the cells is not affected, and the cells exhibit a typical fibroblastic shape. Munir et al. noted that covering the metallic materials with silicon particles favours apatite crystallisation, which later improves the osteogenic cell response of HOB [16]. According to the obtained results, anodising the Ti–15Mo alloy in a 0.5 M K2SiO3 solution can lead to increased bioactivity of the alloy. This may lead to increased ALP activity of the hMSCs, favouring generation of the collagen-rich extracellular matrix by these cells and extracellular matrix calcification.
4. Conclusions In this work, the influence of PEO parameters on the morphology and chemical and phase composition of a Ti–15Mo alloy surface is presented. It was found that the increase in the applied voltage caused the formation of a silicon-enriched porous layer.
cell viability % of change compared to control = 100%
D. Babilas et al. / Materials Letters 100 (2013) 252–256
255
120
TM sample 100 80 60 40 20 0 TM
TM-ANO-0.5-350
TM-ANO-1.0-350
ALP per cell % of change compared to control = 100%
160 140 120 100 80
TM-ANO-0.5-350 sample
60 40 20 0 TM
TM-ANO-0.5-350
TM-ANO-1.0-350
Collagen production per cell % of change compared to control= 100%
600
500
400
300
200
100
TM-ANO-1.0-350 sample
0 TM
TM-ANO-0.5-350
TM-ANO-1.0-350
TM
TM-ANO-0.5-350
TM-ANO-1.0-350
Mineralization per cell % of change conpared to control = 100%
120
100
80
60
40
20
0
at day 10 culture
at day 20 culture
Fig. 3. Cell viability of human MSC after 10- and 21-day cultures; ALP activity of human MSC at day 10 cultures; collagen production, and extracellular matrix calcification of human MSC cultures after 21-day cultures on control and modified Ti–15Mo alloy samples.
This modification also led to an increase of bioactivity of the Ti–15Mo alloy. References [1] Kuphasuk C, Oshida Y, Andres CJ, Hovijitra ST, Barco MT, Brown DT. J Prost Dent 2001;85:195–202.
[2] Donato TAG, de Almeida LH, Nogueira RA, Neimeyer TC, Grandini CR, Caram R, et al. Mater Sci Eng C 2009;29:1365–9. [3] Kumar S, Narayan TSNS. J Dent 2008;36:500–7. [4] Oliveira NTC, Alexio G, Caram R, Guastaldi AC. Mater Sci Eng A 2007;452– 453:727–31. [5] Oliveira NTC, Guastaldi AC. Acta Biomater 2009;5:399–405. [6] Hu H, Liu X, Ding C. Surf Coat Technol 2010;204:3265–71. [7] Liu X, Chu PK, Ding C. Mater Sci Eng R 2004;47:49–121.
256
D. Babilas et al. / Materials Letters 100 (2013) 252–256
[8] Krupa D, Badszkiewicz J, Zdunek J, Smolik J, Słomka Z, Sobczak JW. Surf Coat Technol 2010;205:1743–9. [9] Zhang W, Du K, Yan C, Wang F. Appl Surf Sci 2008;254:5216–23. [10] Krząkała A, Służalska K, Widziołek M, Szade J, Winiarski A, Dercz G, et al. Electrochim Acta 2012. http://dx.doi.org/10.1016/j.electacta.2012.07.075. [11] Simka W, Sowa M, Socha RP, Maciej A, Michalska J. Electrochim Acta 2012. http://dx.doi.org/10.1016/j.electacta.2012.10.130. [12] NIST X-ray photoelectron spectroscopy database; 〈http://srdata.nist.gov/xps/〉.
[13] Blasco T, Camblor MA, Corma A, Perez-Pariente JJ. Am Chem Soc 1993;115:11806–13. [14] Dementjev AP, Ivanova OP, Vasilyev LA, Naumkin AV, Nemirovsky DM, Shalaev DYJ. Vac Sci Technol A 1994;12:423–7. [15] Martins Júnior JRS, Nogueira RA, Araújo RO, Donato TAG, Arana-Chavez VE, A. P.R.A. Claro, et al. Mater Res 2011;14:107–12. [16] Munir G, Koller G, Di Silvio L, Edirisinghe MJ, Bonfield W, Huang JJR. Soc Interface 2011;8:678–88.