Elements-Added Diamond-Like Carbon Film for Biomedical Applications

Elements-added diamond-like carbon films for biomedical applications were investigated. .e aim of this work was to study the effects of the elemental contents (silicon and silicon-nitrogen) in a DLC film on its properties for biomedical applications. Pure DLC, Si-DLC, and Si-N-DLC films were prepared fromC2H2, C2H2 : TMS, and C2H2 : TMS :N2 gaseousmixtures, deposited on an AISI 316L substrate using the plasma-based ion implantation (PBII) technique. .e structure of films was analyzed using Raman spectroscopy. .e chemical composition of films was measured using energy dispersive X-ray spectroscopy (EDS). .e average surface roughness of films was measured by using a surface roughness tester. .e hardness and elastic modulus of films were measured by using a nanoindentation hardness tester. .e friction coefficient of films was determined using a ball-on-disk tribometer. .e surface contact angle was measured by a contact angle measurement. .e corrosion performance of each specimen was measured using potentiodynamic polarization. .e biocompatibility property of films was conducted using the MTT assay cytotoxicity test. .e results indicate that the Si-N-DLC film shows the best hardness and friction coefficient (34.05GPa and 0.13, respectively) with a nitrogen content of 0.5 at.%N, while the Si-DLC film with silicon content of 14.2 at.%Si reports the best contact angle and corrosion potential (92.47 and 0.398V, respectively). .e Si-N-DLC film shows the highest cell viability percentage of 81.96%, which is lower than the uncoated AISI 316L; this is a considerable improvement. All specimens do not demonstrate any cytotoxicity with approximate viabilities between 74% and 107%, indicating good biocompatibilities.


Introduction
Implantable biomedical devices are one of the most popular methods of medical field to treat human illnesses.ere are many kinds of biomedical devices, such as vascular stents, artificial joints, artificial knees, and bone plates [1].Most biomedical devices are made from stainless steel and titanium alloys because of the favorable mechanical and biocompatibility properties of these materials [2][3][4][5].
However, problems arising from material deterioration are often detected after long-term use.Corrosion is unavoidable due to various ions in the body reacting electrochemically with the surface of these metallic materials [6].A metal device may release metal into the body, causing allergic reactions as thrombogenicity of the blood in complicated and aggressive physiological environments.us, ideal biomedical devices provide biocompatibility that prevents metallic ion release [6].Surface coatings can improve both the mechanical properties and biocompatibility of biomedical devices that are in direct contact with blood and tissue.
Coating film technology has been studied for biomedical applications and includes diamond-like carbon (DLC) film.DLC films have many excellent properties including high hardness, low friction coefficient, good corrosion resistance, and good biocompatibility properties [7][8][9][10].Additionally, DLC films can be doped with certain elements, such as hydrogen (H-DLC), fluorine (F-DLC), and sulfur (S-DLC), to improve performance.Previous studies have shown that DLC films doped with silicon (Si-DLC) are significantly improved corrosion properties due to the formation of a passive film on their surfaces [11][12][13][14][15][16].Moreover, silicondoped DLC films deposited by the sputtering method with the Si concentration varied from 4 to 16 at.%can reduce platelet adhesion on the material surface by modifying the hydrophobicity of the material [17].Additionally, the corrosion resistance of a pure DLC lm can be increased via silicon and nitrogen doping by increasing the number of sp 3  sites in the lm [18].
Plasma-based ion implantation (PBII) has been developed to improve DLC lm properties and fabricate three-dimensional (3D) materials with complex shapes.In this technique, low working temperature avoids lm quality degradation, such as loose and rough surface structure, and avoids DLC graphitization caused by normal CVD technique which is performed at higher working temperature [19,20].
Currently, there is no report on the deposition of Si-and Si-N-added DLC lms on the AISI 316L substrates by the PBII method aimed at comparing the mechanical, tribological, and corrosion performance, especially to compare the cell viability percentage.
In this paper, plasma-based ion implantation (PBII) was utilized to prepare di erent elements including silicon-and silicon-nitrogen-added DLC lms (henceforth denoted as Si-DLC and Si-N-DLC lms).e aim of the study was to compare and study the e ects of element contents on the deposition and properties of the lms for biomedical applications.

