Increased reactivity and in vitro cell response of titanium based ...

4 downloads 0 Views 2MB Size Report
Here we investigate the formation of hydroxide groups on sand blasted and acid etched titanium and titanium–zirconium alloy surfaces after anodic polarization ...
J Mater Sci: Mater Med (2013) 24:2761–2773 DOI 10.1007/s10856-013-5020-4

Increased reactivity and in vitro cell response of titanium based implant surfaces after anodic oxidation M. S. Walter • M. J. Frank • M. F. Sunding • M. Go´mez-Florit M. Monjo • M. M. Bucko • E. Pamula • S. P. Lyngstadaas • H. J. Haugen



Received: 5 November 2012 / Accepted: 26 July 2013 / Published online: 3 August 2013 Ó Springer Science+Business Media New York 2013

Abstract In the quest for improved bone growth and attachment around dental implants, chemical surface modifications are one possibility for future developments. The biological properties of titanium based materials can be further enhanced with methods like anodic polarization to produce an active rather than a passive titanium oxide surface. Here we investigate the formation of hydroxide groups on sand blasted and acid etched titanium and titanium–zirconium alloy surfaces after anodic polarization in an alkaline solution. X-ray photoelectron spectroscopy shows that the activated surfaces had increased reactivity. Furthermore the activated surfaces show up to threefold increase in OH- concentration in comparison to the original surface. The surface parameters Sa, Sku, Sdr and Ssk were more closely correlated to time and current density for titanium than for titanium–zirconium. Studies with MC3T3-E1 osteoblastic cells showed that OH- activated surfaces increased mRNA levels of osteocalcin and collagen-I.

1 Introduction

M. S. Walter  M. J. Frank  M. Monjo  S. P. Lyngstadaas  H. J. Haugen (&) Department of Biomaterials, Institute for Clinical Dentistry, University of Oslo, PO Box 1109 Blindern, 0317 Oslo, Norway e-mail: [email protected] URL: www.biomaterials.n

M. Go´mez-Florit  M. Monjo Department of Fundamental Biology and Health Sciences, Research Institute on Health Sciences (IUNICS), University of Balearic Islands, 07122 Palma, Spain

M. S. Walter  M. J. Frank Institute of Medical and Polymer Engineering, Chair of Medical Engineering, Technische Universita¨t Mu¨nchen, Boltzmannstrasse 15, 85748 Garching, Germany M. F. Sunding Department of Physics, University of Oslo, 0316 Oslo, Norway

Recently a titanium–zirconium alloy was introduced as a new dental implant material. This alloy has higher strength than c.p. titanium, the metal most commonly used for dental implants. The alloy contains 13–17 at.% of zirconium in a combination of two recognized biocompatible metals. Studies showed significantly increased elongation and fatigue strength for titanium–zirconium in comparison to pure titanium [1]. It is not only outstanding in its mechanical features, but also performs very well in terms of bone tissue response and is hence suitable for critical implantation sites and low density bone [2]. Titanium– zirconium alloys are furthermore favorable due to a low surface cytotoxicity, good suitability for magnetic resonance imaging and hemocompatibility [3, 4]. Various approaches have been used to increase the surface roughness of dental implants in order to promote bone-tissue attachment. These include the creation of a

M. M. Bucko Department of Ceramics and Refractory Materials, Faculty of Materials Science and Ceramics, AGH University of Science and Technology, Al Mickiewicza 30, 30-059 Krako´w, Poland E. Pamula Department of Biomaterials, Faculty of Materials Science and Ceramics, AGH University of Science and Technology, Al Mickiewicza 30, 30-059 Krako´w, Poland

M. F. Sunding Department of Synthesis and Properties, SINTEF Materials and Chemistry, 0314 Oslo, Norway

123

2762

porous surface structure [5–7] and plasma spraying techniques [8]. Blasting of the titanium surface with Al2O3 [9] or TiO2 [10–17] particles are also commonly applied methods. TiO2 blasting techniques have in particular shown clinical success [16–18]. Blasting techniques are also used in combination with acid etching [19, 20] to create a complex micro structured surface [21]. Other approaches include multi-step etching techniques that result in an anisotropic surface with high frequency of irregularities [22–24]. Several processing techniques are known to modify the surface chemistry of titanium [25, 26]. Acid etching removes the natural oxide layer on titanium as well as on zirconium and furthermore implements liberated hydrogen ions into the metal surface. In addition, acid etching can create a highly complex surface architecture suitable for bone growth [27]. Hydrogen rich surfaces are commercially available in sand blasted and acid etched implant systems such as OsseospeedÒ or SLActiveÒ. These surface also show high surface free energy and wettability [21]. Micro and nano surface architecture and chemical modifications can also be created on titanium surfaces by the use of anodic oxidation [28, 29]. Anodic oxidation can cause hydroxide formation on titanium surfaces [6, 30]. An external layer of compounds containing titanium hydroxide is believed to improve the biological performance of implants made of titanium [31]. The clinical potential for anodically oxidized titanium implants has also been confirmed in an in vivo study suggesting that the presence of a hydroxide layer promotes faster healing and better bone attachment [32]. In the literature, several techniques are reported for formation of hydroxide groups on titanium: Shibata et al. [33] found significantly increased osteoblast response from MC3T3 cells after anodic oxidation in NaCl; Reverchon et al. [34] used hydrolysis of Ti-tetraisopropoxide in supercritical carbon dioxide and generated titanium hydroxide nanoparticles. It has also been reported that hydrogen peroxide in alkaline conditions (HOO–) corroded titanium and formed Ti(OH)4 on the implant surface [35]. The aim of this study was to investigate whether it is possible to deploy OH- groups onto sand-blasted and acid etched surfaces of both titanium and titanium–zirconium, as well as to determine their influence on surface morphology and on in vitro cell response. X-ray photoelectron spectroscopy (XPS) was employed to analyze the surface chemistry. The surface morphology was examined by field emission scanning electron microscope (FE-SEM) and blue-light profilometry. Cell culture experiments were performed with MC3T3-E1 cells to analyze the surface cytotoxicity as well as osteoblast growth and differentiation.

