chapter 2 plasma surface modification of nickel titanium ...abstract: nickel-titanium (niti) shape...

Advances in Biomedical Sciences and Engineering, 45-68 45

S C Tjong (Ed.) All rights reserved - © 2009 Bentham Science Publishers Ltd.

CHAPTER 2 Plasma Surface Modification of Nickel Titanium Shape Memory Alloys Kelvin Wai-Kwok Yeung1,*, Kenneth MC Cheung1, Keith Dip-Kei Luk1, Shuilin Wu2, Xuanyong Liu2, Chenglin Chu2, Chi Yuan Chung2, Paul Kim-Ho Chu2

1Division of Spine Surgery, Department of Orthopedics & Traumatology, The University of Hong Kong, Hong Kong 2Department of Physics and Materials Science, City University of Hong Kong, Hong Kong *Corresponding author; E-mail: [email protected] Abstract: Nickel-titanium (NiTi) shape memory alloys are very potential for surgical implantation due to two unique properties: super-elasticity and shape memory effect. These advantages cannot be seen in current biomedical metallic materials such as medical grade titanium alloys and stainless steels. However, nickel ion release remains a major concern particularly for large implants placed on the spine or joints, as fretting is always expected at such implant junctions. Therefore, an advanced surface treatment using plasma immersion ion implantation (PIII) technology to tackle this issue

BODY RESPONSE TO NICKEL TITANIUM SHAPE MEMORY ALLOYS Nickel titanium (NiTi) shape memory alloys are promising materials for surgical implants in orthopedics due to their unique shape memory effects (SME) and super-elasticity (SE) that other common orthopedic materials such as stainless steels and titanium alloys do not possess. Their bulk mechanical properties are also closer to those of human cortical bones than that of 316L stainless steels and titanium alloys. In terms of wear resistance, the materials are better than CoCrMo alloys used in bone trauma fixation [1]. The high corrosion resistance, excellent biocompatibility, and cytocompatibility of NiTi satisfy the criteria required by biomedical applications [2-14]. The biocompatibility of NiTi alloys has been widely investigated. Studies have confirmed that NiTi alloy has excellent biocompatibility similar to medical grade 316L stainless steel [1, 5, 15-27]. Although nickel is carcinogenic, its properties alter when it is formed into an alloy with titanium. Nickel is chemically bonded to titanium via very strong intermetallic bonds thereby drastically reducing the possibility of nickel liberation. Biocompatibility studies in rabbits have demonstrated the materials safety and Miyazaki et al. [12] have reported that minimal amounts of nickel are released around the tissues and that new bone formation is observed at the surface of NiTi implants. Putters et al. [19] have studied the effects of increased exposure to NiTi, nickel or titanium in cell cultures. The results reveal that nickel inhibits mitosis in human fibroblasts, but NiTi and titanium do not. Filip et al. [16] have report that the histological evaluation of the implant and adjacent bone tissue interface next to NiTi alloy implant shows optimal bio-compatibility to bone tissues.

46 Advances in Biomedical Sciences and Engineering Yeung et al.

CONTROVERSAL ISSUES REGARDING THE USE OF NICKEL TITANIUM ALLOY IN MEDICINE Despite the good biocompatibility of NiTi alloys, some negative effects have also been pointed out. For example, Berger-Gorbet et al. have found that the process of osteogenesis and osteonectin synthesis activity in NiTi alloys are unfavorable compared to stainless steels and titanium alloys [28]. Castleman and Motzkin [29] have evaluated the acceptance of NiTi in vitro and found that NiTi and titanium would significantly reduce cell growth based on control cultures, but 316L stainless steel and Co-Cr alloy do not. Jia et al. [30] have reported that the cell death rate is severe in the presence of NiTi alloys. These findings appear to contradict a study in which titanium is recommended for use in vivo and in vitro [31]. These problems are believed to stem from the poor corrosion resistance of the materials thereby leading to an increase in the cytotoxicity. It is most likely that some toxic components released from the substrate cause the cell death rather than apoptosis [32]. Shih et al. [33] have reported that the supernatant and corrosive products from NiTi may result in the death of smooth muscle cells, especially when the amount of released nickel is higher than 9 ppm. A few other studies have reported that nickel ions [20, 21] leached from the alloys cause allergic reactions in nickel hyper-sensitive patients [34-37]. While the homogeneity of the materials microstructures and the surface morphology may alter the anti-corrosion ability of NiTi alloys, there is no doubt that the corrosion resistance and anti-wear properties of the materials must be enhanced before the materials can be widely used clinically, especially as orthopedic implants with couplings where fretting is expected.

SUEFACE TREATMENT OF NICKEL TITANIUM ALLOY A number of surface coatings or surface modification schemes have been studied to enhance the corrosion resistance as well as the biocompatibility of nickel titanium alloy. For instance, Liu et al. [38] have reported that deposition of a TiO2 film on NiTi alloys using the sol-gel technique improves the anti-corrosion properties. However, the surface mechanical properties have not been reported. Therefore, the practicality of the use of a coating is questionable if the materials are to be used in orthopedic implants where fretting and abrasive wear are expected. Alternatively, Villermaux et al. [39] have studied the use of excimer laser surface treatment to improve the anti corrosion properties of NiTi alloys. This method has been shown to reduce the number of corrosion pits and sizes and increase the thickness of the oxide layer. However, the biological effects of the materials have not been evaluated.

PLASMA IMMERSION ION IMPLANTATION OF BIOMATERIALS In the early 1980’s, plasma immersion ion implantation (PIII) (Figure 1) was first introduced. This surface coating technique has a number of advantages such as high

Plasma Surface Modification of Nickel Titanium Shape….. Advances in Biomedical Sciences and Engineering 47

efficiency, large coating area, batch processing, and small instrument footprint, as compared to conventional beam-line ion implantation [40].

Fig. 1. Schematic diagram of plasma immersion ion implantation setup. The use of this advanced technology in biomaterials has been widely investigated. One of the major benefits is to improve the surface of less biocompatible materials to make them more compatible with living tissues without altering the favorable properties of the substrates [41]. This treatment provides a convenient, economical, and efficient way to generate a biocompatible artificial surface on conventional materials rather than synthesize new materials with comparable biocompatibility. For instance, the hemocompatibility and mechanical durability of the current artificial heart valves coated with low-temperature isotropic carbon (LTIC) thin film for cardiac valve transplantation is far from adequate. Therefore, patients require administration of anticoagulants to prevent embolism. These valves may also fail due to the brittleness of the LTIC coating. Recently, PIII has been applied to remedy those defects in artificial heart valves. Fabrication of TiN, Ti(Ta15)O2, and diamond-like carbon (DLC) thin films on the artificial heart valves using PIII have produced superior haemocompatility and mechanical durability compared to the LTIC coating


[42, 43]. It is known that titanium nitrides have excellent mechanical and chemical properties, for instance, good wear resistance, inactivity with a number of chemical substances and outstanding hardness [41, 44]. Titanium oxides are known to be fairly compatible with living tissues [38, 42, 45-47] and also inactive to many chemical reactions. In the surface coating industry, incorporation of O and N into Ti alloys is commonly used to improve the mechanical and corrosion properties of the substrates by various methods [48-51].

