Experimental investigations into corrosion behaviour of DMLS manufactured Ti6Al4V alloy in different biofluids for orthopedic implants

doi:10.21203/rs.3.rs-4828859/v1

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Experimental investigations into corrosion behaviour of DMLS manufactured Ti6Al4V alloy in different biofluids for orthopedic implants

https://doi.org/10.21203/rs.3.rs-4828859/v1

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Ti6Al4V alloy is widely used in the biomedical implants industry due to the excellent combination of biocompatibility, mechanical strength & corrosion properties. However, long-time use of Ti6Al4V implants may result in degradation at surge due to corrosive conditions of the human body. In this study, a three-electrode electrochemical cell was used to investigate the corrosion behavior of Ti6Al4V alloy parts manufactured using direct metal laser sintering (DMLS) and conventional cast technique. The corrosion behaviour of Ti6Al4V alloy was used to investigate in three different biofluids (corrosion media) physiological saline solution (PSS), simulated body fluids (SBF), and phosphate-buffered saline (PBS)using the potentiodynamic polarization (PDP), and electrochemical impedance spectroscopy (EIS) test. The relevant results showed the DMLS-produced Ti6Al4V alloy has ~ 22% higher corrosion resistance than the conventionally cast Ti6Al4V alloy in all three biofluids. Additionally, the DMLS-produced Ti6Al4V alloy in SBF has the lowest corrosion rate of 3.44 x10^− 4 mm per year and exhibits the highest polarization resistance (Rp) of 3364 Ω.cm² due to the formation of a stable passive film. The corrosion resistance of the DMLS and cast produced Ti6Al4V alloy in all three biofluids followed the order of PSS < PBS < SBF.

Corrosion

DMLS

Biomedical

Biofluids

Wettability

ANOVA

The corrosion behavior is an important characteristic to determine the service life for any metallic implants. The corrosion behavior is influenced by many factors, such as material, manufacturing process, microstructure, biofluids etc.[1–3]. Any of these parameters can affect the performance of the implants, which may be harmful to the human body. Titanium (Ti) and its alloy, particularly Ti6Al4V, possesses desirable properties for implant manufacturing like suitable mechanical properties, high corrosion resistance, and excellent biocompatibility. All these properties are the primary requirement for the implant of different domains like orthopedic, hip, oral etc. However, after long-term use, the Ti6Al4V-made implants may fail prematurely due to fretting corrosion [4]. Researchers have suggested that Ti6Al4V orthopedic implants which are exposed to the corrosive medium may release toxic elements such as aluminum (Al) and vanadium (V) ions that may results into the illnesses such as Neuropathy, Osteomalacia, and Alzheimer. This continuous degrading of the material may also result in premature failure of implants [5].

The salinity of the body fluid also acts as a potent corrosive agent for Ti6Al4V-based implants. It hinders the formation of a stable oxide layer of TiO₂ on its surface, further resulting in an unstable corrosion resistance. Henceforth, it is necessary to improve the corrosion resistance of Ti6Al4V alloy to reduce the damage and increase the service life of the implant. Many authors have reported the electrochemical corrosion behaviour of titanium and its alloy in the presence of various biocompatible fluids such as artificial saliva, NaCl, Ringer solution azide, etc. [6, 7]. Razavi et al.[8] evaluated the corrosion behavior of Ti6Al4V alloy in an HCl solution and reported that, corrosion resistance can be also be enhanced by laser nitriding. They concluded that the formation of TiN during laser nitriding over the surface results in stable passivation over the surface. Ibris et al. [9] investigated the corrosion behaviour of various titanium alloys such as Ti7Al4.5V, and Ti5Al2.5Fe alloy in simulated body fluid (SBF) and noticed that although alloying elements do not influence the charge transfer process, however, they can improve the thickness of the protective layer. Licausi et al. [10] compared the corrosion behavior of TiAl4V alloy fabricated by powder metallurgy and casted route in an artificial saliva solution. They concluded that the formation of a passivation layer over the surface plays a crucial role in improving the corrosion resistance of orthopedic implants. They have also suggested that this layer can be developed during manufacturing and may break very easily from the surface in NaCl and phycological solutions [11]. As mentioned earlier, the corrosion resistance of any alloy is highly affected by its alloying element, manufacturing process, and microstructure [12]. For any specific alloy such as Ti6Al4V, SS316L, etc., varying its alloying composition is not possible; therefore, by changing the manufacturing process, the corrosion resistance can be enhanced.

Looking at the merits and demerits of the different fabrication techniques, additive manufacturing (AM) appears to be more suitable for bioimplants as it offers better customization and short cycle time for this customization[13]. Due to the high applicability of AM technique, several researchers have performed biomedical investigations on AM processed parts [14, 15]. Afzali et al. [16] reported that the Laser Engineered Net Shaped (LENS) produced Ti6Al4V alloy had a lower corrosion rate as compared to the cast and powder metallurgy parts due to the presence of martensitic microstructure. Even the corrosion behavior of the EBM-produced Ti6Al4V alloy was found to be different than the other AM processes. This may be due to the different grain morphologies obtained from different AM routes. The researchers have also found that by changing the microstructure with suitable heat treatment, the corrosion properties of AM produced Ti6Al4V can be altered.