Materials and Methods
A schematic of the PBII apparatus used in this study is shown in Figure 1 [18].e inner dimensions of the vacuum chamber are 600 × 630 × 200 mm 3 , and the chamber produced a residual pressure of approximately 1 × 10 −4 Pa.Plasma was generated from a radio frequency (RF, 13.6 MHz) glow discharge and a negative high-voltage pulse power supply connected to the sample holder.
Medical-grade AISI-type 316L stainless steel was used for the substrate material.Prior to plasma coating, specimens were polished down to 2000 grits using standard abrasive paper and the mirror was polished with 1 μm diamond paste and then cleaned in an ultrasonic cleaning bath.Substrates were sputter-cleaned with Ar + for 20 min to remove surface contaminants and surface oxides using a −10 kV pulse bias voltage.e DLC lm interlayer was then deposited with CH 4 for 60 min to improve adhesion between the lm and the substrate using a bias voltage of −20 kV. e RF power was set at 300 W, and the pressures of the sputter-cleaning and interlayer deposition were both 1 Pa.A pulse width of 5 μs, a pulse delay of 25 μs, and a pulse frequency of 1 kHz were used during the sputter-cleaning and interlayer deposition processes.Pure DLC, Si-DLC, and Si-N-DLC lms were deposited using the PBII technique.e pure DLC lm was used with a C 2 H 2 precursor gas.e Si-DLC lm was prepared from gaseous mixtures of C 2 H 2 and TMS at three di erent ow rate ratios (1 : 2, 1 : 4, and 1 : 6).e Si-N-DLC lm was deposited from gaseous mixtures of C 2 H 2 : TMS : N 2 at three di erent ow rate ratios (14 : 1 : 2, 14 : 1 : 4, and 14 : 1 : 6).e bias voltage and deposition pressure of all of the lms were set at −5 kV and between 2 and 2.5 Pa, respectively.e RF power was set at 300 W, and the pulse frequency was set to 1 kHz at a pulse width of 5 μs and a pulse delay of 25 μs.e thickness of each lm was approximately 500 nm.e deposition conditions for the DLC, Si-DLC, and Si-N-DLC lms are shown in Table 1.
e structures of the lms were analyzed using Raman spectroscopy (JASCO NRS-1000 DT; beam diameter 4 μm and wavelength 532 nm).
e chemical composition of lms was measured using energy dispersive X-ray spectroscopy (EDS).e average surface roughness (R a ) at the top surface of the lms was measured using a surface roughness tester (MAHR; MARSURF PS 10; length of travel 4.8 mm, and cuto length 0.8 mm).e hardness and elastic modulus of lms were measured using a nanoindentation hardness tester using a diamond ball indenter (Berkovich-type) with an indentation load of 1,000 μN.
e friction coe cients of the lms were determined using a ball-on-disk tribometer under ambient air conditions.In the friction test, a dry sliding test was performed using a ball indenter made of AISI 440C (SUS 440C, diameter of 6.0 mm) under a normal applied load of 1 N, a rotational radius of 4 mm, a linear speed of 31.4 mm/s, and 3,000 frictional rotations.e surface contact angle was measured by using a contact angle measurement device from Kyowa Interface Science Co., Ltd. using 1 μL of distilled water.
e corrosion performance of each specimen was measured using potentiodynamic polarization in a simulated physiological Ringer's solution (ASTM F2129-17) at a pH of 7.5.A Pt sheet and Ag/AgCl were used as the counter and the reference electrodes, respectively.e potential voltage was varied from −1 V up to +1 V at a scanning rate of 1 mV/s.e cytotoxicity was selected for testing from the most abundant element added to the AISI 316L substrate; these were the pure DLC, Si-DLC 1 : 6, and Si-N-DLC 14 : 1 : 6 based on the ISO 10993-5 guidelines using the Dulbecco's modi ed Eagle medium (DMEM) dilution method.All samples were placed in sterile Duran bottles and then sterilized at 121 °C for 15 min.L929 broblast cells (mouse broblasts) at 1 × 10 5 cells/mL were grown in a 96-well plate at a volume of 100 μL/well and then kept in an incubator (37 °C, 5% CO 2 ) for 24 hours.e cell