PLASMA IMMERSION ION IMPLANTATION OF NITI ALLOY PIII has been investigated to modify the surface mechanical and biological properties of nickel titanium alloy. Cheng et al. [52] have implanted tantalum by PIII into NiTi substrates and reported that the treated surfaces possess better corrosion resistance. It has also been found that the oxygen-implanted NiTi alloys possess better corrosion and wear resistance than the untreated NiTi alloys [53, 54] and that oxygen plasma immersion ion implantation can significantly reduce Ni leaching from the NiTi alloys [54]. Carbon and nitrogen PIII can also significantly suppress the Ni concentration on the superficial surface of NiTi alloy [55]. The mechanical strength of nickel titanium alloys is enhanced and TiC and TiN layers which are shown to be bio-compatible [43, 56-59]. Table 1 Implantation and annealing parameters.

Sample

NiTi without

implantation

NiTi with nitrogen

implantation NiTi with carbon

implantation NiTi with oxygen

implantation

Gas type Control

N2 C2H2 O2 RF - 1000W - 1000W

High voltage - -40kV -40kV -40kV

Pulse width - 50µs 30µs 50µs Frequency - 200Hz 200Hz 200Hz

Duration of implantation (min)

- 240 90 240

Base pressure -

7.0×10-6 Torr 1×10-5 Torr 7.0x10-6 Torr

Working pressure -

6.4×10-4 Torr 2.0×10-3 Torr 6.4x10-4 Torr

Dose -

9.6×1016 cm-2 5.5×1016 cm-2 1.0x1017 cm-2

Annealing pressure -

8.0×10-6 Torr 1.0×10-5 Torr 8.0x10-6 Torr

Annealing temperature (ºC) -

450 600 600

Duration of annealing (h) -

5 5 5


In our experiments, the circular NiTi bars with 50.8% Ni (SE508, Nitinol Device Company, Fremont, USA) were cut into disks of 5 mm in diameter and 1 mm in thickness. They were grinded, polished to a shiny surface, and then ultrasonically cleaned with acetone and ethanol before implantation was conducted in our plasma immersion ion implanter [55-57]. The implantation parameters are shown in Table 1. All the treated samples were ultrasonically cleaned again after PIII. To compare the biocompatibility with the conventional use implantable materials, two medical grade metals were also examined. These were stainless steel (5.5mm diameter spinal rod, ISOLA Spinal Instrumentation System, DePuyAcroMed, Raynham, Massachusetts, USA) and titanium (5.5mm diameter spinal rod, Universal Spinal System, Synthes® Inc, USA), which were similarly prepared as 5mm diameter and 1mm discs for cell culturing. The depth profile and surface chemical composition of these two metals were also examined.

CHARACTERIZATION OF PIII SURFACE TREATED NICKEL TITANIUM ALLOY

Surface Chemical Composition and Depth Profile Analysis The elemental depth profile and surface chemical composition of the plasma-treated samples were determined by X-ray photoelectron spectroscopy (XPS) using a Physical electronics PHI 5802 system (Physical Electronics, Minnesota, USA). The energy of the Ar ion beam was 4 keV and the sputtered area was 2 mm × 2 mm. Owing to the coarse surface, it was very difficult to measure the sputtered depth directly. Therefore, the depth scales in the depth profiles were approximated by using a sputtering rate of 22.6nm/min calibrated by a SiO2 source under similar conditions. The elemental depth profiles and surface chemical composition analysis of the untreated NiTi, nitrogen, acetylene and oxygen implanted samples are shown in Figures 2, 3, 4 and 5, respectively. The profiles are plotted on a depth scale based on a sputtering rate calculated from a SiO2 reference under similar conditions. Since it is known that the sputtering rate changes in the surface region and will be different in the samples to that of SiO2, the thicknesses of the implanted zones are approximate, but comparison among different samples is nevertheless more valid. Figure 2 shows the absence of a transitional layer before PIII treatment and Figure 3 shows that a 100 nm thick titanium nitride surface layer is formed. X-ray diffraction (XRD) and high resolution XPS analyses (data not shown) reveal that TiN is the only secondary phase present in the N-implanted layer. With regard to the acetylene-implanted layer, a titanium carbide layer with increasing Ti to C stoichiometric ratios is detected underneath the surface oxide layer. Figure 4 shows a 75 nm thick titanium carbide layer beneath a 25 nm surface oxide. The titanium chemical states are analyzed using high-resolution XPS and oxides of Ti2+, Ti3+, and Ti4+ are found in the


implanted layer. Our results show that a titanium oxide layer about 120nm thick beneath a 25nm surface oxide is formed after the treatment (Figure 5). It should be noted that in all cases, the nickel contents are suppressed to low levels near the surface compared to the untreated NiTi control (Figure 2).

0

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70

80

0 50 100 150 200Depth (nm)

Ato

mic

con

cent

ratio

n (%

)

OxygenTitaniumNickel

Fig. 2. Depth profile of NiTi alloy without surface treatment.

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70

80

0 50 100 150 200Depth (nm)

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cent

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n (%

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NitrogenOxygenTitaniumNickel

Fig. 3. Depth profile of NiTi alloy after nitrogen PIII treatment.


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cent

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n (%

)CarbonOxygenTitaniumNickel

Fig. 4. Depth profile of NiTi alloy after acetylene PIII treatment.

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0 50 100 150 200 250 300Depth (nm)

Ato

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cent

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OxygenTitaniumNickel

Fig. 5. Depth profile of NiTi alloy after oxygen PIII treatment. With respect to the surface chemical composition analysis and depth profiles of the stainless steel and titanium samples, the stainless steel sample contains 61% iron, 11% chromium, 12% nickel and other impurities such as oxygen, carbon and molybdenum in small amounts on the surface as shown in Figure 6. The composition of each element remains stable to a depth of 117 nm. The titanium samples show a higher percentage of carbon and oxygen and small amounts of titanium and


aluminum on the surface (Figure 7). The titanium and aluminum concentrations gradually increase to 70% and 6%, respectively at a depth of 117nm, whereas the oxygen concentration reduces to 25%. Carbon also remains in very low quantity at that depth.