From the literature review it can be concluded that (i) No study has been reported on the corrosion behaviour of DMLS-produced Ti6Al4V alloy in various biofluids such as physiological saline solution, phosphate buffered saline and simulated body fluid, particularly for biomedical applications. Therefore, to fill the research gap the aim of the present study is to: (i) Evaluate the electrochemical corrosion behavior of DMLS-produced Ti6AL4V alloy in the physiological saline solution (PSS), simulated body fluids (SBF) and phosphate-buffered saline (PBS) to determine its performance for orthopedic implants applicability. (ii) Electrochemical corrosion behaviour of DMLS-produced Ti6Al4V alloy was also compared with the conventionally cast Ti6Al4V alloy from which commercial implants are made. The corrosion behavior of cast and DMLS-produced Ti6Al4V alloy was measured with the three-electrode electrochemical cell using responses such as open circuit potential (OCP), potentiodynamic polarization (PDP) and electrochemical impedance spectroscopy (EIS). The characterization of the surface of Ti6Al4V alloy before and after the electrochemical study was also performed using the scanning electron microscope (SEM) and X-ray diffraction (XRD). Additionally, the analysis of variance (ANOVA) was also performed to evaluate the efficacy of the manufacturing process and corrosive media on the corrosion rate of Ti6Al4V alloy. This will help us to understand the corrosion behaviour of DMLS-produced Ti6Al4V alloy in different biofluids.

In this study, corrosion tests were performed on cast and DMLS produced Ti6Al4V alloy samples. The DMLS samples were produced using Ti6Al4V powder with the particle size of 30–45 µm. The conventionally cast samples of the same alloy procured from the commercial supplier (Megha Steel, Bombay, India) were also tested as a reference for comparing the corrosion behavior with the DMLS-produced parts. The chemical composition (wt%) of both conventionally cast and DMLS produced Ti6Al4V samples was measured using energy dispersive spectroscopy (EDS) are shown in Table 1.

Table 1

Chemical composition of (wt%) of Titanium alloy produced by conventionally cast and DMLS produced sample (based on the supplier data sheet)
	Composition (wt. %)								Harness (VHN)	Porosity (%)
Sample	Al	V	C	Fe	O	N	H	Ti	Harness (VHN)	Porosity (%)
Cast	5.3	3.6	0.10	0.25	0.13	0.06	0.012	Balance	$\:{336}_{+12}^{-12}$	~ 0.15
DMLS	5.9	4.2	0.08	0.30	0.20	0.05	0.015	Balance	$\:{420}_{+20}^{-20}$	~ 1.43

2.1. Sample preparation

All the DMLS produced Ti6Al4V alloy samples were fabricated using EOS INT M280 machine with the following parameters: laser power 340 W, scanning speed 1250 mm/s, hatching distance 0.2 mm, layer thickness 60 µm. Cubical-shaped samples with 10 mm side length were prepared for both cast and DMLS. In order to avoid any oxidation, all the samples were fabricated in a controlled argon gas atmosphere of 20 PSI pressure. Further, all the samples were annealed at 800°C for 2h under the presence of argon, followed by furnace cooling to remove the internal stresses. These process parameters were opted based on the available literature for achieving maximum density in fabricated Ti6Al4V samples [17–19]. Further all the samples were polished using the silicon carbide (SiC) paper up to grit size of 2500 followed by the suspended alumina paste to obtain the roughness value of 1 ± 0.5 µm for microstructure and corrosion tests. For all the corrosion tests the cylindrical samples with an exposed area of 1cm² were used. Additionally, microhardness and density were also measured for both samples using standard procedures.

2.2. Microhardness and density measurement

The microhardness was measured using the microhardness testes (Walter UHL Asslar, Germany Model: VMHT VMH 002) with 2N load and 25 s dwell time. At least 5 indents were taken across the top surface of the specimen, and the average value was reported. The porosity was also measured using the Archimedes’ principle. The pore size was measured using the scanning electron microscope (SEM) images in the ImageJ software. The results of both density and hardness tests were mentioned in Table 1.

2.3. Microstructure and surface morphology characterization

For microstructural characterization, the polished samples (cast and DMLS produced) were etched by the kroll etchant using 2% of hydrofluoric acid (HF) and viewed under scanning electron microscope (SEM) (Oxford. Germany). Apart from microstructure the phases present in both the samples were also studied with the help of X-ray diffractometer (XRD) equipment (Bruker, USA Model: D2-Phaser ). All the XRD analysis were performed at room temperature with the following settings: wavelength of (1.54 Å) (CuKα radiation), 2θ angle (20^o – 90^o), scanning speed (1–2 sec). Further, X’Pert High Score Plus analysis software was used for the plotting of the XRD data. The peaks corresponding to different phases and their crystallographic planes in an XRD plot were characterized employing the JCPDS database.

2.4. Surface wettability measurement

The contact angle was measured (before and after the corrosion test) using the contact angle meter (Holmarc Opta Mechatronics Ltd, Kochi, India Model: HO-ED-M-01) to evaluate the wettability behaviour. The wettability test was performed using a 20 µl of all three biofluids. All tests were performed five times at various regions of polished and corroded samples to minimize the randomness error in results.

