J Mater Sci: Mater Med (2016)27:31 DOI 10.1007/s10856-015-5646-5

BIOMATERIALS SYNTHESIS AND CHARACTERIZATION

Original Research

Fabrication and characterization of gold nanoparticle-loaded TiO2 nanotube arrays for medical implants Yu Bai1 • Yulong Bai1 • Cunyang Wang1 • Jingjun Gao1 • Wen Ma1

Received: 22 October 2015 / Accepted: 7 December 2015 Ó Springer Science+Business Media New York 2015

Abstract Au nanoparticles (AuNPs) are successfully assembled on TiO2 nanotube (TN) arrays through electrochemical deposition technology to improve the surface characteristics of TN arrays as an implant material. The loading amount of AuNPs can be controlled by adjusting the deposition time of electrochemical deposition. The effect of the amount of the loaded AuNPs on surface roughness and surface energy is systematically investigated on the basis of various characterizations. Results show that the increase in the loading amount of AuNPs on the TN arrays can increase surface roughness and decrease surface energy. Potentiodynamic polarization tests indicate that AuNP-modified TNs possess a higher corrosion resistance than unmodified TNs. Corrosion resistance increases as the amount of the loaded AuNP increases. In vitro cell culture tests are performed on the basis of cell morphology observations and MTT assays. Osteoblast cell adhesion and proliferation ability on the AuNP-modified TN surface are greater than those on the unmodified TN surface. The sample fabricated at the deposition time of 90 s exhibits an optimum cell performance. This work can provide a new platform to develop the surface chemistry of TN arrays and to fabricate titanium-based implant materials to enhance bioactivity.

& Yu Bai [email protected] 1

School of Materials Science and Engineering, Inner Mongolia University of Technology, Hohhot 010051, Inner Mongolia, China

1 Introduction Titanium (Ti) has been extensively used to develop orthopedic or dental implants because of its excellent biocompatibility, mechanical properties, corrosion resistance, and low biotoxicity in vivo [1]. However, the integration of Ti implants in bones remains unsatisfied because of its bioinertness [2]. Therefore, surface treatment methods, such as micro-arc oxidation, ion implantation, anodization, alkali and heat treatments, and electrochemical deposition, have been applied. Excellent surface characteristics can be obtained by improving their wettability, surface roughness, corrosion resistance, and bioactivity through these treatments [3–8]. Thus far, studies have been conducted to form self-organized TNs on Ti implant surface through anodization [9, 10]. Compared with other modification methods, electrochemical anodization is considered as an optimum approach because the nanoscale form of tubes can be easily controlled by adjusting voltage, current, time, temperature, and electrolyte components [11–15]. The TN formation can be applied to enhance bioactivity [16, 17]. Schneider et al. reported that the adhesion and proliferation of osteoblasts on TN arrays are significantly enhanced [18]. Oh et al. also demonstrated that nanotubular TiO2 surfaces can be used for bone implants that favor bone cell growth, cell differentiation, and apatite growth compared with a smooth controlled Ti surface [19]. However, modification can be carried out by adsorbing or loading some bioactive materials on the surface of TN arrays to improve the functionality of TN arrays for practical bioapplications. For example, Zhao et al. revealed that osteogenic activity has been upregulated by strontium-loaded TN arrays on a Ti substrate [20]. Indira et al. incorporated zirconium (Zr) ions onto TN arrays by a simple dip-coating method for biomedical implants [21].