123

J Mater Sci: Mater Med (2013) 24:2761–2773

2 Materials and methods The study was conducted on coin shaped samples with a diameter of 4.39 mm and a height of 2 mm. Materials used were commercially pure grade IV titanium (Ti) and a titanium–zirconium alloy (TiZr) with 13–17 % zirconium (Zr). The surfaces were grit blasted with large-grit (0.25–0.5 mm) and etched in a mixture of hydrochloric and sulfuric acid at 125–130 °C for 5 min (SBAE). The samples were handled under nitrogen cover gas and stored in 0.9 % NaCl solution to achieve a surface composition comparable to the commercially available SLActiveÒ surface (Institut Straumann AG, Basel, Switzerland), as described in the literature [21, 36]. The anodic oxidation setup consisted of a power supply (Protek Dual DC power, Korea) connected to a titanium anode and platinum cathode, a datalogger (NI DAQPad, National Instruments, Asker, Norway) and a magnetic stirrer with heating (IKA-RET Control Visc C, VWR, Kaldbakken, Norway). pH monitoring was performed by a pH electrode (Schott N62, SCHOTT Instruments GmbH, Mainz, Germany) that was driven by a power supply (Xantrex XDL 56-4P, Burnaby, Canada). Temperature measurements were done with a Pt100 device (Pt100, IKA Labortechnik, Staufen, Germany). For the anodic oxidation, the coins were treated in 200 ml of 0.5 M NaCl solution (VWR, Kaldbakken, Norway) adjusted to pH of 8 with 0.2 M NaOH (VWR, Kaldbakken, Norway) and kept at stable pH levels over the process runtime by a SCHOTT Titroline alpha plus titrator with TA50 plus attachment (SI Analytics GmbH, Mainz, Germany). The titrator was used with 0.01 M NaOH. The temperature was set to 20 °C and stirring speed was set to 350 rpm. The coins were unpacked under laminar flow and individually washed with water (deionized, reversed osmosis and autoclaved) for 10 s. The titanium electrode with the sample coins was inserted into the electrolyte; a transfer time of 2 min was kept between insertion of the electrode and the samples under laminar flow and the process start. Samples of both materials were prepared in six different conditions. Process parameters of 0.5 and 1 h duration were used with current densities of 0.5, 1 and 3 mA/cm2. Furthermore Ti SBAE and TiZr SBAE coins were included as control. After the process the coins were kept in the electrolyte for transferring them to the laminar flow workbench. Coins were unmounted under laminar flow, washed for 5 s in water and air dried in laminar flow for 30 min. The coins were placed in sterile Eppendorf tubes (Axygen, Union City CA, USA) after the process awaiting further analysis. Storage of the coins was dry and dark at 4 °C. In vitro coins were gamma-irradiated with 30.2 kGy at

J Mater Sci: Mater Med (2013) 24:2761–2773

room temperature (22 °C) with Co-60 source (Institutt for Energiforsknings, Kjeller, Norway). 2.1 Profilometer A PLl 2300 (Sensofar-Tech S.L., Terrassa, Spain) blue light laser profilometer and interferometer using a 509 EPI (Nikon, Tokyo, Japan) confocal objective was used to assess an extended topography of 2 9 2 images. Each had a viewing area of 253 9 190 lm at 20 percent overlapping. Eight images of each material were analyzed. The advanced topography software Sensomap 4.1 Plus (Sensofar-Tech S.L., Terrassa, Spain) for dimensional and surface state metrology was used to process the measured data. The parameters that were believed to be most relevant in accordance with Lamolle et al. [37] were the total surface area in percent of a completely flat surface (Sdr), average roughness (Sa), skewness of the height distribution (Ssk), kurtosis of the height distribution (Sku). 2.2 SEM The nano and micro analysis of the surface was done with a FEI Quanta 200 (FEI Hillsboro, Oregon, USA) FE-SEM. The samples were platinum sputtered prior to analysis and mounted on a 45° aluminum rack with conductive carbon tape. Images were acquired with a working distance between 5 and 7 mm. Acceleration voltages of 3–5 kV were applied to achieve magnifications of 30,000 and 50,0009. High vacuum (HV) operation mode was used for the examination of all samples. 2.3 Fluorescence microscope A Leica DMRB (Leica Microsystems CMS GmbH, Wetzlar, Germany) Fluorescence Microscope was used for the optical assessment of the coin surface. Main focus was homogeneity and colour shade differences in comparison to Ti and TiZr SBAE coins. The samples were illuminated by external lightning. A 1.259 objective was used, which in combination with the camera system added up to a total magnification of 12.59. The software for image capturing was set at an exposure time of 59 ms and ISO 100 sensitivity. 2.4 XPS The XPS analysis was carried out on an Axis UltraDLD XP spectrometer (Kratos Analytical, Manchester, UK) using monochromatic Al Ka radiation (hm = 1486.6 eV). Survey spectra were recorded in a range between 1,100 and 0 eV binding energy (BE). Detail spectra were acquired in the energy regions of O 1s, Ti 2p, C 1s and Zr 3d. The

2763

instrument resolution was 1.1 eV for the survey scans and 0.55 eV for the detail scans. The analysis area was 300 9 700 lm. Samples were mounted on a sample bar with conductive carbon tape. The energy shift due to surface charging was below 1 eV based on the C 1s peak position relative to established BEs, therefore the experiment was performed without charge compensation. All BEs were energy referenced based on the C 1s peak related to C–C and C–H bonds at 285 eV BE, as described in several papers for oxide layers on titanium [38–40]. The offset for the C–O peak was set to ?1.5 eV BE and for O–C=O to ?4.0 eV BE from the C–C peak, based on the average of the C 1s spectra and reference values for organic compounds [41]. The Na KLL peak was found at around 496 eV kinetic energy as described in the literature [39, 42]. The O 1s photoelectron peak was divided in three individual components, each 1.1–1.5 eV distant to each other as reported earlier by McCafferty et al. [38]. FWHM was set up between 1.1 and 2.0 eV. The determination of the OH concentration in respect to the O 1s photoelectron peak was done according to McCafferty et al. [38]. In contrast to the literature, the concentration was examined with regard to the peak area instead of the peak height. This method was believed to result in higher accuracy, as the calculated area is a more accurate measurement of the total amount of a compound than the peak height (see Eqs. 1–5). Furthermore, the Equations displayed were slightly modified in comparison to McCafferty et al. for improved applicability. %(C) and %(O) are the total concentrations of C and O in the sample calculated to a total of 100 % with Eqs. 1 and 2. A(O) and A(C) are the total peak areas after Shirley background subtraction of the oxygen and carbon peaks, respectively, corrected for the transmission function of the instrument. RSF(O) and RSF(C) are the relative sensitivity factors of oxygen and carbon. A(OH-)meas is the area of the O 1s peak at the energy typical for OH- groups in titanium hydroxide and for C–O bonds in organic matter. It is obtained from peak fitting, corrected for the transmission function of the instrument. The contribution from C–O bonds to A(OH-)meas is obtained according to Eq. 3. A(C–O*) (* marks the considered atom) is the area of the oxygen peak related to C–O bonds and x(C*–O) is the fraction of C–O related signal in the carbon peak. The C 1s peak detected at approximately ?4 eV relative to the peak attributed to C–C bonds fits for C–O–C*=O bonds but is at a too low BE to relate to HO– C*=O bonds [41]. This means that a C–O bond is related to each O–C=O group (the C*–O–C=O carbon) and that A(O– C*=O) has thus to be disregarded in the determination of A(C–O*). This differs from the procedure described by McCafferty et al. [38].