2.5. Electrochemical corrosion testing

The corrosion behaviour of the Ti6Al4V alloy samples were performed on the electrochemical workstation (Princeton Applied Research Oak Ridge, US Model: K0235,). The electrochemical corrosion test was performed according to the ASTM G59 standards. The corrosion setup consists of a three-electrode electrochemical cell, biofluid, potentiostat, and monitor as shown in Fig. 1.

The Ti6Al4V sample was made as the working electrode (W.E.), platinum wire as a counter electrode (CE) and saturated calomel electrode (SCE) as a reference electrode (RE). The electrochemical workstation was used to perform the OCP, PDP and EIS.

2.5.1 Open circuit potential (OCP)

The OCP represents the steady-state potential developed at the surface of the W.E. (i.e. Ti6Al4V alloy sample). OCP data was recorded over the period of 1800s with scan rate of 0.5V/s, for stabilizing the curve in different biofluids.

2.5.2 Potentiodynamic polarization (PDP)

The Potentiodynamic polarization (PDP) test was carried out to determine the corrosion rate and passivation behavior of both the cast and DMLS-produced samples in different biofluids. During PDP, the potential of the W.E. was varied between − 0.5V to + 1.2 V with respect to the open circuit potential (OCP) at a scanning rate of 0.5 millivolt per second (mV/s). The corrosion responses, including corrosion potential (E_corr) (V) and corrosion current density (i_corr) (µA/cm²) were further used to estimate the corrosion rate (CR) (mm/y) of the corroded samples using the following equation [4]

CR = K$\:\times\:\frac{{EW\times\:i}_{corr}}{d\times\:A}$(i)

where K is the corrosion constant (0.00327 mm/y) [4], EW is the electrochemical equivalent weight for Ti6Al4V was measured 11.06 gm/equivalent using the standard methods [20], i_corr is the corrosion current density (µA/cm²), and d is the density of cast and DMLS produced Ti6Al4V alloy are 4.39 and 4.43 g/cm³ respectively.

The anodic (βa) and cathodic (βc) were obtained from the polarization plot. These values were further used to determine the polarization resistance (Rp) (Ratio of △E/△I) by using the following equation[21]:

$\:\frac{\varDelta\:E}{\varDelta\:I}\left(Rp\right)=\frac{1}{2.303}\times\:\frac{{\beta\:}_{a}\times\:{\beta\:}_{b}}{{\beta\:}_{a}+{\beta\:}_{c}}$ (ii)

2.4.1 Electrochemical impedance spectroscopy (EIS)

The EIS observation were recorded between the frequency range of 0.01Hz to 1000 KHz with a scanning rate of 10 mV/s. The recorded EIS data of each sample was analyzed using the Zsimpwin software. It is further used to fit the recorded EIS data using the electrochemical equivalent circuit model. As mentioned earlier that the test was performed in three different biofluids, such as Psychological Saline Solution (PSS), Simulated Body Fluid (SBF), and Phosphate-Buffered Saline (PBS). All biofluids/solutions were supplied by the Himeida Pvt Ltd. Mumbai, India. The chemical composition of all three biofluids are tabulated in Table 2.

Table 2

Chemical composition of the biofluids
Biofluids	PSS (g/l) [22, 23]	PBS (g/l)[24]	SBF(g/l) [25]
NaCl	12	8.5	8.036
KCl	--	0.22	0.45
CaCl₂	--	1.1	0.293
KH₂PO₄3H₂O	--	0.136	0.230
Na₂HPO₄. H₂O	0.58	0.141	0.248
Na₂CO₃	--	--	0.035
MgCl₂. 6H₂O	0.296	--	0.311
NaHCO₃	1.714	--	0.352
Na₂SO₄	--	--	0.072
Glucose	1.00	1.00	1.00
pH level	4.0	7.4	7.4

All electro-chemical corrosion test was performed in identical conditions of temperature (37°C) and relative humidity 10%. The temperature of 37°C was chosen to mimic the human body environment. All the electrochemical tests were conducted minimum of three times for the evaluation of the repeatability of the results.

2.6. Surface analysis of the corroded samples

The morphology of sample’s surface was obtained using SEM and elemental analysis was carried out using EDS. Additionally the XRD was also performed on corroded samples to investigate the phases formed during corrosion. In order to determine the phase on the corroded surface, XRD was performed between 20° to 90° with a step size of 0.039°.

2.7. ANOVA

The analysis of variance (ANOVA) was performed to assess the effect of the manufacturing process and biofluids on the corrosion rate using the MINITAB 15 software. The outcomes of the ANOVA were used to analyze the efficacy of both factors on the corrosion rate. Table 3 shows the level of both the factors used for ANOVA.

Table 3

Levels for chosen control factors
Control factor	Level I	Level II	Level III
Manufacturing process	Casting	DMLS	-
Biofluids	PSS	PBS	SBF

This section includes the findings of all the corrosion experiments along with microstructural and phase analysis. The effect of two parameters (biofluids and manufacturing process) on the corrosion rate using ANOVA analysis, and wettability results were also discussed.