Results and Discussion
3.1.Structure of Films by Raman Spectroscopy.Raman spectroscopy was used to investigate the DLC lm's structure due to its nondestructive mechanism.e Raman spectra measured are shown in Figures 2-4 and indicate that the Raman spectra of pure DLC, Si-DLC, and Si-N-DLC lms deposited at di erent ow rate ratios onto the fabricated AISI 316L substrates.
e Raman spectra in the wavelength of 1,000 cm −1 -1,800 cm −1 were deconvoluted into Gaussian D and G peaks. e Raman spectrum was measured by curve tting procedure using two Gaussian distribution peak shapes as show in Figure 2.
e peak position and I D /I G ratio of the integrated areas by the D and G band in the DLC lms fabricated with various gaseous mixtures and gaseous ratios are presented in Table 2.It has been obtained that the position of G band is related to the bond-angle disorder or sp 3 bonding content, while the I D /I G ratio is related to disorder [16,21].e lms in this experiment show a broad spectrum composed of a D band (1,350 cm −1 ) and a G band (1,580 cm −1 ), which are similar to the peaks produced by conventional DLC lms.
e results shown in Figures 3 and 4 and Table 2 show that the G peak and I D /I G ratio of pure DLC decrease due to silicon and silicon-nitrogen incorporation.For the Si-DLC lm, the G peak tends to decrease from 1467 cm −1 (C : Si/1 : 2) to 1456 cm −1 (C : Si/1 : 6), while the I D /I G ratio decreases from 0.15 to 0.11 as the silicon content increases.is result is likely due to the altered microstructure as more sp 3 structures form as the silicon content increases.Additionally, longer destrained bonds vibrate at a lower frequency, leading to the G band shifting to a lower frequency.is outcome can be partially attributed to a reduction in compressive stress when silicon is introduced into the lms [12].
For the Si-N-DLC lm, the G peak tends to increase slightly from 1521 cm −1 (C : Si : N 14 : 1: 2) to 1524 cm −1 (C : Si : N 14 : 1 : 6), while the I D /I G ratio increases from 0.55 to 0.58 due to silicon-nitrogen incorporation.
ese results were likely caused by unstable sp 3 cluster formation and the increasing number or size of the lm's graphitic structure due to nitrogen incorporation.ese results thus suggest that the formation of a higher sp 2 and lower sp 3 clusters in the lm is caused by an increase in the silicon-nitrogen content [22].Advances in Materials Science and Engineering 3

Relative Atomic Content of Films.
e relative atomic contents measured on the top surface of the films deposited on the stainless-steel substrate are shown in Table 3. e carbon (C), silicon (Si), nitrogen (N), oxygen (O), and iron (Fe) concentrations were measured using the EDS, and values are shown in units of atomic percent (at.%).Concentrations were normalized to a total of 100 at.% and ignored the contribution of hydrogen because the hydrogen content could not be measured by EDS. e results also show that Fe appeared in all films, likely due to the thin film thickness (approximately 500 nm).It is assumed that the penetration depth of the electrons during the test is higher than the film thickness [18].
As shown in Table 3, the DLC film exhibits the highest C content (91.1 at.%), while the Si-DLC film shows that the Si content increased from 13.3 to 14.2 at.% by changing the gas flow rate from 1 : 2 to 1 : 6.
e concentration of the Si-N-DLC film shows that the N content increased from 0.2 to 1.0 at.% when the gas flow rate changed from 14 : 1:2 to 14 : 1:6.In addition, a part of the EDS broad spectrum of Si-N-DLC film at the ratio of 14 : 1:6 is shown in Figure 5.

Surface Hardness and Elastic Modulus of Films.
e hardness and elastic modulus of a film are important for biomedical application because contact pressures in some applications, such as artificial knees and joint replacements, may cause excessive material wear.e wear resistance of a material is directly related to its hardness and elastic modulus.
Each film's hardness and elastic modulus were determined on its top surface using a nanoindentation hardness test and a diamond ball indenter (Berkovich-type) with an indentation load of 1,000 μN. e hardness and elastic modulus of pure DLC, Si-DLC, and Si-N-DLC are shown in Figures 6 and 7 and Table 4.
As shown in Table 4, the hardness and elastic modulus of the films increased from those of pure DLC (25.46 GPa and 184.51 GPa, respectively) as the silicon content increased.For Si-DLC, the hardness and elastic modulus increased from 29.16 GPa (13.3 at.%Si) to 31.15GPa (13.7 at.%Si) and 209.76 GPa to 216.92 GPa when the TMS flow rate ratio increased.e increased hardness and elastic modulus due to the change in microstructure was observed via the Raman analysis; the formation of sp 3 site structures increased the hardness and elastic modulus as the silicon content increased [18,23].However, the hardness and elastic modulus of the 1 : 6 Si-DLC film (14.2 at.%Si) decreased to 26.05 and 185.85 GPa, respectively.C 2 H 2 and TMS used during fabrication: when the TMS flow rate increased, the hydrogen content in the film increased.e polymeric sp 3 C−H bonds established due to the high hydrogen concentration decreased the film's hardness [23].Si-DLC films are likely to be saturated with hydrogen, especially when using a TMS precursor.
e same trend verifies the reliability of our results.
e Si-N-DLC film exhibited the high hardness and elastic modulus values 34.05 GPa and 220.97 GPa (0.5 at.%N), respectively, despite its number and size of sp 3 clusters being lower than those of the Si-DLC film.e formation of C−N bonds yielded high hardness and elastic modulus values because C−Si bonds are weaker than C−N bonds.However, when the nitrogen was increased to 1.0 at.%N, the hardness and elastic modulus decreased to 26.28 GPa and 185.61GPa, respectively, which was likely due to the lower bond strength of the N−Si network [24,25].