123

2764

J Mater Sci: Mater Med (2013) 24:2761–2773

Equation 4 subtracts the contributions of C–O* from the peak OHmeas to find the area A(OH-) of the OH- component related to inorganic hydroxides in the oxygen peak. Eq. 5 calculates x(OH-), the fraction of OH- in relation to the total amount of oxygen related to inorganic compounds (oxide and hydroxide of Ti and TiZr, respectively). x(OH-) is thus a good parameter to evaluate the effect of the treatments on the anodic oxidation of the implant material surfaces. 0 1 AðOÞ RSF B  ðO Þ C % (O) ¼ @ ð1Þ AðOÞ þ AðCÞ A RSFðOÞ RSFðCÞ 0

1 AðCÞ RSF B  ðCÞ C % (C) ¼ @ A AðCÞ A ðOÞ þ RSFðCÞ RSFðOÞ AðC  O Þ ¼ xðC  OÞ 

 % ðCÞ  AðOÞ % ðO Þ



AðOH Þ ¼ AðOH Þmeas  AðC  O Þ xðOH Þ ¼

ð2Þ

AðOH Þ AðOH Þ þ A ðO2 Þ

ð3Þ ð4Þ ð5Þ

2.5 Phase composition

differentiation and cell toxicity. The same number of cells was cultured in parallel in plastic tissue culture dishes in all experiments. Trypan blue stain was used to determine total and viable cell number. For the experiments, MC3T3-E1 cells were maintained for 14 days on the implants in a-MEM supplemented with 10 % FBS and antibiotics. Culture media was changed every other day. Culture media was collected after 24 h to test the toxicity of the treatments and of the different implant surfaces (LDH activity). To study cell differentiation, cells were harvested after 14 days and collagen type I (Coll-I) and osteocalcin (OC) gene expression were analyzed using real-time RT-PCR. The presence of LDH activity in the culture media was used as an index of cell death. LDH activity was determined spectrophotometrically after 30 min incubation at 25 °C using 50 ll of culture media and adding 50 ll of the reaction mixture, by measuring the oxidation of NADH at 490 nm in the presence of pyruvate, according to the manufacturer’s kit instructions (Cytotoxicity Detection kit, Roche Diagnostics). Positive high control (100 %) was cell culture media from cells seeded on plastic culture dishes and incubated with Triton X-100 at 1 %. Negative low control (0 %) was cell culture media from cells seeded on plastic culture dishes without any treatment. Results were presented using Eq. 6. exp: value  low control  100 high control  low control

Phase composition of the samples was determined by grazing incident X-ray diffraction (GIXD) measurements in an Empyrean diffractometer (PANalytical B.V., Almelo, the Netherlands) with a CuKa radiation employing a secondary graphite monochromator and the PIXcel 2D detector. The incident angle was 1° and the penetration depth for this angle was less than 0.8 lm which means that whole layer and a part of the metallic substrate were analyzed. The qualitative analysis was based on the database from the International Center for Diffraction Data (ICDD). The data for the quantitative analysis and the cell parameters were performed using Rietveld refinement and structure models from the International Center for Structure Data. Crystallite sizes were calculated based on X-ray line broadening and Sherrer equation.

Cytotoxicity ð%Þ ¼

2.6 In vitro cell study

Table 1 Primer sequences used for real-time RT-PCR analysis

Cell culture conditions used were as described earlier by Lamolle et al. [43]. Cells were subcultured 1:5. All experiments were performed in the same passage of the MC3T3-E1 cells (passage 16). Ti and TiZr samples treated at 1 h with a current density of 3 mA/cm2 were examined in vitro as well as untreated control samples (Ti SBAE and TiZr SBAE). The coins were placed in a 96-well half area plate (4.5 mm well) and 7 9 103 cells were seeded on each coin to study cell

123

ð6Þ

Total RNA was isolated with Tripure (Roche Diagnostics), following the instructions of the manufacturer. RNA was quantified using a spectrophotometer set at 260 nm (NanoDrop Technologies, Wilmington, DE, USA). Real-time PCR was performed for two housekeeping genes, 18S ribosomal RNA (18S rRNA), glyceraldehyde-3phosphate dehydrogenase (GAPDH), and two target genes, collagen type I and OC. The primer sequences are detailed in Table 1. The same amount of total RNA (0.1 lg) from each sample was reverse transcribed to cDNA at 37 °C for 60 min in a final volume of 20 ll, using High Capacity

Gene

Primer sequence

18S rRNA

S 50 -GTAACCCGTTGAACCCCATT-30 A 50 -CCATCCAATCGGTAGTAGCG-30

GAPDH

S 50 -ACCCAGAAGACTGTGGATGG-30 A 50 -CACATTGGGGGTAGGAACAC-30

Coll-I

S 50 -AGAGCATGACCGATGGATTC-30 A 50 -CCTTCTTGAGGTTGCCAGTC-30

OC

S 50 -CCGGGAGCAGTGTGAGCTTA-30 A 50 -TAGATGCGTTTGTAGGCGGTC-30

J Mater Sci: Mater Med (2013) 24:2761–2773

RNA to cDNA kit (Applied Biosystems). Each cDNA was diluted 1/4. Real-time PCR was performed in the LightCycler 480Ò (Roche Diagnostics, Mannheim, Germany) using SYBR green detection. Each reaction contained 7 ll LightcyclerFast Start DNA Master PLUS SYBR Green I (containing Fast Start Taq polymerase, reaction buffer, dNTPs mix, SYBR Green I dye and MgCl2), 0.5 lM of each, the sense and the antisense specific primers and 3 ll of the cDNA dilution in a final volume of 10 ll. The amplification program consisted of a preincubation step for denaturation of the template cDNA (5 min 95 °C), followed by 45 cycles consisting of a denaturation step (10 s 95 °C), an annealing step (10 s 60 °C) and an extension step (10 s 72 °C). After each cycle, fluorescence was measured at 72 °C. A negative control without cDNA template was run in each assay. Real-time efficiencies were calculated from the given slopes in the LightCycler 480 software using serial dilutions, showing all the investigated transcripts high real-time PCR efficiency rates, and high linearity when different concentrations are used. PCR products were subjected to a melting curve analysis on the LightCycler and subsequently 2 % agarose/TAE gel electrophoresis to confirm amplification specificity, Tm and amplicon size, respectively. Relative quantification after PCR was calculated by dividing the concentration of the target gene in each sample by the mean of the concentration of the two reference genes (housekeeping genes) in the same sample using the Advanced relative quantification method provided by the LightCycler 480 analysis software version 1.5 (Roche Diagnostics, Mannheim, Germany). 2.7 Statistics Analysis of the raw data was carried out with Excel 2007 (Microsoft, Redmond, USA) and SigmaPlot 11.0 (Systat Software Inc, Chicago, USA). All datasets were examined for parametric or non-parametric distributions (ShapiroWilktest). One-way ANOVA was executed for the statistical evaluation with multiple pairwise comparison with Holm Sidak test of the single batches for parametric distributions. ANOVA on ranks was performed for nonparametric distributions with multiple pairwise comparison with Dunn’s test. Pairwise comparisons for in vitro study were assessed with Student-t test for parametric distribution of cytotoxicity and OC gene expression and nonparametric with Mann–Whitney test for Coll-I and ALP gene expression. Differences were considered significant (*) for P B 0.05 and highly significant (**) for P B 0.01, and labeled respectively. All parametric data are presented as mean values ± standard deviation and non-parametric data as median ± interquartile range.