3.1 Analysis of phases and microstructure

The XRD pattern of the conventionally cast and DMLS-produced Ti6Al4V alloy indicates the presence of peaks corresponding to the α and ß phases (Fig. 2). The position of peak and the intensities are nearly identical, suggesting a similarity in the hexagonal closed packed (HCP) (α and ß phase). The DMLS Ti6Al4V samples exhibit relatively intense peaks corresponding to the α’, while the weak peak shows ß phase. The results suggest that the amount of ß phase was higher in the conventionally cast Ti6Al4V alloy as compared to the DMLS-produced Ti6Al4V alloy [14].

Figure 3 depicts the SEM images of the cast (Fig. 3a) showing the coarse grain microstructure that is composed of the $\:\alpha\:$ and $\:\beta\:$ phase with a longitudinal pockets lamellar structure [26]. The average grain size was measured ~ 26 µm. The grain boundary was not clearly visible and the primary equiaxial $\:\alpha\:\:$phase with secondary fine acicular fine $\:\alpha\:$ phase can also be observed. The microstructure of DMLS produced Ti6Al4V alloy (Fig. 3c), showed the needle-shaped structure composed of α’+ß phase. The $\:\alpha\:{\prime\:}$ grains were formed due to the high cooling rate and made the non-equilibrium martensitic needle-shaped structure known as widmanstatten microstructure. The non-equilibrium microstructure (α’+ß phase) is metastable and is sensitive to corrosion.

Some pores (~ 3–4 µm) are also present in the microstructure of DMLS-produced Ti6Al4V alloy (Fig. 3b). The average grain size of the α’ phase in DMLS produced samples was measured ~ 10 µm. The microstructure of the DMLS-produced Ti6Al4V consists of finer grain as compared to the cast Ti6Al4V alloy that, results in higher yield strength. Figure 3b and 3d shows the EDS analysis of the cast and DMLS-produced Ti6Al4V alloy. The α’ phase is higher in the DMLS-produced Ti6Al4V alloy than the cast Ti6Al4V alloy. It is due to the higher concentration of aluminum, which enhances the stability of α’ phase [27, 28].

3.2 Open circuit potential (OCP)

Figure 4 shows the variation of open circuit potential (OCP) with respect to dipping time for the cast and DMLS fabricated Ti6Al4V alloy in all three different biofluids namely, PSS, PBS, and SBF.

These curves can be categorized into two zone (i) in the first zone the OCP value rapidly increases due to dissolution of metal ions from the surface due to the attack of Cl⁻ ions. (ii) in the second zone, the value of OCP gradually increased towards the positive potential, indicating the formation of a protective oxide layer on the W.E. (Ti6Al4V alloy) surface. It is due to the reaction of released metal ions with the oxygen present in the biofluids, which further results into the deposition of oxides on the sample surface. The layer thickness of this oxide increases with time, which stabilizes the OCP value [14] after 1800 sec, indicating formation of uniform passive layer of oxide on the Ti6Al4V alloy surface [29]. This layer of oxide acts as a barrier for further electrochemical reaction, thus enhancing its corrosion resistance [30]. A fluctuating or negative OCP value indicates a higher corrosion rate, while a constant and positive OCP value suggests a lower corrosion rate. The more positive value of OCP denotes the presence of oxide film on the Ti6Al4V alloy surface, which improves the corrosion resistance [31]. The mean values of the OCP in steady conditions for all the cases are presented in Table 4.

Table 4

OCP value of cast and DMLS produced Ti6Al4V alloy in different biofluids.
Manufacturing process/ biofluids	PSS	PBS	SBF
Cast Ti6Al4V (mV)	-0.12	-0.07	~ 0.01
DMLS – Ti6Al4V (mV)	-0.09	-0.05	~ 0.03

The value of OCP for DMLS-produced Ti6Al4V seems to be slightly higher than the cast Ti6Al4V alloy in all the biofluids/electrolyte media. The OCP value for both samples tested in the SBF solution was recorded as lower than that of the PBS and PSS. It has been reported that the pH level of biofluids highly influence the corrosion behaviour of alloy, and in specifically, higher chloride (Cl) concentrations result into more active potential [16]. The lowest value of OCP was observed in the PSS solution due to the presence of the higher chloride concentration as compared to the PBS and SBF solutions. Moreover, the higher concentration of Na₂HPO₄ and KH₂PO₄ in SBF react with the Ti alloy surface and act as a barrier due to the formation of passive film [16][32]. From these results it can be concluded that biofluids and the manufacturing process greatly affect the corrosion behaviour of the alloy. The value of OCP in the different biofluids have shown the following order: SBF < PBS < PSS.

3.3 Potentiodynamic polarization (PDP)

The Potentiodynamic polarization (PDP) plot shows a similar corrosion mechanism observed from OCP test results. Both types of samples show passivation during anodic polarization in all the biofluids.

Firstly, all three biofluids show (Fig. 5), that the passive layer on the working electrode (Ti6Al4V alloy) surface begins to dissolve due to the initial increment in potential. This can be considered as the initiation of the corrosion on the Ti6Al4V alloy surfaces. This phenomenon was followed by the passivation region, where the potential increased and the current density remained constant. A wider passive region in the polarization plot is advantageous and shows better corrosion resistance, excellent material stability and superior performance in all three biofluids [33–35]. During the passivation process, a stable layer of oxide film was developed on the W.E., which inhibits further dissolution in the aqueous solution. It can be seen (Fig. 5) in the polarization plot that DMLS-produced T6Al4V alloys have a higher passivation range. Moreover, a wider passive region was also found in SBF followed by the PBS and lowest in the PSS biofluids.