Friction Coefficient of Films.
e friction coefficients of pure DLC, Si-DLC, and Si-N-DLC films were determined under ambient air conditions and are shown in Figure 8.
e effects of the silicon and silicon-nitrogen contents on the friction coefficients of the Si-DLC and Si-N-DLC films were investigated.
e results indicate that all coating films can increase the low friction coefficient of AISI 316L.
e friction coefficient of pure DLC was 0.17.It was found that increasing the silicon content in the film increased the film's friction coefficient.
e Si-DLC film's friction coefficient increased from 0.22 (13.3 at.%Si) to 0.30 (14.2 at.%Si) when the TMS flow rate ratio increased, which reduced the carbon content in the film.In general, carbon acts as a solid lubricant that reduces friction between ballfilm contact areas [26,27].Additionally, for the Si-DLC film, the friction coefficient also increased due to the film's lower surface hardness and elastic modulus.e contact area at the ball-film interface actually increased as the film hardness and elastic modulus decreased, resulting in the film surface having difficulty supporting the high load induced by the friction coefficient [28].e Si-N-DLC film exhibited the lowest friction coefficient (0.13 at 0.5 at.%N) due to the film's high hardness and elastic modulus.e film's low friction coefficient can be attributed to the lubricating effect of the sp 2 cluster in the film coatings, as shown in Table 2 [29].

Contact Angle of Films.
e contact angles measured on the top surfaces of the films are shown in Table 5.A volume of 1 μL of distilled water was released onto the top surface of the films under atmospheric conditions at room temperature.
Many studies have shown that surface roughness and surface hydrophilicity strongly increase the blood compatibility of implants: a smooth and hydrophobic surface is beneficial to the blood compatibility [21,30,31].

Advances in Materials Science and Engineering
Contact-angle studies play an important role in understanding the surface phenomena of the substrates.
e polished smooth surface of AISI 316L reported a contact angle value of 93.47 °, describing the hydrophobic nature of AISI 316L [32,33].e contact angle value reduced to 79.33 °for the pure DLC film.For the Si-DLC film, the contact angle increased from 88.93 °(13.3 at.%Si) to 92.47 °(14.2 at.%Si) as silicon content increased.is was likely due to the decreasing surface energy of the films as the silicon content increased and more strongly influenced the   6 Advances in Materials Science and Engineering hydrophobic property of the film [34].Moreover, the contact angle was relatively unaffected by surface roughness, as shown in Table 3.
For the Si-N-DLC film, the contact angle was found to be related to the nitrogen and carbon atomic concentration ratio.e contact angle decreased as the nitrogen content increased, as shown in Table 5. e contact angle of the Si-N-DLC film shifted slightly as the gas flow rate changed from 14 : 1 : 2 to 14 : 1 : 6. e Si-N-DLC surface was thus more hydrophilic than the pure DLC and Si-DLC films, which might be due to an increase in the dispersive and polarity components of surface energy from the high number of nucleophilic electrons produced by the formation of N-H and C�N networks [35].