2765

A correlation study was carried out in the course of this work to assess the connection between the surface characterization parameters and time or current. A two tailed Spearman correlation test was carried out, which produces the Spearman correlation coefficient, r. The coefficient can vary between -1 and ?1, which would represent perfect negative, respectively perfect positive correlation. The results were interpreted as follows: no correlation if |r| \ 0.3; correlation if 0.3 \ |r| \ 0.5; and strong correlation if 0.05 \ |r| \ 1 [44].

3 Results 3.1 Surface topography The profilometer analysis of the anodically oxidized Ti SBAE surfaces revealed changes to several surface structure parameters (Fig. 1). Ti SBAE samples had a Sa of 1.89 lm. Sa values were significantly elevated for Ti at 3 mA/cm2 at 0.5 and 1 h, as well as for 1 mA/cm2 and 1 h, showing Sa values of 2.18, 2.01 and 2.07 lm respectively. TiZr samples showed significant increase in Sa for 0.5 mA/ cm2 at treatment times of 0.5 and 1 h furthermore for 3 mA/cm2 and 1 h treatment, with Sa values of 2.05, 1.92 and 1.99 lm respectively. TiZr SBAE samples had a Sa value of 1.86 lm The Sa value is an indicator for the surface roughness, defined as the average height deviation of a surface from the mean square. Sku showed significant increase from a Sku of 3.24 in the Ti SBAE to a Sku of 4.63 in the 3 mA/cm2 at 1 h treatment group. The values for TiZr were significantly elevated to a Sku of 4.19 for 0.5 mA/cm2 and 0.5 h of treatment group in comparison to TiZr SBAE with a Sku of 3.16. All Sku values were ranging above 3, indicating the presence of steep summits on the surface. The analysis of the Sdr value, which shows the percental increase in the total surface area compared to a completely flat surface, was significant for all Ti groups in comparison to control with P \ 0.05. Highest increase was found for the 1 h and 1 mA/cm2 group with an Sdr of 107.5 % in comparison to 65.1 % for Ti SBAE. For TiZr the 0.5 mA/cm2 and 0.5 h treatment group increased significantly to 77.9 % in comparison to TiZr SBAE with 54.1 %. The Ssk value for Ti was significantly increased in comparison to control for 3 mA/cm2 at 0.5 h as well as for all 1 h treatments, with values between 0.155 and -0.076 compared to Ti SBAE with -0.396. TiZr showed significant increase to 0.246 for 0.5 h at 0.5 mA/cm2 treatment in comparison to TiZr SBAE with -0.213. The Ssk value indicates a surface composed of peaks emerging from a relatively flat surface for positive values of Ssk, whereas a negative Ssk indicates the presence of a wide plateau with fine deep valleys.

123

2766

The correlation study revealed strong positive (r [ 0.65) and highly significant (P \ 0.0001) correlations between surface roughness (Sa) and current density and time for Ti (Table 2). The correlation between time and Sa was also highly significant (P = 0.0056) and positive (r = 0.48) for TiZr, but no correlation was found for the current density. Kurtosis (Sku) showed strong (r = 0.512) and highly significant (P = 0.0029) correlation to current density for Ti, whereas no significant correlation (r \ 0.3) was found for TiZr. The Sdr value showed strong positive (r [ 0.6) and highly significant (P [ 0.001) correlation between Ti and time and current density. For TiZr a correlation between Sdr and time was found significant (P = 0.0237) and positive (r = 0.39). The skewness of the height distribution (Ssk) showed statistically highly significant (P \ 0.001) and strong (r [ 0.5) correlations to time and current density in the Ti group. No significant correlation with TiZr SBAE was found for this parameter. The anodically oxidized Ti samples showed structural changes on micron and sub-micron levels when compared to the untreated Ti SBAE surfaces (Fig. 2). TiZr SBAE samples expressed a rough surface morphology in the SEM analysis. Anodically oxidized TiZr

J Mater Sci: Mater Med (2013) 24:2761–2773

samples showed a smoother surface compared to the TiZr SBAE surface. The oxidized TiZr surface had less distinct structural changes in comparison to anodically oxidized Ti and showed lower porosity (Fig. 2). As the treatment produced distinct differences in coin color, a low resolution microscopic analysis was used to display these differences. Figure 3 shows the comparison between anodically oxidized Ti and anodically oxidized TiZr. The latter showed a deep blue color after the anodic Table 2 Spearman rank values for surface and treatment parameters Sa

Sku

Sdr

Ssk

Titanium Time

0.66**

0.25

0.63**

0.59**

Current density

0.77**

0.51**

0.67**

0.66**

Titanium–zirconium Time

0.48**

-0.06

0.40**

0.24

Current density

0.29

-0.03

0.35

0.34

No correlation for |r| \ 0.3; correlation for 0.3 \ |r| \ 0.5; and strong correlation for 0.5 \ |r| \ 1; * P \ 0.05, ** P \ 0.01. Only |r| [ 0.3 and P \ 0.05 are displayed. Process parameters of 0.5 and 1 h duration were used with current densities of 0.5, 1 and 3 mA/cm2

Fig. 1 Surface analysis for Ti and TiZr. One way ANOVA was performed group wise (Ti and TiZr) against control (*P B 0.05 compared to control). a Sa, b Sku, c Sdr, d Ssk. Parameters are displayed as mean ± SD

123

J Mater Sci: Mater Med (2013) 24:2761–2773

2767

Ti SBAE

25 µm

5 µm

1 µm

5 µm

1 µm

5 µm

1 µm

5 µm

1 µm

Ti Hydrox.

25 µm

TiZr SBAE

25 µm

TiZr Hydrox.

25 µm

Fig. 2 Images of anodically oxidized (hydrox) and untreated Ti SBAE and TiZr SBAE. Ti SBAE and TiZr SBAE are displayed as a reference in three magnifications. Anodically oxidized samples were treated at 3 mA/cm2 for 1 h and are displayed in three different magnifications

oxidation. This color change was homogenous over the surface structure. The anodically oxidized Ti surface had darkened in comparison to the untreated Ti SBAE surface, but did not develop a distinct color. 3.2 Surface chemistry The element analysis of Ti surfaces with XPS showed increased carbon levels for rising anodic oxidation times (Fig. 4a). Nevertheless, the sample treated for 1 h at 3 mA/ cm2 showed the lowest carbon concentration and highest oxygen content amongst the treated groups. The highest carbon level with nearly 80 at.% was found in the samples anodically oxidized for 1 h at 1 mA/cm2. In contrast, titanium and oxygen levels were lowest for this group.

Chlorine was found in all samples in concentrations up to 2 at.% except for the Ti SBAE sample. Oxygen and titanium levels showed the opposite devolution in comparison to carbon. Some coins furthermore showed F, P, Ca and N in low concentrations (B1.6 at.%). Analysis of the TiZr coins revealed high carbon levels increasing with treatment time (Fig. 4b). The 1 h and 1 mA/cm2 group showed lowest carbon and highest oxygen concentrations of the anodically oxidized samples. This group furthermore was found to have the highest surface titanium levels of all TiZr groups. The highest carbon levels were reached for the group treated for 1 h at 3 mA/cm2, the lowest carbon content for the TiZr SBAE sample. Inversely corresponding to the carbon levels, oxygen was lowest for the sample treated for 1 h at 3 mA/cm2.