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They found that Zr–TNs possess higher bioactivity and corrosion resistance than other control samples. Gold nanoparticles (AuNPs) have been extensively used to manipulate cell behavior because these particles can be easily synthesized and can be versatile in surface functionalization [22]. Nevertheless, the potential toxicity of AuNPs has been greatly considered. Nanoparticle size plays a critical role in the rate and extent of cellular uptake. Conner et al. investigated the toxicity of AuNPs with different sizes, such as 4, 12, and 18 nm in diameter, on human leukemia cell line [23]. They conducted MTT assays and found that nanoparticles are not cytotoxic. Similar results have been obtained when 3.5 nm gold nanospheres are used on immune system cell lines. AuNPs enter cells through endocytosis without inducing toxicity; these AuNPs also reduce the level of reactive oxygen species [24]. Pan et al. reported that 1.4 nm AuNPs trigger necrosis, mitochondrial damage, and oxidative stress on the examined cell lines [25]. Interestingly, they did not find any evidence of cellular damage caused by 15 nm AuNPs bearing the same surface group. These results indicate the possible size-dependent toxicity of AuNPs. In general, AuNPs with a larger diameter ([5 nm) are considered nontoxic to cellular functions and are chemically inert. Furthermore, the chemical reactivity of AuNPs with a diameter of\3 nm is very different from that of larger AuNPs [26]. Several methods, such as magnetron sputtering, photoreduction, and lyophilization, have been applied to load AuNPs on supported TiO2. Jun et al. loaded the AuNPs on the surface of TN arrays through direct current magnetic sputtering technology and investigated their opto-electronic response under UV light [27]. Although this approach can result in the homogeneous dispersion of nanoparticles on TN arrays, the direct current magnetic sputtering technology requires a specific instrument and a spectrally pure Au target, which is very expensive. Neupane et al. employed lyophilization combined with vacuum drying method to deposit gelatin-stabilized AuNPs on TNs. However, AuNPs generally enlarge and easily aggregate on surfaces [28]. Electrochemical deposition can be applied to fabricate AuNPs loaded on TN arrays because of its simplicity, low equipment costs, and convenient control mechanisms to obtain a homogeneous surface appearance. In this process, Au3? ions from a HAuCl4 precursor can be reduced to metallic Au; metallic Au can be subsequently deposited onto a nanotube surface. In our work, 30 nm AuNP-loaded TN arrays were prepared through anodization combined with electrochemical deposition. The effects of the loading amount of AuNPs on surface morphology, wettability, roughness, and corrosion resistance were explored. The relationship between implant surface and osteoblastic behavior was also investigated systematically.

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2 Materials and methods 2.1 Sample preparation Commercially available pure Ti plates with a size of 20 mm 9 10 mm 9 1 mm were prepared and polished with #400–1200 SiC paper. Before anodization was performed, the plate was cleaned by ultrasonically washing with ethanol and deionized water (DW) for 10 min; the plate was etched in a HF:HNO3:H2O solution with a volume ratio of 1:1.5:6 for 10 s, washed with DW, and dried at 40 °C. 2.2 Fabrication of TN arrays Anodization was performed at 20 V for 40 min by connecting the Ti plate and a platinum foil to the anode and the cathode of a DC power supply, respectively. The electrolyte solution was prepared by mixing 79 wt% glycerol with 20 wt% DW and 1 wt% NH4F. After anodization was completed, the sample was ultrasonically washed with DW for 10 s and dried at room temperature. In the heat treatment, the as-prepared TNs were heated up to 500 °C at a rate of 5 °C/min in the electric furnace (SX-2-1300, Test Electric Co., Shanghai, China), maintained at this temperature for 2 h, and cooled to room temperature. 2.3 Preparation and characterization of AuNPloaded TN arrays The AuNP-loaded TN arrays were prepared through potentiostatic electrodeposition by using an electrochemistry station (Zennium, Zahner Co., Germany). A three-electrode setup was applied with TNs, a platinum sheet, and a saturated calomel electrode (SCE) as working, counter, and reference electrodes, respectively. AuNPs were deposited onto the crystallized TN arrays in the HAuCl4 (1.5 mmol/L) aqueous solution at the applied potential of -0.9 V. The loading amounts of AuNPs were controlled by adjusting the deposition time. The deposition time setting was 5, 15, 30, 60, 90, and 120 s. The samples were denoted correspondingly as xAuNP-TNs, where x = 5, 15, 30, 60, 90, 120). After deposition was completed, the samples were washed with DW and dried at room temperature. The morphological characteristics and components of the samples were observed through field emission scanning electron microscopy (FE-SEM, Quanta, FEG-650) with an energy dispersive spectroscopy (EDS). In the cross-sectional view, the back of the sample was scratched deeply with a diamond cutter. The crystalline structure was identified using an X-ray diffractometer (XRD, Dmax-2500/ PC, Rigaku Co., Japan) with Cu Ka incident radiation, a tube voltage of 40 kV, and a current of 100 mA. The