Table 5

Electrochemical corrosion parameters obtained from the potentiodynamic polarization plot for the cast and DMLS produced Ti6Al4V alloy in different biofluids
Manufacturing process	Cast Ti6Al4V			DMLS Ti6Al4V
Biofluids	PSS	PBS	SBF	PSS	PBS	SBF
i_corr(µA/cm²)	0.0988 ± 0.084	0.0707 ± 0.067	0.05771 ± 0.055	0.0758 ± 0.046	0.0528 ± 0.039	0.0417 ± 0.023
E_corr (V)	-0.4010 ± 0.03	-0.3820 ± 0.05	-0.3713 ± 0.07	-0.3025 ± 0.07	-0.2827 ± 0.06	-0.1917 ± 0.09
Corrosion rate (1x10^− 4mm/y)	8.05 ± 0.5	5.77 ± 0.85	4.70 ± 0.20	6.23 ± 0.35	4.35 ± 0.20	3.44 ± 0.15
Rp (Polarization resistance) (ΩA.cm²)	1083 ± 35	1592 ± 30	1796 ± 25	2193 ± 35	3084 ± 30	3364 ± 25
Ip (pitting current) (µA/Cm²)	0.5610 ± 0.092	0.5690 ± 0.079	0.5636 ±0.055	0.5654 ± 0.042	0.5853 ± 0.096	0.5690 ± 0.014
E_pit (V)	0.1754 ± 0.058	0.2269 ± 0.056	0.2885 ± 0.074	0.4118 ± 0.049	0.4533 ± 0.028	0.6552 ± 0.037
Passivation region	-0.5764 ± 0.009	-0.6089 ± 0.058	-0.6599 ± 0.036	-0.7143 ± 0.004	-0.7361 ± 0.003	-0.8470 ± 0.002

Table 5 shows the values of corrosion potential (E_corr), corrosion current density (i_corr) (µA/cm2), and polarization resistance (R_p). These values were obtained from the slope of anodic and cathodic regions of the polarization plot, which is further used to calculate the corrosion rate (mm/y) using the Eq. (1) [31] for cast and DMLS produced Ti6Al4V alloy in all biofluids. A lower value of i_corr and higher E_corr indicate the higher stability of oxide layer formation which further inhibits corrosion and gives a lower corrosion rate [34].

It can be concluded from Table 5 that the DMLS produced samples have the highest value of E_corr and minimum value of i_corr (for all three biofluids) than the cast produced Ti6Al4V alloy. The DMLS-produced Ti6Al4V alloy consists of α’ martensitic microstructure and a small amount of ß phase, which exhibits lower corrosion rate [36, 37].

As compared to the corrosion behaviour of Ti6Al4V alloy in biofluids (Table 5), the SBF has the minimum corrosion rate followed by the PBS, whereas PSS has the highest value of corrosion rate. It is mainly because calcium (Ca) and phosphate ions are present in the SBF, which interact with the Ti ions and form a complex TiO₂ on the surface. This further results into the higher corrosion resistance of Ti6Al4V alloy [36].

It is noticed that the value of Icorr and Ecorr in SBF are shown to be higher than in the PBS and PSS. However, the R_p shows the opposite tendency for both cast and DMLS-produced Ti6Al4V alloy. A higher value of R_p denotes the high corrosion resistance in the specific medium.

The pitting potential (E_p) and pitting current density (i_p) was computed using the potentiodynamic polarization plot (Fig. 5). The E_p and i_p represent the degradation rate on the Ti6Al4V alloy surface. The higher value of E_p and lower value of i_p shows better stability of the passive film [34]. The value of E_p also dependent on the type of biofluids. The SBF had the highest E_p value followed by the PBS and PSS. This result suggests that calcium and Phosphate ions are less conducive to pitting as compared to the chloride ions, despite the high Ecorr in PSS. The critical damage in the PSS is due to the presence of Cl- ions that may induce cracking. Overall, the corrosion parameters represent that the DMLS-produced Ti6Al4V alloy has better corrosion resistance as compared to the cast produced Ti6Al4V alloy in all three biofluids.

3.4 Electrochemical Impedance spectroscopy (EIS)

Figure 6 and 7 shows the EIS plot (Nyquist and Bode plot) of the cast and DMLS-produced Ti6Al4V alloy in all three biofluids. Nyquist plot (Fig. 6), show the relation between the real (Z’) and imaginary (Z”) impedance for both the samples (cast and DMLS produced Ti6Al4V alloy) in PSS (Fig. 6a), PBS (Fig. 6b) and SBF (Fig. 6c). An equivalent circuit or Modified Randle model (Fig. 6d) was employed to fit the EIS data for cast and DMLS produced Ti6Al4V alloy in different biofluids for analyzing the corrosion behaviour.

The equivalent circuit was constructed using biofluid’s resistance ‘R_s’ (Ω) and oxides film resistance ‘R_f’ (Ω). Additionally, the oxide layer film capacitance ‘C_f’(farad) was used along with double layer capacitance “C_dl’(farad) accommodate the effect of passive film (formed due to interaction of polished Ti6Al4V surface with atmosphere). Due to presence of non-uniform oxide layer on surface, the behaviour of double layer capacitance is non ideal and this was incorporated through constant phase elements ‘Q’ (µF/Cm²).