123

2768

J Mater Sci: Mater Med (2013) 24:2761–2773

Fig. 3 Optical microscopy analysis of Ti SBAE (a) and TiZr SBAE (b), anodically oxidized Ti (c) and anodically oxidized TiZr (d)

Fig. 4 XPS element analysis Ti (a) and TiZr (b). Analysis of elements detected on Ti and TiZr samples, comprising all treatment groups and SBAE samples as control

Titanium concentrations were especially high for the two samples with the lowest carbon concentration—SBAE and the anodically oxidized surface treated for 1 h at 1 mA/ cm2. The zirconium content at the surface of the alloy decreased upon hydroxylation from 2.93 at.% in the SBAE sample to 1 at.% in the anodically oxidized samples. Furthermore the surfaces of some samples showed low concentrations of Na, Cl, F and Ca (B1.58 at.%). The majority of carbon on the surface was bound as C– H or C–C in all samples. These compounds had up to 85 % relative share of total carbon on the surface for Ti and up to 90 % for TiZr in the anodically oxidized groups. In contrast, the SBAE samples had a maximum concentration of 65 % for Ti and 60 % for TiZr. The second biggest part in

123

carbon was C–O and the smallest part O–C=O for all samples. O2- was the dominant component in the oxygen peak. In Table 3 the relative OH- concentrations on the surface of the implant material and the relative increase of OH- in comparison to the untreated Ti SBAE and TiZr SBAE samples are displayed. The relative OH- concentration decreased from 20 % in the Ti SBAE sample, to 14–18 % in most anodically oxidized samples. The hydroxide concentrations detected in the Ti samples treated for 1 h at 0.5 and 1 mA/cm2 are not considered due to unreliable constitution of the spectra. The TiZr SBAE sample showed 9 % OH-. The values observed in the anodically oxidized TiZr samples ranged from 4 to 27 %—the

J Mater Sci: Mater Med (2013) 24:2761–2773

2769

Table 3 Relative OH- in Ti and TiZr Treatment

OH of total O (%)

Change of OH (%)

Ti

TiZr

Ti

SBAE

19.5

8.6

0.5 h, 0.5 mA/cm2

18.1

18.0

-7

110

0.5 h, 1 mA/cm2

18.4

10.5

-6

22

2

-27

-53

0.5 h, 3 mA/cm

TiZr

14.3

4.0

1 h, 0.5 mA/cm2



13.0



51

1 h, 1 mA/cm2



19.5



126

1 h, 3 mA/cm2

14.9

26.9

-24

212

-

The table displays the share of OH in percent of total oxygen, as well as the increase in OH- in percent, compared to untreated Ti SBAE and TiZr SBAE Fig. 5 X-ray diffraction patterns of untreated TiZr SBAE and anodically oxidized TiZr

latter corresponding to a three-fold increase in OH- concentration compared to TiZr SBAE. 3.3 Phase composition Figure 5 shows X-ray diffraction patterns while Table 4 shows phase composition and parameters of the unit cell of TiZr as received and TiZr anodically oxidised. Phase analysis of both samples revealed a presence of titanium– zirconium alloy with the hexagonal a-Ti structure. The layer on the metallic surface was composed mostly of the cubic (Fm3m) titanium hydride with a little non-stoichiometry, TiH2-x, where x was about 0.03. Broadening of the (111) line of this phase showed that the crystallite size was about 18 nm. The balance phase was the orthorhombic (Cccm) titanium hydride, TiH2-x, with non-stoichiometry about x = 0.3. The size of crystallites of this phase was 21.5 nm. 3.4 In vitro The graph in Fig. 6 displays the cytotoxicity obtained in the in vitro examination. Ti and TiZr both exhibited higher cell toxicity for the anodically oxidized samples in comparison to their SBAE references, although these values were not significant against the reference tissue culture plastic (TCP) surface. For TiZr there was a significant increase in cytotoxycity (P = 0.014) but not for Ti (P = 0.071). The TiZr SBAE samples showed significantly lower (P B 0.001) toxicity compared to Ti SBAE. Figure 7a displays relative mRNA levels of Collagen I, which was significantly increased (P B 0.001) for both materials in the anodically oxidized group. TiZr SBAE had significantly lower (P = 0.028) Collagen I mRNA levels in comparison to Ti. Anodically oxidized TiZr showed significantly lower (P = 0.003) Collagen I mRNA levels in comparison to anodically oxidized Ti. In Fig. 7b significant

increase (P = 0.003) in OC expression was observed for Ti in the anodically oxidized group, whereas the levels decreased significantly (P = 0.042) for TiZr. The OC expression in the anodically oxidized TiZr group was found to be significantly lower (P B 0.001) in comparison to anodically oxidized Ti.

4 Discussion Particularly the factors Sa and Sku emphasize the appearance of topographical changes in the micron level. Increased Sa can either constitute due to erosion in the valleys or accumulation on the peaks. The increase in Sku and the tendency of Ssk to become more positive with increasing time and current density suggests erosion in the valleys was more dominant over the accumulation of peak structures. This is also in accordance with the work of Videm et al., who suggested liberation of Ti4? in pits on the surface. In that theory the re-addition would most likely occur at the peaks of the surface. Changes in the surface topography as observed in the profilometer analysis (Fig. 1) were more significant for longer treatment times and higher current densities, as described by the correlation study (Table 2). Yet this also confirmed that for TiZr the impact of current density was less important for topographical changes. With rising levels of Sa, Ssk and Sku a positive effect on the in vivo performance of anodically oxidized implants could be observed, as these factors have been associated with improved bone healing in earlier investigations [37]. An increase of total surface area indicated by Sdr could furthermore enhance the potential for biomechanical interlocking and the bone-implant interface as indicated by Klokkevold et al. [45]. It is noteworthy that the correlation study revealed no influence of current density on any of the surface parameters for TiZr, but

123

2770 Table 4 Phase composition and parameters of the unit cell of TiZr and TiZr anodic oxidised

J Mater Sci: Mater Med (2013) 24:2761–2773

Sample TiZr

TiZr anodically oxidized

Component

˚) a (A

TiZr-hexagonal

24

2.9766

Cubic TiH2

72

4.4299

Orthorhombic TiH2

4

4.1929

TiZr-hexagonal

27

2.9795

Cubic TiH2

70

4.4335

Orthorhombic TiH2

3

4.2001

Fig. 6 In vitro response of MC3T3-E1 cells to SBAE and anodically oxidized (hydrox) Ti and TiZr. LDH activity measured from culture media of MC3T3-E1 cells seeded on implants (Ti and TiZr) SBAE, or anodically oxidized (hydrox) after 24 h. Hydrox surfaces were compared against SBAE (*P \ 0.05) within each material. Some differences were found when comparing Ti versus TiZr surfaces (#P \ 0.05) in each treatment group

strong correlations for all analyzed parameters of Ti (Table 2). The SEM images of anodically oxidized Ti showed ‘‘cone shaped’’ structures growing on the surface (Fig. 2). As the TiZr alloy comprises a different structure, differences in diffusion speed and conductivity are probable and could serve as an explanation for the lack of these cone shaped structures on TiZr [2, 46]. On the other hand, TiZr coins turned into a deep dark blue after anodic oxidation whereas Ti coins were found darkened with a slightly blue shine (Fig. 3). This phenomenon was also observed by Sul et al. [47] and Pan et al. [48] who found the blue color after anodizing and linked it to a thickening oxide layer. Even though the literature correlates the blue color with the appearance of higher oxide concentration on the surface, the anodically oxidized TiZr SBAE samples showed in average lower total oxygen content in comparison to Ti SBAE. The chemical surface analysis with XPS revealed a high surface coverage with carbon (Fig. 4), which can be attributed to the influence of environmental carbon contamination [18, 49, 50]. As this contamination was dominant to a bigger extent on the anodically oxidized samples than on SBAE, it is assumed that the anodically oxidized surface has a higher attractive potential and therefore a