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scanning angle ranged from 20° to 70°, with a scanning rate of 3°/min. The arithmetical mean roughness (Ra) measurements of the samples were quantified using a confocal laser scanning microscope (CLSM, Carl Zeiss Jena, LSM-700); with this instrument, the surface roughness was measured five times along a 2 mm length on 9500 and at three different positions of each sample. 2.4 Corrosion resistance Electrochemical investigations were performed using an electrochemistry station to evaluate the corrosion resistance ability of the samples in a simulated body fluid (SBF). A conventional three-electrode cell was utilized for electrochemical measurements with a platinum sheet, SCE, and sample as counter, reference, and working electrodes, respectively. The SBF was produced by adding 0.185 g/L calcium chloride dihydrate, 0.0977 g/L magnesium sulfate, and 0.350 g/L sodium hydrogen carbonate to Hanks’ balanced solution (H2387, Sigma Chemical Co., USA). The pH of the SBF was controlled at 7.4. The potentiodynamic polarization measurements were performed at a scan rate of 4 mV/s. Corrosion current density (Icorr) and corrosion potential (Ecorr) were obtained by plotting the linear anodic and cathodic curves on the basis of the classical Tafel analysis. Furthermore, the corrosion rate (VL) can be calculated using the following equation [29]: VL ¼ 3:27  103 

A  Icorr nq

ð1Þ

where A, n, and q represent metal relative atomic mass, metal valence, and metal density in g/cm3, respectively. 2.5 Contact angles and surface energy Static contact angles were detected via a liquid drop method through contact angle measurement (JC2000DM, Beijing Zhongyikexin Technology Co., China) at room temperature. Two different liquids, namely, DW and glycol, were used to carry out the measurements. A droplet of the liquid (5 lL) was applied to the surface of the sample, and an image of the droplet was captured immediately after stabilization was achieved. The droplet profile was automatically fitted and analyzed with the software supplied by the manufacturer. At least three samples were used in each group, and different points were selected in each sample. The surface energies of the samples were calculated using ‘‘Owens model,’’ which employs the long-range dispersion (Lifshitz–van der Waals, cd) and short-range polar (hydrogen bonding, cp) components of surface energy. The polar (cp) and dispersive (cd) components of the surface energy of each sample surface can be obtained using the following equations:

cL ð1 þ cos hÞ ¼ 2ð

qffiffiffiffiffiffiffiffiffi qffiffiffiffiffiffiffiffiffi cdS cdL þ cPS cPL Þ

c ¼ cP þ cd

31

ð2Þ ð3Þ

where cL and cS correspond to the liquid surface energy and solid surface energy, respectively. Two different liquids, namely, DW and glycol, with known values of the polar and dispersive components of the surface energy were used in this study. 2.6 Cell culture The mouse osteoblast cells (MC3T3-E1) used in this study were obtained from Procell Ltd. (Wuhan, China). The culture medium was prepared by adding 10 % fetal bovine serum (FBS; Gibco Co., USA) and 500 unit/mL penicillin– streptomycin (Procell Ltd., China) into a-MEM (Gibco Co., USA). The cultured cells were seeded on the samples with an initial density of 5 9 104/cm2. The cells were incubated at 37 °C in a humidified 5 vol. % CO2 atmosphere. 2.7 Cell morphology The cells cultured for 1 day were washed with phosphate buffer solutions (PBS) to remove non-adherent cells, and adherent cells were fixed with 2.5 % glutaraldehyde (Sigma-Aldrich, USA) in PBS solution for 2 h at 4 °C and post-fixed in 1.0 % osmium tetroxide (Sigma-Aldrich, USA) for 2 h at 4 °C. The samples were subsequently washed twice with PBS, dehydrated in a graded ethanol series (30, 50, 70, 80, 90, 95, and 100 % v/v) for 15 min each, and dried at room temperature. The dried samples were examined using SEM. 2.8 Cell viability assay Cell viability was determined by methylthiazole tetrazolium (MTT) assay after the cells were cultured for 2 and 4 days. Triplicate samples per group were evaluated. aMEM (18 mL) without FBS was mixed with 2 mL of 5 mg mL-1 MTT solution. The MTT mixture (1 mL) was used in each sample in 24-well plates and incubated for 4 h. The solution was then aspirated, and the dark blue crystals remaining in the wells were dissolved in 1 mL of dimethyl sulfoxide (Sigma-Aldrich, USA). Afterward, 100 lL of the solution was transferred to a new 96-well plate, and five data points were obtained from each sample. The optical density of the solution in each well was measured at a wavelength of 570 nm by using a microplate reader (Epoch 2, BioTek, USA).