0

10

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30

40

50

60

70

80

0 20 40 60 80 100 120

Depth (nm)

Ato

mic

con

cent

ratio

n (%

)

FeCrNiOCMo

Fig. 6. Depth profile of medical grade stainless steel.

0

10

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60

70

80

0 20 40 60 80 100 120

Depth (nm)

Ato

mic

con

cent

ratio

n (%

)

TiOAlC

Fig. 7. Depth profile of medical grade titanium.


Surface Hardness Analysis The surface hardness of all the samples was measured by nano- indenter (MTS Nano Indenter XP, USA). Nano-indentation tests were conducted on five randomly-selected areas to determine the average hardness and Young’s modulus of the treated and control samples.

0 20 40 60 80 100 120 140 160 1800

20

40

60

80

100

120

140

160

024681012141618202224

Youn

g's

mod

ulus

(GPa

)

Depth (nm)

NiTi Control

Young's modulus Nano-hardness

Nan

o-ha

rdne

ss (G

Pa)

Fig. 8. Hardness and modulus profile of NiTi alloy without surface treatment.

0 20 40 60 80 100 120 140 1600

20

40

60

80

100

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140

160

024681012141618202224

o40kV N-implanted 450 C-annealed

Youn

g's

mod

ulus

(GPa

)

Depth (nm)


Nan

o-ha

rdne

ss (G

Pa)

Fig. 9. Hardness and modulus profile of NiTi alloy after nitrogen PIII treatment. The hardness and the modulus profiles of the control, nitrogen-, acetylene- and oxygen-implanted samples are shown in Figures 8, 9, 10 and 11, respectively and the results are summarized in Table 2. The hardness of the control sample is 4.5 GPa and


the Young’s modulus is 57GPa. All the surface-treated samples possess higher surface hardness and Young’s modulus than the control.

0 20 40 60 80 100 120 140 1600

102030405060708090

100110120130140150160

024681012141618202224

40kV C-implanted 600 oC-annealedYo

ung'

s m

odul

us (G

Pa)

Depth (nm)


Nan

o-ha

rdne

ss (G

Pa)

Fig. 10. Hardness and modulus profile of NiTi alloy after acetylene PIII treatment.

0 20 40 60 80 100 120 140 1600

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120

140

160

024681012141618202224

Youn

g's

mod

ulus

(GPa

)

Depth (nm)

40kV O-implanted 600 oC-annealed


Nan

o-ha

rdne

ss (G

Pa)

Fig. 11. Hardness and modulus profile of NiTi alloy after oxygen PIII treatment.


Table 2 Young’s modulus and hardness of the control and treated samples surfaces.

Sample NiTi NiTi implanted with nitrogen

NiTi implanted with acetylene

NiTi implanted with oxygen

Young's modulus (GPa) 57 150 – 65 110 – 70 150 – 55 Hardness (GPa) 4.5 11 – 5 9.5 – 4.5 9 – 3.5 In the nitrogen-implanted sample, the maximum hardness is 11 GPa at 40 nm from the surface, which gradually decreases to 5 GPa at 165 nm. The Young’s modulus, of about 150 GPa at the topmost surface, decreases to 70 GPa gradually with depth. In the acetylene-implanted sample, the maximum hardness is 9.5 GPa at around 30 nm from the surface and gradually diminishes to 4.5 GPa at 150nm. The Young’s modulus shows a maximum value of 110 GPa at the topmost layer and then decreases gradually to a generally constant value of 70 GPa between 110 and 150 nm from the surface. The lower hardness value in the first 30 nm of this sample is probably due to surface moisture or oxide. In the oxygen-implanted sample, the modulus is found to be 150 GPa near the surface and progressively decreases to 55 GPa between 130 and 160nm. The hardness is 9 GPa at 20 nm and drops to 3.5 GPa at 160 nm. It should be noted that these values are higher than that of 57 GPa for the control sample. Our results suggest that the Young’s modulus of the nitrogen-implanted sample is 163 GPa or 14% higher than that of the substrate, whereas the hardness is 144 GPa or 11% higher throughout the measurement. Hence, the nitrogen-implanted alloy is mechanically stronger than the substrate. With regard to the acetylene-implanted sample, the hardness of the treated layer between 20 and 150 nm is 110 GPa or 11% greater than that of the substrate, whilst the Young’s modulus is 92 – 23% higher throughout the depth of the measurement. In the oxygen-implanted sample, the hardness at 20 – 70 nm is 100 GPa or 11% higher and the modulus at 0 – 120 nm is 163 GPa or 5% higher than that of the untreated substrate. Thus, the mechanical properties of all the treated layers are superior to those of the untreated substrate. Electrochemical Corrosion Analysis Electrochemical tests were conducted to evaluate the corrosion resistance properties of the surface-treated and untreated samples. Tests based on ASTM G5-94 (1999) and G61-86 (1998) were performed using a potentiostat (VersaStat II EG&G, USA) in a standard simulated body fluid (SBF) [78] at a pH of 7.42 and temperature of 37 ± 0.5°C. The solution was prepared by using analytically pure reagents and de-ionized water. The ions concentrations in the SBF are listed in Table 3. The surface area of each sample was 0.181 cm2. A cyclic potential ranging between -400 mV and +1600 mV was applied at a scanning rate of 600 mV per hour. Before the electrochemical tests, the medium was purged with nitrogen for 1 hour to remove dissolved oxygen and nitrogen purging continued throughout the measurements.


Table 3 Ion concentration of SBF in comparison with human blood plasma.