123

Phase composition (mass%)

˚) b (A

˚) c (A 4.7255

4.2859

4.5626 4.7269

4.2834

4.5674

higher reactivity [28]. The decrease in oxygen is an effect from the subsequent coverage of the anodically oxidized surface with carbon, attenuating its signal in the spectra. Confirmation for that can be found as well in the lowered titanium and zirconium levels in the surface of these samples [51]. The oxide thickness is also in the range of, or larger than the probing depth of XPS. An increased oxide thickness will thus not affect the spectra. The OH- levels showed a small decrease in the Ti samples upon anodic oxidation (Table 3). No clear trend is apparent on the amount of formed hydroxide groups in function of neither the duration of the treatment nor of the employed current. The OH- levels registered on the anodized TiZr samples were increased in comparison to the TiZr SBAE sample. A decrease in OH- concentration with an increase in sample current is observed at 0.5 h and 3 mA/cm2 (Table 3). The absence of consistent effects of treatment time and current density on the OH- concentration on the Ti and TiZr samples can indicate that the high level of surface contamination prohibits reliable quantification of the XPS data. The differences in the surface roughness between the samples attenuate further the significance of the data. As reported by Ohtsuka et al. [52] our study supports the hypothesis that anodic oxidation effectively creates OHgroups on the sample surfaces. These groups are believed to have a high potential for attraction of carbon if exposed to air. Tugulu et al. [39] could demonstrate a strong increase in carbon contamination during storage of hydroxylated Ti surfaces. Furthermore Feng et al. [53] reported higher surface energy, and increasing carbon coverage for heat treated titanium. Therefore, an increase in the OH- concentration would also increase the potential for the attraction of carbon, and as a result more total carbon would be observed, if no cover gas or other protection is applied [18]. If a construction of the surface oxides with an inner layer of titanium oxide and an outer layer containing OH- groups is assumed, increased coverage of the outer layer with carbon would result in lower oxygen and titanium levels in the XPS analysis [51]. This layer construction is described in a similar way by Sul et al. [48].

J Mater Sci: Mater Med (2013) 24:2761–2773

2771

Fig. 7 In vitro response of MC3T3-E1 cells to SBAE and anodically oxidized (hydrox) Ti and TiZr surfaces. a Gene expression levels of collagen-I, b OC in MC3T3-E1 cells seeded implants (Ti and TiZr) SBAE, or anodically oxidized (hydrox) after 14 days of cell differentiation. Values represent the median ± IQR for (a) and

mean ± SD for (b). Hydrox surfaces were compared against SBAE (*P \ 0.05) within each material. Some differences were found when comparing Ti versus TiZr surfaces (#P \ 0.05) in each treatment group

The efficiency of surface activation by anodic oxidation seems to be lower on Ti in comparison to TiZr, if the assumption is made that more OH- subsequently attracts more environmental carbon [53]. Figure 4 shows increase in carbon concentration upon anodic oxidation treatment. It is noteworthy that the effect of the anodic oxidation on the surface topography and chemistry differs between Ti and TiZr. This suggests that Zr affects the electrochemical reactions during the hydroxylation. As pointed out by Martins et al. alloying of Zr can increase microstructure homogeneity and induce a finer microstructure. Furthermore they proved enhanced electrochemical corrosion behavior [54]. This could result in a more homogenous surface after anodic oxidation and furthermore a higher concentration of OH- on the TiZr surfaces. At the same time the finer structure on TiZr might be the reason why Ti shows cone shaped structures and TiZr does not. The characterization of the compounds established under anodic oxidation is difficult to perform with analytical methods. Theoretical considerations suggest the event of local oxide breakdown and pitting with the release of Ti4? ions into the electrolyte. Due to low solubility of these ions in pH 8 a further reaction with liberated OH- to TiO2? is likely. These titanyl ions are then available for the formation of solid products with liberated O2-, OH- and Clon the Ti surface. In that respect it seems also likely that the addition of negative organic molecules could replace the OH- in the reaction with Ti4?, which would give room for addition of different molecules and even drugs. Despite the obvious differences between Ti and TiZr, these considerations are believed to be valid for both materials. The biological performance of Ti and TiZr surfaces was evaluated in vitro in terms of cytotoxicity and differentiation of MC3T3-E1 cells. TiZr was found to be less cytotoxic compared to Ti. Furthermore, a decrease in cell viability of cells cultured on TiZr upon anodic oxidation

treatment was observed. It is important to note that even though LDH activity was increased for the anodically oxidized groups, its cytotoxicity was not significantly higher than the reference surface TCP. With regards to osteoblast differentiation, the Ti surface anodically oxidized at 3 mA for 1 h showed increased Coll-I and OC mRNA levels compared to the Ti SBAE surface. Coll-I is an early marker expressed during the proliferation stage of osteoblastic cells. It is associated with the extracellular matrix biosynthesis [55]. In a later stage OC is expressed, which is related to calcium deposition in the extra cellular matrix under its mineralization period [43, 56]. These two markers are related to major developmental stages of osteoblasts, a period of growth and proliferation (collagen type I) and higher differential stages that include mineralization (osteocalcin) [55]. This suggests that anodically oxidized Ti surfaces provide the best properties for supporting osteoblast differentiation, without significant negative effects on cell viability. In this respect, it was found in accordance with the literature that the increase in OHgroups probably triggers better osteoblast attachment and higher cell activity [53, 57]. Furthermore higher surface area and porosity on the anodically oxidized Ti surfaces play an important role for osteoblast response [53, 58, 59]. Therefore, the performance of anodically oxidized TiZr has to be seen in context with lower surface area and roughness, lowering the osteoblast response compared to oxidized Ti. Nevertheless the beneficial effects of surface OHgroups on TiZr can be seen with respect to increased Coll-I expression.

5 Conclusion In this paper the suitability of anodic oxidation for improvement of surface chemistry and topography could

123

2772

be demonstrated on titanium based implant materials. Surface topography of Ti was strongly influenced by time and current of anodic oxidation. The Ti surface was more susceptible to topographical and morphological changes due to anodic oxidation. Both Ti and TiZr expressed high surface reactivity, demonstrated by substantial carbon coverage. Yet, increase in OH-groups due to anodic oxidation was more distinct on TiZr. The in vitro study confirmed the production of a highly reactive surface with anodically oxidized Ti expressing significantly increased collagen type-I and osteocalcin mRNA levels compared to the Ti SBAE surface on MC3T3-E1 cells. It was concluded that anodic oxidation of Ti is more promising due to its comprehensive chemical and structural changes, while on oxidized TiZr chemical changes were more dominant. Yet, TiZr demonstrated high surface reactivity and furthermore increased in vitro Coll-I gene expression levels. Anodic oxidation of titanium based materials seems to be a suitable way to activate surfaces and embraces potential for use in biomedical applications.