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2.9 Data analysis

Au3þ þ 3e ! Au

All experiments were carried out in triplicate and were expressed as the mean ± standard deviation (SD). Single-factor analysis of variance was used to assess the statistical significance of the results. A P value \ 0.05 was considered significant.

3 Results and discussion Figure 1 shows the SEM images of the surface and cross section of TN arrays fabricated at 20 V in NH4F/glycerol/ H2O electrolyte solution for 40 min. The TNs with a diameter of 80 nm and a length of 500 nm were uniformly arranged on the Ti substrate. The TN arrays were strongly adherent to Ti metal base because the nanotube layer was very difficult to be removed by scraping or bending of the Ti substrate [30]. Figure 2 shows the surface morphological characteristics of the AuNP-TNs observed through SEM. AuNPs were electrochemically deposited on the TN arrays at -0.9 V with a duration of (a) 5, (b) 15, (c) 30, (d) 60, (e) 90, and (f) 120 s in 1.5 mmol/L HAuCl4 solution. The AuNPs with a size of approximately 30 nm are homogeneously distributed on the surface of the TN arrays. The nanoparticles preferentially nucleate at the boundaries between TNs, and this phenomenon may be related to high current densities. AuNPs are not found on pure Ti sheet without TNs. This result is consistent with that found by Yang et al. [31] who proposed that highly uniform TN arrays provide a homogeneous electrodepositing environment for the nucleation and growth of Au crystals. In an electrolyte solution, [AuCl4]- exists in an electroionization reaction described as follows: ½AuCl4  Au3þ þ 4Cl

ð4Þ

Au3? near the TN electrode can obtain electrons and is thus reduced into Au; this phenomenon can be depicted as Eq. (5).

ð5Þ

concentration around the TN electrode As the Au decreases, a concentration gradient between the bulk solution and the TN surface is produced. Under the concentration gradient force, numerous Au3? move toward the TN electrode. As a consequence, more reduced Au crystals form on the surface of TN arrays. At a specific deposition potential, the loading amount of AuNPs increases as duration time is prolonged. Table 1 shows the components of the surface through EDS analysis in three different places; the result revealed that the loading amount of AuNP gradually increases as the deposition time is extended. At 5 s, the loading amount of AuNPs is 1.87 wt%. When the time is increased to 120 s. 4.76 wt% AuNPs are deposited on the TN arrays. Although the loading amount increases in the prolonged duration time, the size of AuNPs does not have evidently change (Fig. 2); this finding indicates that AuNP does not agglomerate. Figure 3 illustrates the crystalline analysis results of the samples corresponding to Ti, TNs, annealed TNs, and AuNP-TNs. Curve (a) shows the Ti metal phase (JCPDS No. 01-1198). Curve (b) reveals the Ti peaks observed on the as-formed sample; this finding indicates that the nanotubes were amorphous before annealing occurred. Except the Ti peaks, some distinct peaks can be found on the XRD spectrum of the annealed TN arrays [Curve (c)]. According to the JCPDS card (No: 21-1272), these peaks are well indexed to the anatase phase of TiO2. The anatase phase with nanoscale topography elicits a good biological effect on cell adhesion, spreading, proliferation, and differentiation [32]. Curves (d–i) correspond to the XRD patterns of AuNP-loaded TN arrays fabricated at the duration times of 5, 15, 30, 60, 90, and 120 s, respectively. In addition to the characteristic peaks assigned to Ti and TiO2 (anatase), the XRD pattern in Curves (g–i) exhibit sharp reflections or characteristic peaks, which are indexed to the cubic Au (JCPDS No. 04-0784); this observation reveals that Au crystals are formed. The XRD peak intensity of Au strengthens as duration time is prolonged; this finding