Concentration (mM) Na+ K+ Ca2+ Mg2+ HCO3

- Cl- HPO42- SO4

2- SBF 142.0 5.0 2.5 1.5 4.2 148.5 1.0 0.5 Blood plasma 142.0 5.0 2.5 1.5 27.0 103.0 1.0 0.5

Table 4 Essential results from the electrochemical tests. Sample Control N-treated C-treated O-treated Ecorr (mV) -231 -163 -114 -27

Eb (mV) 272 1120 1170 867

Surface Area (cm2) 0.181 0.181 0.181 0.181

Table 4 lists some of the essential readings from our electrochemical tests in lieu of the more complicated potentiodynamic curves. Ecorr and Eb represent the corrosion potential and the breakdown potential respectively. Higher Ecorr and Eb values represent better corrosion resistance. The Ecorr and Eb values of the control sample are -231 mV and 272 mV, respectively. The Ecorr values measured from the nitrogen-, acetylene- and oxygen-implanted samples are -163 mV, -114 mV, and -27 mV, respectively. The Eb values of the nitrogen-, acetylene- and oxygen-implanted sample are 1120 mV, 1170 mV, and 867 mV, respectively. All of the surface-treated samples exhibit higher Ecorr and Eb values than the untreated sample. These results suggest that the corrosion resistance of the implanted samples is enhanced. Ion Leaching and Surface Morphology Analyses after Corrosion Testing Nickel and/or titanium ions leached from the substrates during the electrochemical tests were determined to assess the effectiveness of the plasma-treated surface to impede out-diffusion from the substrate. The solutions taken from each sample after the corrosion test were analyzed for Ni and Ti concentrations using inductively coupled plasma mass spectroscopy (ICPMS) (Perkin Elmer, PE SCIEX ELAN6100, USA). In order to evaluate the surface morphology of the nitrogen-, acetylene-, and oxygen-implanted samples, the electrochemical tests were performed on another similar set of samples with the aim of comparing the surface morphologies of exposed and unexposed areas in the same sample. The unexposed area was coated with a commercial nail polish that was removed with acetone after the test. The surface morphologies of each sample before and after the electrochemical test were studied using scanning electron microscopy (SEM, JEOL JSM-820, Japan). Table 5 displays the amounts of Ni leached from the surface-treated and untreated samples after the electrochemical tests, as determined by inductively coupled plasma mass spectrometry (ICPMS). The Ni and Ti concentrations in the control sample are


30.2324 and 0.1575ppm respectively. The Ni concentrations in the nitrogen-, acetylene- and oxygen-implanted samples are 0.0117, 0.0082 and 0.0123ppm, respectively. The Ti concentrations are 0.0527 and 0.0057ppm in the nitrogen- and acetylene-implanted samples, respectively. The Ti concentration in the oxygen-implanted sample is undetectable. Our results reveal that the amounts of Ni leached from all the treated samples are significantly reduced. The leached amount is only about 0.03 to 0.04% of that of the control samples.

Exposed area

Un-exposed area Holes

Un-exposed

Exposed area

Holes

Un-exposed

Exposed area

Un-exposed

Exposed area

Holes

A

D

B

C

Fig. 12. Microscopic view of the treated and untreated NiTi samples after electrochemical testing under scanning electron microscopy (SEM) examination. (A) NiTi alloy without surface treatment, (B) with nitrogen PIII implantation, (C) with acetylene PIII implantation, and (D) with oxygen PIII implantation. Table 5 Amounts of Ni and Ti ions detected in SBF by ICPMS after electrochemical tests.

Sample Ni content (ppm) Ti content (ppm) Ni Ti

Control 30.2324 0.1575 N-treated 0.0117 0.0527

C-treated 0.0082 0.057

O-treated 0.0123 Not detectable

Holes


The surface morphologies of the samples after the electrochemical tests are shown in Figure 12. The diameter of the holes on the treated samples surfaces is about 25-30µm, whereas considerably larger holes with more irregular shapes are found on the control sample surface, unequivocally showing that surface treatments can effectively enhance the anti-corrosion capabilities. In Vitro Evaluation of PIII Treated Nickel Titanium Alloy and Conventional Medical Materials To investigate the cyto-compatibility of the plasma-treated and untreated samples, osteoblasts isolated from calvarial bones of 2-day-old mice that ubiquitously expressed an enhanced green fluorescent protein (EGFP) were used, cultured in a Dulbecco’s Modified Eagle Medium (DMEM) (Invitrogen) supplemented with 10% (v/v) fetal bovine serum (Biowest, France), antibiotics (100 U/ml of penicillin and 100 µg/ml of streptomycin), and 2mM L-glutamine at 37°C in an atmosphere of 5% CO2 and 95% air. The specimens (1 mm thick and 5 mm in diameter) were fixed to the bottom of a 24-well tissue culture plate (Falcon) using 1% (w/v) agarose. A cell suspension consisting of 5,000 cells was seeded onto the surface of the untreated NiTi samples, the three types of plasma-implanted samples (oxygen, nitrogen, and acetylene), commercial medical grade stainless and titanium, and wells without any metal discs serving as a control for normal culturing conditions. Cells were grown in 1 ml of medium and changed every three days. Cell attachment was examined after the second day of culture, and cell proliferation examined after 4, 6 and 8 days of culture. Four samples were taken at each time point for statistical analysis. In our study, cells were allowed to reach confluence during the examination period. To determine the cell number, the attached cells were released by digestion with trypsin-EDTA (Invitrogen) and counted using a haematocytometer (Tiefe Depth Profondeur, Marienfeld, Germany). Cell viability was assessed by staining with 0.2% Trypan blue (Sigma). The number of cells was expressed as a mean value ± standard deviation (SD). The data were analyzed using unpaired two-sample t-test and the statistical analysis was performed using the SPSS program (SPSS for Windows, Release 11.0.0). The attached living EGFP-expressing osteoblasts were visualized using a fluorescent microscope (Axioplan 2, Carl Zeiss, Germany) with a 450-490 nm incident filter and the fluorescenct images emitted at 510 nm captured using a Sony DKS-ST5 digital camera. All the plasma-implanted samples are well tolerated by the EGFP-expressing osteoblasts as shown in Figure 13. After culturing for 2 days, the cells started to attach to and proliferate on all the samples except for the stainless steel samples. After 4 days, cell proliferation on the untreated NiTi alloy samples was slightly higher than that of the nitrogen, oxygen and acetylene PIII samples, and about two


times higher than the stainless steel and titanium samples. However, the nitrogen PIII samples exhibited the highest degree of cell proliferation among the samples after 6 and 8 days of culturing. Cell proliferations on the acetylene-implanted, oxygen-implanted and titanium samples were lower than that on the NiTi control sample after 6 and 8 days, but the difference was not significant. Stainless steel showed the least number of viable cells through the period of culturing. A small number of dead cells emerged after 8 days of culturing. The total number of viable cells observed on the untreated, nitrogen, acetylene, and oxygen PIII, stainless steel and titanium samples after 2 and 8 days of culturing are shown in Figures 14 and 15, respectively. These results clearly demonstrate that cells can attach to and proliferate on the surfaces of all the samples tested, indicating that there are no immediate cyto-toxic effects.

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2 4 6 8Day

Num

ber o

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ble

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(X10

000)

NiTi

NiTi-N

NiTi-O

NiTi - C Implanted

S.S.