The double layer formed in this case is highly porous in nature and hence a constant phase element (Q) was used to define the capacitance. Moreover, the charge transfer resistance (Rct) in the double layer is negligibly low, due to which Rct was not used in the fitted equivalent circuit. This could also be confirmed from the very small diameter of the double-layer curve present in the Nyquist plot at high frequency. This is in line with the previous work by Valereto et al. [38]. Alternately, a large diameter curve in the low frequency region confirms the formation of highly stable and defect free passive layer, resulting in fitting the circuit using a pure capacitance and film resistance. Hence, the equivalent circuit (Fig. 6d) was used to fit the Nyquist plot exhibiting lower chi-square value (10^− 3). Warburg impedance ‘W’(Ω) refers to the diffusion at the interface of the oxide layer and biofluid. This modification allows more precise representation of the EIS data and corrosion behaviour of the W.E. (Ti6Al4V alloy) in the biofluids. Furthermore, the n is the dispersion constant which represents the denser concentration of passivation film. The various parameters such as R_s, R_f, Q, C, W, and dispersion constant ‘n’ obtained from the equivalent circuit by fitting the EIS data for Cast and DMLS-produced Ti6Al4V Alloy in all three biofluids are shown in Table 6.

Table 6

Vale of impedance parameters obtained from EIS for the cast and DMLS produced Ti6Al4V alloy in different biofluids.
Samples	Cast Ti6Al4V			DMLS Ti6Al4V
Biofluids	PSS	PBS	SBF	PSS	PBS	SBF
R_s (Ω)	2.181 ± 0.032	2.862 ± 0.041	3.994 ± 0.059	4.582 ± 0.072	6.927 ± 0.098	8.043 ± 0.076
R_f (Ω)	42.713 ± 6	44.791 ± 7	51.092 ± 8	45.059 ± 5	49.182 ± 4	56.199 ± 6
Q (µF/Cm²) (1x10^− 6)	8.04 ± 0.09	9.53 ± 0.12	13.21 ± 0.18	8.88 ± 0.08	9.49 ± 0.11	15.24 ± 0.15
C_f (Farad)(1x10^− 6)	.0111 ± 0.002	.0115 ± 0.002	.0122 ± 0.001	.0114 ± 0.001	.0118 ± 0.002	0.127 ± 0.001
C_d (Farad) (1x10^− 6)	18.712 ± 0.150	6.789 ± 0.200	7.436 ± 0.200	16.087 ± 0.300	6.283 ± 0.200	5.354 ± 0.100
W (Ω)(1x10^− 4)	3.966 ± 0.20	4.288 ± 0.35	5.224 ± 0.15	4.175 ± 0.15	4.916 ± 0.10	6.385 ± 0.20
n	0.436 ± 0.04	0.462 ± 0.02	0.493 ± 0.04	0.464 ± 0.05	0.485 ± 0.03	0.508 ± 0.04

Nyquist plot for cast and DMLS produced Ti6Al4V alloy in PSS (Fig. 6a), PBS (Fig. 6b) and SBF (Fig. 6c) reveal different semicircular arc, indicating the resistance of surface and interference between the biofluid and Ti6Al4V alloy surface [39]. The diameter of the semi-circle arc (as shown inset Fig. 6) represents the oxide layer resistance ‘R_f’. The higher value of R_f indicates the thickness of the compact layer which further contributes to corrosion resistance [40]. Figure 6, shows that DMLS produced Ti6Al4V alloy has the higher corrosion resistance as compared to the cast Ti6Al4V alloy, as its Nyquist plot have shown bigger semi-circle diameter. Moreover, it can also be seen that SBF has the largest semi-circle diameter, showing higher film resistance, followed by the PBS whereas PSS has the smallest film resistance. As mentioned earlier, this variation in the corrosion resistance is mainly because of the different thicknesses of protective oxide layer formed on the Ti6Al4V alloy surface. It is also verified by the R_f value that is shown in Table 6 follows the same order of i_corr (SBF > PBS > PSS) obtained from PDP results.

The formation of the oxide layer over the passive film results into two capacitances as observed from the circuit fitting data. Additionally, double layer capacitance is also formed at the interface of the oxide film and biofluid. However, due to non-uniform surface of oxide layer its behaviour is not ideal which is accommodated by incorporation of Constant phase elements. The ions diffusion can also be observed from the Nyquist plot, which was incorporated by inclusion of Warberg element. From the graph and the obtained constants (obtained after fitting EIS data) it can be observed that the ion diffusion is different for all the cases. From the EIS plot it can also be observed that the proportion of the semicircle and line is different for DMLS and cast produced Ti6A4V alloy. This is mainly due to the formation of the thickness of the oxide layer which results in higher resistance and Warberg impedance. The same trend can also be observed for different biofluids as the portion of same circle as an impedance to inclined line is larger for SBF followed by PBS and PSS, respectively. This can also be observed from the obtained Warburg impedance value for different bio-fluids. Additionally, the SEM image (Fig. 9) also confirms this phenomenon as thinner oxide layer was formed on PSS which provides lesser impedance and higher diffusion.