Fig. 1 SEM images of the surface (a) and the cross section (b) of TN arrays

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Fig. 2 SEM images of a 5AuNP-TNs, b 15AuNP-TNs, c 30AuNP-TNs, d 60AuNP-TNs, e 90AuNP-TNs, and f 120AuNP-TNs

Table 1 Elemental composition of AuNP-TN samples with different loading amount of AuNPs (wt%)

Element

5AuNPs

15AuNPs

30AuNPs

60AuNPs

90AuNPs

120AuNPs

O

32.33

33.85

32.64

32.21

29.84

29.01

Ti

65.80

63.58

64.50

64.84

66.10

66.23

Au

1.87

2.57

2.86

2.95

4.06

4.76

Totals

100.00

100.00

100.00

100.00

100.00

100.00

suggests that the loading amount of Au increases on the TN arrays. Furthermore, the Au peak cannot be observed on 5AuNP-TNs [curve (d)], 15AuNP-TNs [curve (e)], and 30AuNP-TNs [curve (f)], which is probably attributed to the low loading amount of AuNPs on the TN arrays. In conductive human body fluids, electrochemical corrosion is a dominant form interacting with metals. The

potentiodynamic polarization curves of pure Ti, annealed TN, and AuNP–TN samples in SBF are evaluated to investigate the effect of AuNPs on the corrosion behavior of TN arrays, and the typical curves are shown in Fig. 4. Ecorr and Icorr extracted from the Tafel polarization curves are summarized in Table 2. The annealed TN arrays exhibit a higher Ecorr than the pure Ti plate. On the basis of

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T

T

T: Titanium A: Anatase

T T

A

T

Au

Relative intensity (a.u.)

Au A

i h g f e d c b a

20

30

40

50

60

70

2θ / degree

Fig. 3 XRD patterns of different samples: a pure Ti, b TNs, c annealed TNs d 5AuNP-TNs, e 15AuNP-TNs, f 30AuNP-TNs, g 60AuNP-TNs, h 90AuNP-TNs and i 120AuNP-TNs

1.2

h Potential (V vs SCE)

0.8 0.4

f g i

0.0

c

d e

-0.4 -0.8

a

b

-1.2 -1.6 -10

-9

-8

-7

-6

-5

-4

-3

-2

log i (A/cm2)

Fig. 4 Potentiodynamic polarization plot recorded for a pure Ti, b TNs, c annealed TNs, d 5AuNP-TNs, e 15AuNP-TNs, f 30AuNPTNs, g 60AuNP-TNs, h 90AuNP-TNs and i 120AuNP-TNs