Ti

Empty well

Fig. 13. Cell proliferation versus number of days. Effectiveness of the Surface Mechanical and Biocompatibility Enhancement of NiTi Treated by Various Surface Technologies A number of surface coatings or surface modification schemes have been studied to enhance corrosion resistance. For instance, Liu et al. [38] have reported that the deposition of a TiO2 film on NiTi alloys using the sol-gel technique improved the anti-corrosion properties, although the surface mechanical properties were not reported. Therefore, the practicality of the use of a coating is questionable if the materials are to be used in orthopedic implants where fretting and abrasive wear are expected. Alternatively, Villermaux et al. [39] have studied the use of excimer laser surface treatment to improve the anti corrosion properties of NiTi alloys. This


Fig. 14. Microscopic view (100x magnification) of NiTi alloy with or without implantations after 2 days of cell culturing showing the EGFP-expressing mouse osteoblasts. Proliferation clusters are obviously seen on the surfaces. (A) NiTi without implantation, (B) N implantation, (C) O implantation, (D) C implantation, (E) stainless steel and (F) pure titanium.

A

100μm

B

100μm

C

100μm

D

100μm

E

100μm

F

100μm


Fig. 15. Microscopic view (100× magnification) of NiTi alloy with or without implantations after 8 days of cell culturing showing the EGFP-expressing mouse osteoblasts. Proliferation clusters are obviously seen on the surfaces. (A) NiTi without implantation, (B) N implantation, (C) O implantation, (D) C implantation, (E) stainless steel and (F) pure titanium.

A

100μm

B

100μm

C

100μm

D

100μm

E

100μm

F

100μm


method is shown to reduce the number of corrosion pits and their sizes and also to increase the thickness of the oxide layer. However, the biological effects of the materials have not been evaluated. Recently, PIII has been used for surface modification of biomaterials. Cheng et al. [80] implanted tantalum by PIII into a NiTi substrate, and reported that the treated surfaces possessed better corrosion resistance. Although tantalum is well compatible with living tissues, the price for this uncommon metal is high, and therefore other alternatives have been explored. It was found that oxygen-implanted NiTi alloys possessed better corrosion and wear resistance than the untreated NiTi alloys. Our results are similar to Tan’s findings [53]. Poon et al. 81 also pointed out that oxygen plasma immersion ion implantation could significantly reduce the Ni leaching of NiTi alloys. The mechanical properties and bio-compatibility of acetylene-implanted samples shown in our study are consistent with Mitura’s study of the mechanical properties and bio-compatibility of titanium alloys after carbon plasma deposition [60]. They reported that a TiC layer underneath a carbon layer was found after treatment. Therefore, the mechanical strength of titanium alloy was enhanced and the TiC layer was shown to be bio-compatible. Nitrogen is the most common impurity used in stainless steel, Ti6Al4V and aluminum alloys to enhance their mechanical properties and corrosion resistance [61, 62]. In our experiments, the NiTi alloy with a TiN layer had lower dissolution currents, higher corrosion resistance, higher wear resistance, and least nickel percentage in wear debris than the other samples. Again, our results are in line with the results obtained by Wan et al. 74 in their study of TiN and Ti-O/TiN films fabricated by PIII&D on Ti6Al4V substrate. Additionally, preliminary results show that the new surface layers formed by PIII do not affect the bulk transformation characteristics of NiTi shape memory alloys such as the shape memory effect and super-elasticity [62]. The results reported here suggest that oxygen, nitrogen or acetylene PIII can effectively suppress the leaching of nickel from the NiTi alloys as shown in Table 5. In our previous immersion tests 81, the plasma-treated and control samples were immersed in SBF for several weeks at 37 oC to simulate in vivo conditions. The amount of Ni leached from the treated NiTi to the SBF was found to be reduced by several orders of magnitude compared to the untreated control. In this study, we accelerated surface corrosion by applying a high voltage and the results are similar indicating that the barrier against Ni out-diffusion introduced by our plasma treatment is very strong and can withstand both simulated in vivo conditions and accelerated degradation processes. This enhancement phenomenon can be attributed to the high affinities of Ti towards N, C and O as compared to Ni under high temperature annealing. This provides a motive force to enrich the surface with elements forming a stronger chemical bond. The heat of formation of lowest titanium oxide is –


913kJmole-1 while that of NiO is –244 kJ-mole-1 [40]. The heat of formation of TiN is –305.6 kJmole-1 [59] while nickel nitrides such as Ni3N are unstable with respect to TiN [60]. The heat of formation of TiC is –773 kJmole-1 [61] while NiC is not well established for the Ni-C phase diagram and does not show stable carbides. The term nickel carbide may only represent interstitial solid solutions of C in Ni which possess the NaCl structure [62]. Therefore, the formation of titanium oxide, nitride, and carbide is energetically favored over the nickel counterparts and this is believed to account for the suppression of Ni in the implanted & annealed region. It should be noted that the degree of suppression depends on the implantation parameters, as reported by Tian et al. [63] in their study of the suppression of nickel in the surface of stainless steel after nitrogen PIII. With regard to the hardness and modulus enhancement, our nano-indentation results show that the treated surfaces possess higher Young’s modulus and hardness than the untreated control surface. Hence, the surface mechanical properties of the treated samples are enhanced. The modified surfaces not only possess better corrosion resistance, but also are capable of resisting mechanical scratch. The efficacy of using PIII to strengthen the materials surface mechanical properties such as hardness and elastic modulus [82] depends on the amorphous matrix composition and the size of precipitates [61]. Additionally, the corrosion resistance seems to be directly proportional to the surface conditions of metals. For instance, smooth surfaces usually give rise to higher corrosion resistance. A crack-free surface is always advantageous because of the reduced chance of localization of corrosive agents. Chemically inert materials such as metal oxides, nitrides, or carbides can effectively reduce the permeability of the corrosive agent [79]. The wetting properties also govern the anti-corrosion capability of a material [79]. Compared to the other treated and untreated NiTi alloys, the in vitro cell culture study indicated that the NiTi alloy after nitrogen implantation exhibited good bio-compatibility. The cell proliferation rate on nitrogen treated surfaces appear to be as good if not better than untreated NiTi alloy at later time points [17, 20, 21, 24, 27]. This finding can be explained by Piscanec’s study [64], which reported the growth of a calcium phosphate phase on TiN-coated titanium implants, but no such activities on untreated titanium implants. Surface composition analysis revealed that this layer consisted of mixed precipitates of TiOxNy oxynitride. This coated layer promoted the deposition of Ca ions due to negative charges localized on the surface after surface treatment. Therefore, this coating was favorable for the formation of bone-like materials under in vivo conditions. It was believed that the TiOxNy oxynitride layer also existed on the nitrogen-implanted NiTi alloy. Medical grade stainless steel and titanium are the most conventional implantable materials in medicine and are believed to possess good corrosion resistance and deformability, especially for medical grade commercial Ti. Numbers of studies