The value of n (Table 6) represents the dispersion constant, and the higher value of n means the compact passivation film. From the literature [41] it was found that increasing the dispersion constant n may result high corrosion resistance. DMLS produced Ti6Al4V in SBF shows the maximum dispersion constant value followed by PBS and PSS.

From all these results it can be concluded that the value for DMLS-produced Ti6Al4V alloy has the highest corrosion resistance in SBF compared to the PBS and PSS biofluids, which can also be supported by the EIS results. The dissolution of titanium and further formation of TiO₂ provide stability to the oxide layer on the Ti6Al4V alloy sample surface. The dissolution and formation of TiO₂ continue to enhance the stability of the protective layer on the Ti6Al4V alloy sample surface. On the contrary, in PSS biofluids, the dissolution of TiO₂ is aided and regulated by the presence of Cl⁻ complexes, leading to its faster degradation.

The EIS Bode plot 1 (frequency vs impedance) (Fig. 7a-c), depicts the variation of absolute value of impedance and Bode plot 2 (frequency vs phase angle) (Fig. 7d-f) represents the variation of phase angle. It may be noted from Fig. 7c that DMLS produced Ti6Al4V alloy has the higher value of impedance in the low-frequency zone, indicating higher resistance of its W.E. and cast produced Ti6Al4V alloy has the minimum impedance value, suggesting the lower resistance value of its W.E. Now considering the corrosion behaviour of biofluids according to the biofluids, the SBF shows the highest value of impedance at the low frequency, indicating the higher Rs which further leads to higher corrosion resistance. However, PSS possesses the lowest value of impedance at higher frequency, indicating the lowest resistance due to Cl and P ions which further induces the dissolution of the oxide layer. The observations are verified by the value of R_s and R_f shown in Table 6.

From Fig. 7 (d, e, f) it can be observed that, cast produced Ti6Al4V alloy immersed in PSS has the lowest value of phase angle between 0.01 and 1 Hz, resulting in the poor corrosion resistance due to the faster charger transfer and lower impedance (Fig. 7d). SBF has the highest phase angle value (Fig. 7f) over a wide frequency of range among all three biofluids for both cast and DMLS produced Ti6Al4V alloy followed by the PBS. It has a peak value almost equal to 80°, which indicates a more compact film of oxides over surface than the other two biofluids. It can be concluded that DMLS produced Ti6Al4V alloy in SBF has the highest phase angle in SBF that will provide excellent surface protection against corrosion.

3.5 Corrosion mechanism and morphology of corroded surface

Figure 8 shows the schematic of corrosion mechanism of cast and DMLS produced Ti6Al4V alloy in biofluids. Figure 8a represents formation of very thin passive layer on the cast and DMLS produced Ti6Al4V alloy surface due their contact of the environment (after polishing). This layer acts as a barrier and protects from the surrounding environment and prevents further corrosion..

When these alloys are immersed in biofluid, the chloride ion present in the fluid reacts with this passive film and forms pores in it further reaching to the fresh Ti6Al4V surface. As the external voltage increases, the speed of this reaction increases, leading to the pitting corrosion as depicted in Fig. 8b. The interaction of the fresh surface with chloride ion and water results into the formation oxide layer over surface (Fig. 8c & 8d) [42]. This reaction of Ti, Al and V with the chlorine ion and water, resulting in the formation of oxides are shown below [11][43]:

$$\:Ti+{{2Cl}_{3}^{-}}^{\:}+{2H}_{2}O\rightleftharpoons\:Ti{O}_{2}+{4H}^{+}+6{Cl}_{\:}^{-}$$

$$\:{2Al}^{3+}+{{2Cl}_{4}^{-}}^{\:}+{{2H}_{2}O}^{\:}\rightleftharpoons\:{Al}_{2}{O}_{3}+{6H}^{+}+8{Cl}_{\:}^{\--}$$

$$\:2V+{2Cl{\:}_{3}^{-}}^{\:}+{5H}_{2}O\rightleftharpoons\:{V}_{2}{O}_{2}+{10H}^{+}+6{Cl}_{\:}^{-}$$

The mechanism of pitting corrosion of the cast and DMLS produced Ti6Al4V samples in PSS, PBS and SBF lubricated conditions are confirmed by the SEM images and EDS analysis as shown in Fig. 9.

From SEM images it can be observed that pits were formed on cast as well as DMLS produced Ti6Al4V alloy. These pits are larger for sample corroded in the PSS (Fig. 9a & 9g) which may be due to higher concentration of chlorine ion which results in the dissolution of the oxide layer. Medium and small-sized pits were observed on samples corroded in PBS (Fig. 9c & 9i) whereas very fine pits were noticed on the for SBF case (Fig. 9e & 9k). Apart from fluid, the influence of manufacturing technique on corrosion can easily be noticed as DMLS produced Ti6Al4V alloy surface have relatively smaller pits (Fig. 9g, 9i & 9k) as compared to cast samples (Fig. 9a, 9c & 9e). The higher corrosion resistance of DMLS produced samples may be attributed to smaller grain size and presence of phases α’, which results into formation of relatively stable oxide layer on the surface as compared to casted one [36, 37].