References 1. Bernhard N, Berner S, De Wild M, Wieland M. The binary TiZr Alloy—a newly developed Ti alloy for use in dental implants. Forum Implantol. 2009;5:30–9. 2. Gottlow J, Dard M, Kjellson F, Obrecht M, Sennerby L. Evaluation of a new titanium–zirconium dental implant: a biomechanical and histological comparative study in the mini pig. Clin Implant Dent Relat Res. 2010;14(4):538–45. 3. Niinomi M. Fatigue performance and cyto-toxicity of low rigidity titanium alloy, Ti-29Nb-13Ta-4.6 Zr. Biomaterials. 2003;24(16): 2673–83. 4. Seligson D, Mehta S, Mishra AK, FitzGerald TJ, Castleman DW, James AH, et al. In vivo study of stainless steel and Ti-13Nb13Zr bone plates in a sheep model. Clin Orthop Relat Res. 1997;343:213–23. 5. Cui X, Kim HM, Kawashita M, Wang L, Xiong T, Kokubo T, et al. Preparation of bioactive titania films on titanium metal via anodic oxidation. Dent Mater. 2009;25(1):80–6. 6. Shih YH, Lin CT, Liu CM, Chen CC, Chen CS, Ou KL. Effect of nano-titanium hydride on formation of multi-nanoporous TiO2 film on Ti. Appl Surf Sci. 2007;253(7):3678–82. 7. Sul Y-T, Johansson CB, Petronis S, Krozer A, Jeong Y, Wennerberg A, et al. Characteristics of the surface oxides on turned and electrochemically oxidized pure titanium implants up to dielectric breakdown: the oxide thickness, micropore configurations, surface roughness, crystal structure and chemical composition. Biomaterials. 2002;23(2):491–501. 8. Baik KH. Microstructural evolution and tensile properties of Ti– Al–V alloys manufactured by plasma spraying and subsequent vacuum hot pressing. Mater Trans. 2006;47(4):1198–203. 9. Wennerberg A, Albrektsson T, Lausmaa J. Torque and histomorphometric evaluation of c.p. titanium screws blasted with 25and 75-microns-sized particles of Al2O3. J Biomed Mater Res. 1996;30(2):251–60. 10. Gotfredsen K, Nimb L, Hjorting-Hansen E, Jensen JS, Holmen A. Histomorphometric and removal torque analysis for TiO2-blasted

123

J Mater Sci: Mater Med (2013) 24:2761–2773

11.

12.

13.

14.

15.

16.

17.

18.

19.

20.

21.

22.

23.

24.

25.

26.

27.

titanium implants. An experimental study on dogs. Clin Oral Implants Res. 1992;3(2):77–84. Gotfredsen K, Wennerberg A, Johansson C, Skovgaard LT, Hjorting-Hansen E. Anchorage of TiO2-blasted, HA-coated, and machined implants: an experimental study with rabbits. J Biomed Mater Res. 1995;29(10):1223–31. Mustafa K, Silva Lopez B, Wennerberg A, Arvidson K. Attachment and proliferation of human oral fibroblasts to titanium surfaces blasted with TiO2 particles. A scanning electron microscopic and histomorphometric analysis. Clin Oral Implants Res. 1998;9(3):195–207. Cooper LF, Masuda T, Whitson SW, Yliheikkila P, Felton DA. Formation of mineralizing osteoblast cultures on machined, titanium oxide grit-blasted, and plasma-sprayed titanium surfaces. Int J Oral Maxillofac Implants. 1999;14(1):37–47. Rasmusson L, Roos J, Bystedt H. A 10-year follow-up study of titanium dioxide-blasted implants. Clin Implant Dent Relat Res. 2005;7(1):36–42. Ronold HJ, Lyngstadaas SP, Ellingsen JE. Analysing the optimal value for titanium implant roughness in bone attachment using a tensile test. Biomaterials. 2003;24(25):4559–64. Rønold HJ, Ellingsen JE. Effect of micro-roughness produced by TiO2 blasting. Tensile testing of bone attachment by using coinshaped implants. Biomaterials. 2002;23(21):4211–9. Rønold HJ, Lyngstadaas S, Ellingsen JE. A study on the effect of dual blasting with TiO2 on titanium implant surfaces on functional attachment in bone. J Biomed Mater Res. 2003;67A(2): 524–30. Zhao G, Raines AL, Wieland M, Schwartz Z, Boyan BD. Requirement for both micron- and submicron scale structure for synergistic responses of osteoblasts to substrate surface energy and topography. Biomaterials. 2007;28(18):2821–9. Roccuzzo M, Bunino M, Prioglio F, Bianchi SD. Early loading of sandblasted and acid-etched (SLA) implants: a prospective splitmouth comparative study—one-year results. Clin Oral Implants Res. 2001;12(6):572–8. Cochran DL, Buser D, ten Bruggenkate CM, Weingart D, Taylor TM, Bernard JP, et al. The use of reduced healing times on ITI implants with a sandblasted and acid-etched (SLA) surface: early results from clinical trials on ITI SLA implants. Clin Oral Implants Res. 2002;13(2):144–53. Rupp F, Scheideler L, Olshanska N, de Wild M, Wieland M, Geis-Gerstorfer J. Enhancing surface free energy and hydrophilicity through chemical modification of microstructured titanium implant surfaces. J Biomed Mater Res A. 2006;76(2):323–34. Khang W, Feldman S, Hawley C, Gunsolley J. A multi-center study comparing dual acid-etched and machined-surfaced implants in various bone qualities. J Periodontol. 2001;72(10): 1384–90. Wen HB, Wolke JGC, de Wijn JR, Liu Q, Cui FZ, de Groot K. Fast precipitation of calcium phosphate layers on titanium induced by simple chemical treatments. Biomaterials. 1997; 18(22):1471–8. Wen HB, Liu Q, De Wijn J, De Groot K, Cui F. Preparation of bioactive microporous titanium surface by a new two-step chemical treatment. J Mater Sci Mater Med. 1998;9(3):121–8. Ellingsen JE, Lyngstadaas SP. Increasing biocompatibility by chemical modification of titanium surfaces. Bio-implant interface: improving biomaterials and tissue reactions. Florida: CRC Press; 2003. Lausmaa J. Mechanical, thermal, chemical and electrochemical surface treatment of titanium. Titanium in Medicine. New York: Springer; 2001. p. 231–66. Schneider GB, Perinpanayagam H, Clegg M, Zaharias R, Seabold D, Keller J, et al. Implant surface roughness affects osteoblast gene expression. J Dent Res. 2003;82(5):372.