revealed that they are also compatible with living tissues [59, 65-70]. However, Schmidt et al. [71] reported that stainless steel gave poor corrosion resistance under physiological conditions that resulted in nickel (Ni) and chromium (Cr) ions release. The Ni ion is highly toxic to living tissue and reported to be carcinogenic, while the Cr ion may cause impairment of osteoblast proliferation and differentiation, and cytokine release [20, 72, 73]. These findings may explain the insignificant growth of osteoblasts in stainless steel samples in our study, since Cr and Ni elements have been detected on the stainless steel samples. For the titanium samples, our study suggested that the cell viability on this metal was higher than that of stainless steel. However, it still appeared to be less than that on untreated NiTi alloy and all PIII treated NiTi alloys, particularly the untreated NiTi and nitrogen-implanted samples. Kapanen et al. [20] suggested that the apoptosic cell count found in NiTi alloy was less than that in commercial titanium. Although the surface modified and untreated NiTi alloys exhibit good bio-compatibility, a number of conditions such as surface free energy, surface stress, surface morphology, wettability [21, 32, 74, 75], as well as interfacial free energy [18, 76] could also affect the rate of cell attachment and proliferation, and we are therefore conducting further work to study these factors and the change of the surface properties after stressing the materials. It should be noted that we have only demonstrated short-term cell viability. It is possible that the nickel or its compounds may penetrate and leach out through the surface barrier layer over a longer time to the detriment of the cells. The long-term effect of these treated NiTi alloys on cell viability is being tested using both in vitro and in vivo studies.

CONCLUSION The mechanical properties and corrosion resistance in NiTi alloys are improved by conducting C2H2, N, or O plasma immersion ion implantation. Leaching of Ni and Ti ions from the substrate is significantly reduced. Cell culture experiments suggest that all the plasma-treated samples are well tolerated by EGFP-expressing osteoblasts. No immediate cytotoxic effects were found. Long-term in vitro and in vivo bio-compatibility studies are presently being conducted.

REFERENCES [1] Ryhanen J. Biocompatibility Evaluation of NiTi Shape Memory Alloy: University of Oulu;

1999. [2] Chen MF, Yang XJ, Hu RX, Cui ZD, Man HC. Bioactive NiTi shape memory alloy used as

bone bonding implants. Mater Sci Eng C 2004; 24(4): 497-502. [3] Cragg AH, De-Jong SC, Barnhart WH, Landas SK, Smith TP. Nitinol intravascular stent:

results of preclinical evaluation. Radiology 1993; 189(3): 775-8. [4] Dai K, Chu Y. Studies and applications of NiTi shape memory alloys in the medical field in

China. Biomed Mater Eng 1996; 6(4): 233-40.


[5] Es-Souni M, Es-Souni M, Brandies HF. On the transformation behaviour, mechanical properties and biocompatibility of two NiTi-based shape memory alloys: NiTi42 and NiTi42Cu7. Biomaterials 2001; 22(15): 2153-61.

[6] Gil FX, Planell JA. Thermal Cycling and Ageing Effects in Ni-Ti Shape Memory Alloys used in Biomedical Applications. J Biomech 1998; 31:135.

[7] Kong X, Grabitz RG, van Oeveren W, Klee D, van Kooten TG, Freudenthal F, et al. Effect of biologically active coating on biocompatibility of Nitinol devices designed for the closure of intra-atrial communications. Biomaterials 2002; 23(8): 1775-83.

[8] Kujala S, Ryhanen J, Danilov A, Tuukkanen J. Effect of porosity on the osteointegration and bone ingrowth of a weight-bearing nickel-titanium bone graft substitute. Biomaterials 2003; 24(25): 4691-7.

[9] Kujala S, Ryhanen J, Jamsa T, Danilov A, Saaranen J, Pramila A, et al. Bone modeling controlled by a nickel-titanium shape memory alloy intramedullary nail. Biomaterials 2002; 23(12): 2535-43.

[10] Yang P, Zhang Y, Ge M. Internal Fixation with NiTi Shape Memory Alloy Compressive Staples in Orthopedic Surgery. Chinese Med J-Peking 1987; 100(9): 712-4.

[11] Ryhanen J, Kallioinen M, Serlo W, Peramaki P, Junila J, Sandvik P, et al. Bone healing and mineralization, implant corrosion, and trace metals after nickel-titanium shape memory metal intramedullary fixation. J Biomed Mater Res 1999; 47(4): 472-80.

[12] Miyazaki S. Medical and dental application of shape memory alloys. In: K.Otsuka, C.M.Wayman, editors. Shape Memory Materials. 1st ed: Cambridge University Press, 1998. p. 267-281.

[13] Sanders JO, Sanders AE, More R, Ashman RB. A preliminary investigation of shape memory alloys in the surgical correction of scoliosis. Spine 1993; 18(12): 1640-6.

[14] Schmerling MA, Wilkov MA, Sanders AE, Woosley JE. Using the shape recovery of nitinol in the Harrington rod treatment of scoliosis. J Biomed Mater Res 1976; 10(6): 879-892.

[15] Trepanier C, TK L, Tabrizia M, Yahia L. In Vivo Biocompatibility study of NiTi Stents. Second International Conference on Shape Memory and Superelastic Technology 1997; California, USA; 1997. p. 423-428.

[16] Filip P, Musialek J, Michalek K, Yen M, Mazanec K. TiAlV/Al2O3/TiNi shape memory alloy smart composite biomaterials for orthopedic surgery. Mater Sci Eng A 1999; 273-275: 769-74.

[17] Bogdanski D, Koller M, Muller D, Muhr G, Bram M, Buchkremer HP, et al. Easy assessment of the biocompatibility of Ni-Ti alloys by in vitro cell culture experiments on a functionally graded Ni-NiTi-Ti material. Biomaterials 2002; 23(23): 4549-55.

[18] El Medawar L, Rocher P, Hornez J-C, Traisnel M, Breme J, Hildebrand HF. Electrochemical and cytocompatibility assessment of NiTiNOL memory shape alloy for orthodontic use. Biomol Eng 2002; 19(2-6): 153-60.

[19] Putters J, Kaulesar S, De Z, Bijma A, Besselink P. Comparative cell culture effects of shape memory metal (Nitinol), nickel and titanium: a biocompatibility estimation. Eur Surg Res 1992 1992; 24: 378-82.