Table 2 Corrosion parameters of different samples

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the XRD result, we found that the annealed TN arrays are transferred to an anatase phase through annealing. The crystallized anatase phase is more efficiently resistant to corrosion than the amorphous phase. After the AuNPs were loaded on the annealed TN arrays, Ecorr of the samples is further enhanced compared with pure Ti and TN arrays. With the AuNP–TN samples, a general trend that Ecorr increases as the loading amount of AuNP increases. Notably, a high Ecorr is beneficial to improve corrosion resistance. Icorr can be generally used to characterize the corrosion rate, and a low Icorr indicates a slow corrosion rate of the sample; this observation may include localized corrosion and uniform corrosion. In pitting corrosion, Icorr is a measure of the corrosion rate within the pit and thus a measure of the pit penetration rate. A decrease in Icorr delays corrosion in the pit and decreases the ionic concentration in the solution. The corrosion rate calculated from Icorr is also listed in Table 2. The 120AuNP–TN sample yields the lowest corrosion rate in SBF (0.2384 9 10-3 mm/a). The superior corrosion resistance of the AuNP–TN sample can be explained by the homogeneous AuNP distribution on the surface of the TN arrays and the inherent inertness of Au. Oliveira et al. showed that the incorporation of gold oxide nanoparticles into diamond-like carbon films has remarkably improved the electrochemical corrosive resistance of aggressive ions toward the surface of these films [33]. Ahmed et al. demonstrated that the addition of AuNPs to chitosanforming biocomposite coat on the NiTi alloy greatly enhances the corrosion resistance in Hanks’ physiological solution [34]. This behavior may be explained by considering the distinct physical and chemical properties of Au and combined effect between AuNPs and chitosan. Figure 5 shows the roughness value (Ra) from the investigated samples. It is seen that the annealed TN arrays show a higher Ra value than the pure Ti plate. The Ra value of Ti and annealed TN arrays is 0.103 and 0.162 lm, respectively. The change in roughness should be closely

Sample

Icorr (lA/cm2)

Ecorr (V vs SCE)

Corrosion rate (mm/a)

Pure Ti

24.0438

-0.8595

0.2070

TNs

19.7424

-0.7886

0.1699

Annealed TNs

1.9759

-0.1310

0.0170

5AuNP-TNs

0.2554

-0.2347

2.1982 9 10-3

15AuNP-TNs

0.1820

-0.2614

1.5665 9 10-3

30AuNP-TNs

0.1542

0.1676

1.3272 9 10-3

60AuNP-TNs

0.1390

0.1675

1.1964 9 10-3

90AuNP-TNs

0.0867

0.2334

0.7462 9 10-3

120AuNP-TNs

0.0277

0.2558

0.2384 9 10-3

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related to anodization and annealing with the anatase phase formation. The average Ra values of 5AuNP-TNs, 15AuNP-TNs, 30AuNP-TNs, 60AuNP-TNs, 90AuNPTNs, and 120AuNP-TNs are 0.294, 0.354, 0.412, 0.471, 0.503, and 0.537 lm, respectively. After the AuNPs are loaded on the TN arrays, the Ra value further increases. The longer duration time or higher Au loading amount also increase the Ra value. The cell response to the surface morphology has been extensively investigated, and the effects of roughness on osseointegration and biomechanical stability have been described in detail [35, 36]. In conclusion, osteoblastic cells likely attach to rough surfaces as the initial stability increases. Figure 6 shows the images of water contact angle on the surfaces of Ti plate, TN arrays, and AuNP–TN arrays.

0.6

Ti

0.1

120AuNP-TNs

90AuNP-TNs

60AuNP-TNs

5AuNP-TNs

* 0.2

30AuNP-TNs

0.3

15AuNP-TNs

0.4

TNs

Surface roughness (μm)

0.5

0.0

Fig. 5 Surface roughness of different samples. Significant difference *P \ 0.05 compared with pure Ti