J Mater Sci: Mater Med (2013) 24:2761–2773 28. Sul YT. The significance of the surface properties of oxidized titanium to the bone response: special emphasis on potential biochemical bonding of oxidized titanium implant. Biomaterials. 2003;24(22):3893–907. 29. Sul YT, Jeong Y, Johansson C, Albrektsson T. Oxidized, bioactive implants are rapidly and strongly integrated in bone. Part 1— experimental implants. Clin Oral Implants Res. 2006;17(5):521–6. 30. Cheng HC, Lee SY, Chen CC, Shyng YC, Ou KL. Influence of hydrogen charging on the formation of nanostructural titania by anodizing with cathodic pretreatment. J Electrochem Soc. 2007; 154(1):13–8. 31. Takadama H, Kim HM, Kokubo T, Nakamura T. An X-ray photoelectron spectroscopy study of the process of apatite formation on bioactive titanium metal. J Biomed Mater Res. 2001; 55(2):185–93. 32. Ellingsen JE, Lyngstadaas SP. Inventor Medical prosthetic devices having improved biocompatibility Patent US20080269910. 2008. 33. Shibata Y, Kawai H, Yamamoto H, Igarashi T, Miyazaki T. Antibacterial titanium plate anodized by being discharged in NaCl solution exhibits cell compatibility. J Dent Res. 2004; 83(2):115–9. 34. Reverchon E, Caputo G, Correra S, Cesti P. Synthesis of titanium hydroxide nanoparticles in supercritical carbon dioxide on the pilot scale. J Supercrit Fluids. 2003;26(3):253–61. 35. Ra¨mo¨ J. Hydrogen peroxide metals-chelating agents; interactions and analytical techniques. Oulu: University of Oulu; 2003. 36. Szmukler-Moncler S, Bischof M, Nedir R, Ermrich M. Titanium hydride and hydrogen concentration in acid-etched commercially pure titanium and titanium alloy implants: a comparative analysis of five implant systems. Clin Oral Implants Res. 2010;21(9):944–50. 37. Lamolle S, Monjo M, Lyngstadaas S, Ellingsen J, Haugen H. Titanium implant surface modification by cathodic reduction in hydrofluoric acid: surface characterization and in vivo performance. J Biomed Mater Res A. 2009;88A(3):581–8. 38. McCafferty E, Wightman J. Determination of the concentration of surface hydroxyl groups on metal oxide films by a quantitative XPS method. Surf Interface Anal. 1998;26(8):549–64. 39. Tugulu S, Lo¨we K, Scharnweber D, Schlottig F. Preparation of superhydrophilic microrough titanium implant surfaces by alkali treatment. J Mater Sci Mater Med. 2010;21(10):2751–63. 40. Lu X, Wang Y, Yang X, Zhang Q, Zhao Z, Weng LT, et al. Spectroscopic analysis of titanium surface functional groups under various surface modification and their behaviors in vitro and in vivo. J Biomed Mater Res A. 2008;84(2):523–34. 41. Beamson G, Briggs D. High resolution monochromated X-ray photoelectron spectroscopy of organic polymers: a comparison between solid state data for organic polymers and gas phase data for small molecules. Mol Phys. 1992;76(4):919–36. 42. Silversmit G, Poelman H, Depla D, Barrett N, Marin GB, De Gryse R. A comparative XPS and UPS study of VOx layers on mineral TiO2 (001)-anatase supports. Surf Interface Anal. 2006; 38(9):1257–65. 43. Lamolle SF, Monjo M, Rubert M, Haugen HJ, Lyngstadaas SP, Ellingsen JE. The effect of hydrofluoric acid treatment of

2773

44. 45.

46.

47.

48.

49.

50.

51.

52.

53.

54.

55.

56.

57.

58.

59.

titanium surface on nanostructural and chemical changes and the growth of MC3T3-E1 cells. Biomaterials. 2009;30(5):736–42. Cohen J. Statistical power analysis for the behavioral sciences. 2nd ed. Hillsdale: Lawrence Erlbaum Associates; 1988. Klokkevold PR, Nishimura RD, Adachi M, Caputo A. Osseointegration enhanced by chemical etching of the titanium surface. A torque removal study in the rabbit. Clin Oral Implants Res. 2002;8(6):442–7. Ferreira EA, Rocha-Filho RC, Biaggio SR, Bocchi N. Corrosion resistance of the Ti–50Zr at.% alloy after anodization in different acidic electrolytes. Corros Sci. 2010;52(12):4058–63. Pan J, Thierry D, Leygraf C. Electrochemical impedance spectroscopy study of the passive oxide film on titanium for implant application. Electrochim Acta. 1996;41(7–8):1143–53. Sul YT, Johansson CB, Jeong Y, Albrektsson T. The electrochemical oxide growth behaviour on titanium in acid and alkaline electrolytes. Med Eng Phys. 2001;23(5):329–46. Lausmaa J. Surface spectroscopic characterization of titanium implant materials. J Electron Spectros Relat Phenomena. 1996;81(3):343–61. Babelon P, Dequiedt A, Mostefa-Sba H, Bourgeois S, Sibillot P, Sacilotti M. SEM and XPS studies of titanium dioxide thin films grown by MOCVD. Thin Solid Films. 1998;322(1–2):63–7. Deligianni D, Katsala N, Ladas S, Sotiropoulou D, Amedee J, Missirlis Y. Effect of surface roughness of the titanium alloy Ti– 6Al–4V on human bone marrow cell response and on protein adsorption. Biomaterials. 2001;22(11):1241–51. Ohtsuka T, Nomura N. The dependence of the optical property of Ti anodic oxide film on its growth rate by ellipsometry. Corros Sci. 1997;39(7):1253–63. Feng B, Weng J, Yang BC, Qu SX, Zhang XD. Characterization of surface oxide films on titanium and adhesion of osteoblast. Biomaterials. 2003;24(25):4663–70. Martins DQ, Oso´rio WR, Souza MEP, Caram R, Garcia A. Effects of Zr content on microstructure and corrosion resistance of Ti–30Nb–Zr casting alloys for biomedical applications. Electrochim Acta. 2008;53(6):2809–17. Stein GS, Lian JB, Stein JL, Van Wijnen AJ, Montecino M. Transcriptional control of osteoblast growth and differentiation. Physiol Rev. 1996;76(2):593–629. Quarles LD, Yohay DA, Lever LW, Caton R, Wenstrup RJ. Distinct proliferative and differentiated stages of murine MC3T3E1 cells in culture: an vitro model of osteoblast development. J Bone Miner Res. 1992;7(6):683–92. Zhu X, Chen J, Scheideler L, Reichl R, Geis-Gerstorfer J. Effects of topography and composition of titanium surface oxides on osteoblast responses. Biomaterials. 2004;25(18):4087–103. Keller J, Stanford C, Wightman J, Draughn R, Zaharias R. Characterizations of titanium implant surfaces. III. J Biomed Mater Res. 1994;28(8):939–46. Zhu X, Chen J, Scheideler L, Altebaeumer T, Geis-Gerstorfer J, Kern D. Cellular reactions of osteoblasts to micron-and submicron-scale porous structures of titanium surfaces. Cells Tissues Organs. 2004;178(1):13–22.

123