[20] Kapanen A, Ilvesaro J, Danilov A, Ryhanen J, Lehenkari P, Tuukkanen J. Behaviour of Nitinol in osteoblast-like ROS-17 cell cultures. Biomaterials 2002; 23(3): 645-50.

[21] Kapanen A, Kinnunen A, Ryhanen J, Tuukkanen J. TGF-[beta]1 secretion of ROS-17/2.8 cultures on NiTi implant material. Biomaterials 2002; 23(16): 3341-6.

[22] Kapanen A, Ryhanen J, Danilov A, Tuukkanen J. Effect of nickel-titanium shape memory metal alloy on bone formation. Biomaterials 2001; 22(18): 2475-80.

[23] Kujala S, Pajala A, Kallioinen M, Pramila A, Tuukkanen J, Ryhanen J. Biocompatibility and strength properties of nitinol shape memory alloy suture in rabbit tendon. Biomaterials 2004; 25(2): 353-8.

[24] Rocher P, El Medawar L, Hornez J-C, Traisnel M, Breme J, Hildebrand HF. Biocorrosion and cytocompatibility assessment of NiTi shape memory alloys. Scripta Mater 2004; 50(2): 255-60.


[25] Ryhanen J, Kallioinen M, Tuukkanen J, Junila J, Niemela E, Sandvik P, et al. In vivo biocompatibility evaluation of nickel-titanium shape memory metal alloy: muscle and perineural tissue responses and encapsule membrane thickness. J Biomed Mater Res 1998; 41(3): 481-8.

[26] Ryhanen J, Niemi E, Serlo W, Niemela E, Sandvik P, Pernu H, et al. Biocompatibility of nickel-titanium shape memory metal and its corrosion behavior in human cell cultures. J Biomed Mater Res 1997; 35(4): 451-7.

[27] Wever DJ, Veldhuizen AG, Sanders MM, Schakenraad JM, van Horn JR. Cytotoxic, allergic and genotoxic activity of a nickel-titanium alloy. Biomaterials 1997; 18(16): 1115-20.

[28] Berger-Gorbet M, Broxup B, Rivard C, Yahia LH. Biocompatibility testing of NiTi screws using immunohistochemistry on sections containing metallic implants. J Biomed Mater Res 1996; 32(2): 243-8.

[29] Castleman L, Motzkin S. The biocompatibility of Nitinol. In: William D, editor. Biocompatibility of clinical implant materials. Florida: CRC press, Inc, 1981. p. 129-154.

[30] Jia W, Beatty MW, Reinhardt RA, Petro TM, Cohen DM, Maze CR, et al. Nickel release from orthodontic arch wires and cellular immune response to various nickel concentrations. Journal of Biomed Mater Res1999; 48(4): 488-95.

[31] Doran A, Law F, Allen M, Rushton N. Neoplastic transformation of cells by soluble but not particulate forms of metals used in orthopaedic implants. Catheter Cardio Inte, 1998; 19: 751-9.

[32] Es-Souni M, Es-Souni M, Fischer-Brandies H. On the properties of two binary NiTi shape memory alloys. Effects of surface finish on the corrosion behaviour and in vitro biocompatibility. Biomaterials 2002; 23(14): 2887-94.

[33] Shih C-C, Lin S-J, Chen Y-L, Su Y-Y, Lai S-T, Wu GJ, et al. The cytotoxicity of corrosion products of nitinol stent wire on cultured smooth muscle cells. J Biomed Mater Res 2000; 52(2): 395-403.

[34] Dalmau L, Alberty H, Parra J. A study of nickel allergy. J Prosthet Dent 1984; 52: 116-9. [35] Lamster I, Kalfus D, Steigerwald P, Chasens A. Rapid loss of alveolar bone associated with

nonprecious alloy crowns in two patients with nickel hypersensitivity. J Periodontol 1987; 58: 486-92.

[36] Espana A, Alonso ML, Soria C, Guimaraens D, Ledo A. Chronic urticaria after implantation of 2 nickel-containing dental prostheses in a nickel-allergic patient. Contact Dermat 1989; 21: 204-6.

[37] Sanford WE, Niboer E. Renal toxicity of nickel in humans. In: Niboer E, Nriagu JO, editors. Nickel and Human Health Current Perspectives. Canada: John Wiley & Sons, Inc., 1992. p. 123-134.

[38] Liu J-X, Yang D-Z, Shi F, Cai Y-J. Sol-gel deposited TiO2 film on NiTi surgical alloy for biocompatibility improvement. Thin Solid Films 2003; 429(1-2): 225-30.

[39] Villermaux F, Tabrizian M, Yahia LH, Meunier M, Piron DL. Excimer laser treatment of NiTi shape memory alloy biomaterials. Appl Surf Sci 1997; 109-110: 62-6.

[40] Conrad JR, Radtke JL, Dodd RA, Worzala FJ, Tran NC. Plasma source ion-implantation technique for surface modification of materials. J Appl Phys 1987; 62(11): 4591-6.

[41] Chu PK, Chen JY, Wang LP, Huang N. Plasma-surface modification of biomaterials. Mater Sci Eng R 2002; 36(5-6): 143-206.

[42] Leng YX, Huang N, Yang P, Chen JY, Sun H, Wang J, et al. Structure and properties of biomedical TiO2 films synthesized by dual plasma deposition. Surf Coat Tech 2002; 156(1-3): 295-300.

[43] Yang P, Huang N, Leng YX, Chen JY, Fu RKY, Kwok SCH, et al. Activation of platelets adhered on amorphous hydrogenated carbon (a-C:H) films synthesized by plasma immersion ion implantation-deposition (PIII-D). Biomaterials 2003; 24(17): 2821-29.


[44] Fu RKY, Tian X, Chu PK. Enhancement of implantation energy using a conducting grid in plasma immersion ion implantation of dielectric/polymeric materials. Rev Sci Instrum 2003; 74(8): 3697-700.

[45] Kuphasuk C, Oshida Y, Andres CJ, Hovijitra ST, Barco MT, Brown DT. Electrochemical corrosion of titanium and titanium-based alloys. J Prosthet Dent 2001; 85(2): 195-202.

[46] O'Brien B, Carroll WM, Kelly MJ. Passivation of nitinol wire for vascular implants--a demonstration of the benefits. Biomaterials 2002; 23(8): 1739-48.

[47] Tan L, Crone WC. Surface characterization of NiTi modified by plasma source ion implantation. Acta Mater 2002; 50(18): 4449-60.

[48] Chu PK, Fu RKY, Zeng X, Kwok DTK. Metallic contamination in hydrogen plasma immersion ion implantation of silicon. J Appl Phys 2001; 90(8): 3743-9.