Fig. 6 Water contact angle images of different samples

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Contact angle values differ from one another; these values range from 0° to 37°. On the surface of pure Ti, the contact angle is 37°, whereas the contact angle evidently decreases on the TN surface. The detailed value is 25°. The variation in contact angles between pure Ti and TN arrays is attributed primarily to their different surface natures. The nanotubular structure of TNs provides a larger surface area and higher porosity; thus water easily infuses into the pores. The contact angles of the AuNP–TN samples are lower than those of the TN sample without AuNPs; this observation may be ascribed to their large surface roughness. Surface roughness is a factor influencing surface wettability. In general, higher roughness value results in enhanced wettability, that is, a lower contact angle. Among the AuNP–TN samples, the water contact angles of 5AuNP-TNs, 15AuNP-TNs, 30AuNP-TNs, 60AuNP-TNs, 90AuNP-TNs, and 120AuNP-TNs are 0°, 5°, 8°, 10°, 11°, and 15°, respectively. The changing trend of the water contact angle increases as the amount of the loaded AuNPs on TN arrays increases; this finding is particularly observed in 90AuNP–TN and 120AuNP–TN sample; this observation is very likely attributed to the increase in the loading amount of Au. As a result, the TN porosity decreases; thus, the contact angle increases. In the electrochemical process, the Au crystals preferentially nucleate and grow at the boundaries between TNs; however, the AuNPs possibly enter the pores of TN arrays as the duration time is extended. When AuNPs cover the pores of TN arrays, water cannot easily infuse into the pores; consequently, the surface wettability decreases. Therefore, differences in contact angle or surface wettability is likely related to the common effect of porosity and surface roughness. It is known that the surface energy has a significant effect on protein adsorption and cell behaviors. Among the different approaches for determination of the solid surface energy, the surface energy component approach is widely used, where the solid surface energy is expressed as a sum of a dispersive surface energy and a polar surface energy. The dispersive surface energy results from molecular interaction due to London forces, while the polar surface energy comprises all other interactions due to non-London forces. Nyilas et al. evaluated the effect of polymer surface molecular structure on blood interfacial phenomena [37]. They found that the primary determinant of surface—induced effects was the polar force contribution. Similarly, both Zhao et al. and Ponsonnet et al. studied the relationship between cell behaviors and the polar component of surface energy [38, 39]. It was shown that the polar component was a major factor affecting cell proliferation. The smaller the polar component, the higher the cell adhesion and proliferation. Motivated by these observations, we calculated the surface energy of different samples and meanwhile separated them into the dispersive and polar

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120 Dispersive component Polar component

110

0.5

0.4

90

*

*

Ti

120AuNP-TNs

30

TNs

90AuNP-TNs

40

60AuNP-TNs

50

30AuNP-TNs

60

15AuNP-TNs

70

20 10 0

Fig. 7 Surface energy of pure Ti, TN and AuNP-TN samples

components using ‘‘Owens model’’. The result is shown in Fig. 7; the polar and dispersive components of surface energy are also clearly visible in Fig. 7. Among the samples, the pure Ti sample exhibits the smallest polar component. The polar component increases after the AuNPs are loaded on TN arrays. The polar component likely attracts the electric dipoles of water. A large polar component corresponds to a small water contact angle. Furthermore, the total surface energy is similar among AuNP–TN samples. The slight differences are possibly attributed to the difference in the polar component, and the dispersive component is exactly the same in all of the AuNP–TN samples. In vitro cell tests are performed to investigate the biocompatibility and bioactivity of the AuNP–TN sample as an implant material. An MTT assay is used to determine the proliferating osteoblast cells and viable cells on pure Ti, TN, and AuNP–TN samples. Figure 8 shows the comparison of the cell densities on the different samples for 2 and 4 days. The optical density of the samples increases as culture time is extended; this finding indicates that cytotoxic effects are not elicited. The number of the cells attached to the AuNP–TN sample is higher than that of the cells on the pure Ti and TN samples because of the synergetic effect of the TN arrays and AuNPs. Previous studies revealed that nanotubular TiO2 surfaces can promote cell adhesion, proliferation, and differentiation. Au is also a well-known biometallic element; thus, the AuNP loading onto TN arrays is favorable for cell behavior. Furthermore, 90AuNP–TN shows a slightly higher cell proliferation ability than the other AuNP-modified TN arrays probably because of structure cues. Cell attachment and proliferation are primarily associated with material surface characteristics, such as surface chemistry, morphology, wettability, and surface energy. The difference in the loading amount of AuNPs on TN arrays causes the variations of surface roughness and surface energy; as a

123

OD570nm

80

5AuNP-TNs

Surface energy (mJ/cm2)