Corrosion Resistance Property of Films.
Biocompatibility is simply defined as the performance of a material to be permitted by the body.A foreign body is created when materials are implanted in a body and is directly related to the corrosion behavior of the material due to the tendency for the alloy to release potential toxic ions from corrosion behavior [36].
In general, samples with greater corrosion potential and lower current density provide better corrosion performance.An improvement in the corrosion resistance of a film is demonstrated by a shift in the polarization curve towards a region of lower current density and greater potential [37].Additionally, variation in the film microstructure allows different corrosion behaviors to occur [38].Generally, a more prevalent sp 3 structure tends to increase electrochemical corrosion resistance [18,38].
Figures 9 and 10 show the potentiodynamic curves for uncoated AISI 316 L, pure DLC, Si-DLC, and Si-N-DLC films.Table 5 summarizes the electrochemical parameters measured from the potentiodynamic polarization curves shown in Figures 9 and 10.All films show higher corrosion potential (E corr ) and lower current density (i corr ) than the uncoated AISI 316 L sample.Pure DLC shows a corrosion potential and current density of −0.010 V and 0.949 × 10 −6 A, respectively.Corrosion potential increased, and current density decreased due to silicon (Si-DLC) and silicon-nitrogen (Si-N-DLC) incorporation.For the Si-DLC film, corrosion potential (E corr ) increased from 0.198 V to 0.398 V and current density (i corr ) decreased from 0.887 × 10 −6 A/cm 2 to 0.280 × 10 −6 A/cm 2 due to silicon incorporation.For the Si-N-DLC film, corrosion potential (E corr ) increased from −0.001 V to 0.198 V and current density (i corr ) decreased from 0.931 × 10 −6 A/cm 2 to 0.570 × 10 −6 A/cm 2 due to silicon-nitrogen incorporation.
us, the best corrosion resistance was observed in the Si-DLC film with a silicon content of 14.2 at.%Si with corrosion potential (E corr ) and current density (i corr ) values of 0.398 V and 0.280 × 10 −6 A/cm 2 , respectively.e greater corrosion resistance of the Si-DLC and Si-N-DLC films than that of the uncoated AISI 316L and pure DLC film shows that the incorporation of silicon and siliconnitrogen into the pure DLC film improves corrosion resistance.
is outcome is likely due to the increased number of sp 3 sites in the film's structure [39] and the creation of a silicon oxide passive film on the film's surface [11].Additionally, the sp 3 structure created via silicon incorporation can reduce the internal stress that causes the corrosion [12].Additionally, the results show more corrosion-resistant potential than the previous report mentioned above [14].e report showed the corrosion potential and current density of Si-DLC up to 1.5 × 10 −5 mA/cm 2 and −0.7 V with high silicon content of 30 at.% [14].

Cytotoxicity Test of Films.
A cytotoxic test was designed according to ISO 10993-5 standard, and L 929 fibroblast cells were used in this method.e cytotoxic test is typically used to evaluate materials for biomedical applications to confirm their implantability.e cell viability percentage of uncoated AISI 316L, pure DLC, Si-DLC, and Si-N-DLC are shown in Figure 11.
Figures 11 and 12 show the viability and morphology of the L929 fibroblast cells after being exposed to the MTT solution.
e Si-N-DLC film exhibited a cell viability of 81.96%, which is lower than the uncoated AISI 316 L (107%) but higher than the pure DLC (78.16%) and the Si-DLC films (74.4%).However, none of the specimens were found to be cytotoxic.Cell viability ranged from 74.4% to 107%, indicating good biocompatibility because viability percentages (live/dead %) were not lower than 50% as per ISO10993-5.Additionally, no visible change in cell morphology was observed; the cells still exhibited a typical fibroblastic morphology, as shown in Figure 12 [40].ese results agree with those reported by Sui et al. [21], Gotzmann et al. [41], Bociaga et al. [42], and Antunes et al. [43].

Conclusions
Elements-added diamond-like carbon films were investigated for biomedical applications.Pure DLC, Si-DLC, and Si-N-DLC films were prepared on an AISI 316L  e Si-N-DLC film showed the best hardness and friction coefficient (34.05 GPa and 0.13, respectively) with a nitrogen content of 0.5 at.%N, while the Si-DLC film with silicon content of 14.2 at.%Si reported the best contact angle and corrosion potential (92.47 • and 0.398 V, respectively).
e Si-N-DLC film showed the highest cell viability percentage of 81.96%, which was lower than the uncoated AISI 316L; this is a considerable improvement.e DLC and Si-DLC films showed cell viability percentages of 78.16% and 74.4%, respectively.All specimens did not demonstrate any cytotoxicity with approximate viabilities between 74% and 107%, indicating good biocompatibilities.

Table 1 :
Deposition of pure DLC, Si-DLC, and Si-N-DLC lms.Film type Gaseous mixture Bias voltage (kV) RF power (W) Deposition pressure (Pa) Gas ow rate ratio Actual gas ow (sccm)

Table 2 :
G peak and ID/IG ratio of pure DLC, Si-DLC, and Si-N-DLC films.

Table 3 :
Relative atomic content and average roughness of pure DLC, Si-DLC, and Si-N-DLC films.

Table 4 :
Mechanical properties and friction coefficient of pure DLC, Si-DLC, and Si-N-DLC films.

Table 5 :
Surface contact angle, corrosion potential (E corr ), and corrosion current density (i corr ) of pure DLC, Si-DLC, and Si-N-DLC films.