[49] Ng BS, Annergren I, Soutar AM, Khor KA, Jarfors AEW. Characterisation of a duplex TiO2/CaP coating on Ti6Al4V for hard tissue replacement. Biomaterials 2005; 26(10): 1087-95.

[50] Burstein GT, Liu C, Souto RM. The effect of temperature on the nucleation of corrosion pits on titanium in Ringer's physiological solution. Biomaterials 2005; 26(3): 245-56.

[51] Cairney JM, Tsukano R, Hoffman MJ, Yang M. Degradation of TiN coatings under cyclic loading. Acta Mater 2004; 52(11): 3229-37.

[52] Cheng Y, Wei C, Gan KY, Zhao LC. Surface modification of TiNi alloy through tantalum immersion ion implantation. Surf Coat Tech 2004; 176(2): 261-5.

[53] Tan L, Dodd RA, Crone WC. Corrosion and wear-corrosion behavior of NiTi modified by plasma source ion implantation. Biomaterials 2003; 24(22): 3931-9.

[54] Poon RWY, Ho JPY, Liu X, Chung CY, Chu PK, Yeung KWK, et al. Anti-corrosion performance of oxidized and oxygen plasma-implanted NiTi alloys. Mater Sci Eng A 2004; In Press, Corrected Proof.

[55] Poon RWY, Yeung KWK, Liu XY, Chu PK, Chung CY, Lu WW, et al. Carbon plasma immersion ion implantation of nickel-titanium shape memory alloys. Biomaterials 2004;In Press, Corrected Proof.

[56] Tang L, Tsai C, Gerberich WW, Kruckeberg L, Kania DR. Biocompatibility of chemical-vapour-deposited diamond. Biomaterials 1995; 16(6): 483-8.

[57] Narayan J, Fan WD, Narayan RJ, Tiwari P, Stadelmaier HH. Diamond, diamond-like and titanium nitride biocompatible coatings for human body parts. Mater Sci Eng B 1994; 25(1): 5-10.

[58] Kwok SCH, Yang P, Wang J, Liu X, Chu PK. Hemocompatibility of nitrogen-doped, hydrogen-free diamond-like carbon prepared by nitrogen plasma immersion ion implantation-deposition. J Biomed Mater Res A 2004; 70(1): 107-14.

[59] Czarnowska E, Wierzchon T, Maranda-Niedbala A. Properties of the surface layers on titanium alloy and their biocompatibility in in vitro tests. J Mater Process Tech 1999; 92-93: 190-4.

[60] Mitura E, Niedzielska A, Niedzielski P, Klimek L, Rylski A, Mitura S, et al. The properties of carbon layers deposited onto titanium substrates. Diam Relat Mater 1996; 5(9): 998-1001.

[61] Girardeau T, Bouslykhane K, Mimault J, Villain JP, Chartier P. Wear improvement and local structure in nickel-titanium coatings produced by reactive ion sputtering. Thin Solid Films 1996 1996/9/1;283(1-2):67-74.

[62] Grant DM, Green SM, Wood JV. The surface performance of shot peened and ion implanted NiTi shape memory alloy. Acta Metall Mater1995; 43(3): 1045-51.

[63] Tian X, Fu RKY, Wang L, Chu PK. Oxygen-induced nickel segregation in nitrogen plasma implanted AISI 304 stainless steel. Mater Sci Eng A 2001; 316(1-2): 200-4.

[64] Piscanec S, Colombi Ciacchi L, Vesselli E, Comelli G, Sbaizero O, Meriani S, et al. Bioactivity of TiN-coated titanium implants. Acta Mater 2004; 52(5): 1237-45.

[65] Shih C-C, Shih C-M, Su Y-Y, Su LHJ, Chang M-S, Lin S-J. Effect of surface oxide properties on corrosion resistance of 316L stainless steel for biomedical applications. Corros Sci 2004; 46(2): 427-41.


[66] Niinomi M. Recent research and development in titanium alloys for biomedical applications and healthcare goods. Sci Technol Adv Mater 2003; 4(5): 445-54.

[67] Sgambato A, Cittadini A, Ardito R, Dardeli A, Facchini A, Dalla Pria P, et al. Osteoblast behavior on nanostructured titanium alloys. Mater Sci Eng C 2003; 23(3): 419-23.

[68] Tsyganov I, Wieser E, Matz W, Reuther H, Richter E. Modification of the Ti-6Al-4V alloy by ion implantation of calcium and/or phosphorus. Surf Coat Tech 2002; 158-159: 318-23.

[69] Lagneau C, Farges J-C, Exbrayat P, Lissac M. Cytokeratin expression in human oral gingival epithelial cells: in vitro regulation by titanium-based implant materials. Biomaterials 1998; 19(11-12): 1109-15.

[70] Guleryuz H, Cimenoglu H. Effect of thermal oxidation on corrosion and corrosion-wear behaviour of a Ti-6Al-4V alloy. Biomaterials 2004; 25(16): 3325-33.

[71] Schmidt C, Ignatius AA, Claes LE. Proliferation and differentiation parameters of human osteoblasts on titanium and steel surfaces. J Biomed Mater Res 2000; 54(2): 209-15.

[72] Swiontkowski M, Agel J, Schwappach J, McNair P, Welch M. Cutaneous metal sensitivity in patients with orthopaedic injuries. J Orthop Trauma 2001; 2: 86-9.

[73] Wang J, Wicklund B, Gustilo R, Tsukayama D. Prosthetic metals interfere with the functions of human osteoblast cells in vitro. Clin Orthop 1997; 339: 216-26.

[74] Bonewald L, Kester M, Schwartz Z, Swain L, Khare A, Johnson T, et al. Effects of combining transforming growth factor beta and 1,25- dihydroxyvitamin D3 on differentiation of a human osteosarcoma (MG-63). J Biol Chem 1992; 267(13): 8943-9.

[75] Ponsonnet L, Comte V, Othmane A, Lagneau C, Charbonnier M, Lissac M, et al. Effect of surface topography and chemistry on adhesion, orientation and growth of fibroblasts on nickel-titanium substrates. Mater Sci Eng C 2002; 21(1-2): 157-65.

[76] Ponsonnet L, Reybier K, Jaffrezic N, Comte V, Lagneau C, Lissac M, et al. Relationship between surface properties (roughness, wettability) of titanium and titanium alloys and cell behaviour. Mater Sci Eng C 2003; 23(4): 551-60.

chapter 2 plasma surface modification of nickel titanium ...abstract: nickel-titanium (niti) shape...

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