100

Pure Ti TNs 5AuNP-TNs 15AuNP-TNs 30AuNP-TNs 60AuNP-TNs 90AuNP-TNs 120AuNP-TNs

0.3

0.2

0.1

0.0

Day 2

Culture time

Day 4

Fig. 8 MTT assay of osteoblast cells on different sample surfaces after culture of 2 and 4 days. Significant difference *P \ 0.05 compared with other samples

consequence, cell proliferation ability differs. The surface roughness is increased by loading AuNPs on TN arrays. Nanoscale surface morphology is observed to promote the adhesion and proliferation of osteoblasts and enhance osseointegration [40]. The AuNPs on TN surfaces provide more adhesion points for proteins to attach to and thus contribute to cell adhesion. Although surface roughness increases as the loading amount of AuNPs on TN arrays increases, the highest cell proliferation ability is not observed on the 120AuNP–TN sample. This observation is possible because the surface energy is also an important factor influencing cell behavior despite surface roughness. In particularly, the polar component of surface energy is closely related to cell adhesion and proliferation. The polar component likely attracts the electric dipoles of water, and this phenomenon represents the interaction of the surface with water. A larger polar component indicates that the surface can adsorb more cell-adhering proteins, such as fibronectin and vitronectin, from serum solution and consequently promote cell attachment and proliferation. However, a contradictory result reveals that the smaller surface energy of the surface is partially responsible for less protein denaturation and cell behavior enhancement. In conclusion, surface energy is an important factor, but roughness can strongly disrupt the relationships between surface energy and cell proliferation. Indeed, the cell performance on different samples is mainly attributed to the common effect of surface roughness and surface energy. Figure 9 shows the morphological characteristics of the osteoblast cells onto the different samples after these cells were cultured for 1 day. The seeded cells adhere to the samples and are well expanded on the material surface. This observation suggests that all of the prepared samples provide a biocompatible environment favorable for cell

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attachment. A typically elongated morphology is observed on the 90AuNP-TN surfaces covered with cells with many filopodium extensions from the cell to the substrate

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(Fig. 10). This result indicates that the interplay between the cell and the AuNP–TN surface can enhance dynamic propagation and osteoblast activation.

Fig. 9 SEM images of the cellular morphology on the surface of a pure Ti, b TNs, c 5AuNP-TNs, d 15AuNP-TNs, e 30AuNP-TNs, f 60AuNPTNs, g 90AuNP-TNs, h 120AuNP-TNs after culture of 1 day

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Fig. 10 SEM images of osteoblast cell filopodia interacting with a 90AuNP-TNs and b TN substrate (control sample) after culture of 1 day

4 Conclusions The surfaces of TNs are modified with AuNPs through a simple electrochemical deposition technology to improve the benefits of TNs as an implant material. SEM, EDS, and XRD results demonstrate that the AuNPs are successfully loaded onto the TN arrays. The AuNP-modified TN arrays exhibit an increased corrosion resistance, surface roughness, and surface energy compared with TNs; this result is mainly attributed to the synergistic effect of TNs and AuNPs. In vitro cell culture test reveals that the osteoblast cell adhesion and proliferation ability are promoted to a greater extent on AuNP-TNs than on TNs; this finding can be attributed to the inherent biocompatibility of the loaded AuNPs and possibly to the pathways of fluid between nanotubes. In addition, the loading amount of AuNPs is an important factor affecting the surface characteristics of AuNP-TNs and their subsequent cell behavior. Among the investigated samples, 90AuNP-TNs show the highest osteoblast proliferation ability, as revealed by the MTT assay. The novel AuNP-modified TNs on a Ti implant surface can be used as an excellent bioactive surface for orthopedic and dental materials. Acknowledgments This work was financially supported by ‘‘The Natural Science Foundation of Inner Mongolia Autonomous Region (2014BS0504)’’ funded by the Technology Department of Inner Mongolia Autonomous Region and by ‘‘The Scientific Research Foundation of the Education Ministry for Returned Chinese Scholars ([2013]693)’’ funded by the Department of International Exchange and Cooperation of the Ministry of Education. Additionally, the work was also funded by the Key Laboratory of Inorganic Coating Materials, Chinese Academy of Sciences.

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