These oxide layers act as a more effective barrier to penetration by chloride ions and thus minimize the corrosion rate. In SBF, the occurrence of pitting corrosion is slightly lower than the PSS medium. This may be because of the presence of sulfate and phosphate ions in the SBF solution that stabilizes its similar pH values [14]. The samples immersed in simulated body fluid exhibit capacitive behaviour at a wide range of frequencies, suggesting the presence of a stable passive region. However, the damage is more significant in the case of conventionally cast PSS. Severe cracks and dissolved surfaces can be observed from SEM image (Fig. 9a), which may be due to the penetration of hydrides [14]. It may also be due to the more penetration of thinner TiO₂ layer by chloride ions in case of casted Ti6Al4V.

The corresponding energy dispersive spectroscopy (EDS) of the corroded cast (Fig. 9b,9d,9f) and DMLS produced Ti6Al4V alloy (Fig. 9h, 9j,9l) shows that pitting corrosion is more prominent in PSS which results in the formation of oxides. Apart from the Ti, Al and V peaks, an additional peak for chlorine also appears because of the chloride ions concentrations in all three biofluids, which further results into the penetration of the oxide layer during the active dissolution. The high intensity of oxygen (O) peaks were also observed in the SBF and PBS biofluids, which may correspond to the complex passive film on the surface [32][44].

3.6 XRD analysis

The XRD pattern of the cast and DMLS produced Ti6Al4V are shown in Fig. 10. The XRD plot revealed that the Ti6Al4V alloy surface mainly comprised of α phases and ß phase which was in agreement with the initial analysis (section 3.1) performed before corrosion test (Fig. 3). This shows that no corrosion by-products (phases or compound) were formed on the Ti6Al4V surface in all three biofluids. The absence of corrosion products was attributed to the presence of the dense passive film on Ti6Al4V alloy surface, which act as a barrier to the corrosion process.

3.7 Surface wettability

To have a better understanding of the interaction between the corroded surface and the biofluids, wettability was conducted. This property plays a key role in the cell adhesion process and bioactivity. Utilizing the contact angle (CA) measurement, the wettability of the surface was estimated.

Figure 11 shows the biofluids droplet alongside the measured contact angle for corroded cast and DMLS produced Ti6Al4V alloy samples in all three biofluids. CA values for DMLS-produced Ti6Al4V samples were lower than those for the cast sample. The highest contact was measured in cast-produced Ti6Al4V in PSS solution and the lowest contact angle was measured in DMLS-produced samples in SBF.

3.8 ANOVA

The analysis of variance (ANOVA) was also performed on the results obtained to investigate the effect of the manufacturing process and biofluids on the corrosion rate. The ANOVA results for the corrosion rate are shown in Tables 7. The P value for all the considered parameters was less than 0.05, which shows their statistical significance on the response. The last column represents the percentage of contribution (P_r) to the overall variation, demonstrating the degree of influence on the response.

Table 7

Analysis of variance for the corrosion rate
Source	DF	Adj SS	Adj MS	F-Value	P-Value	Contribution (P_r) ((%)
Manufacturing Process	1	27.58	27.58	57.53	0	63.9
Electrolyte	2	08.89	4.45	9.27	0.003	20.6
Error	14	06.71	0.478	-	-	15.5
Total	17	43.18	-	-	-	100.0

In this section the effect of both the investigated parameters biofluid and fabrication process were discussed and related with the established theories.

4.1 Effect of biofluids:

As mentioned earlier, the corrosion rate of cast Ti6Al4V samples was found to be higher in PSS than the other two due to the presence of Cl- ion [45]. The Cl ions diffused rapidly in the oxide layer and corroded the substrate, resulting in the formation of pores in the Ti6Al4V specimen and increased the corrosion rate. The high corrosion rate of Ti6Al4V alloy results in pit formation on the surface is illustrated in Fig. 9. In simulated body fluid, the corrosion resistance of Ti6Al4V alloy surface is higher than that of the PSS. It is important to highlight that simulated body fluid simulates the conditions of medical implants while PSS is a more aggressive biofluids with a similar pH level that mimics the conditions of the human body condition. The pH levels are balanced by the presence of calcium and phosphate ions that ensuring the stable passivity in bodily fluids applications.

The corrosion behaviour of conventionally cast and DMLS produced Ti6Al4V alloys has been elucidated by both PDP and EIS analysis. The OCP studies have also conveyed facts regarding the passivation process, that was also observed from the PDP curve in the anodic zone for both cast and DMLS produced specimens in the various biofluids. From the obtained results it can be concluded that the corrosion of the Ti6Al4V alloy is prevented by the oxide layer (TiO₂) formed on the surface. The TiO₂ layer experiences a series of chemical reactions in the different biofluids, that depend on the surface conditions, corrosion current, potential and ions concentration, as well as the pH value. It was also represented that the corrosion resistance of the cast Ti6Al4V alloy was found to be inferior as compared to the DMLS produced Ti6Al4V alloy in all three biofluids. The order of the corrosion passivation current is as follows: ip_(SBF) < ip_(PBS) < ip_(PSS). All these facts are also confirmed by EIS results, as the Rs and R_f values of the conventionally cast and DMLS-produced Ti6Al4V alloy have showed similar trends. This was also confirmed by the SEM images (Fig. 9), that show very fine pitting in the case of SBF as compared to PBS and PSS.

4.2 Effect of fabrication process

Table 4 represents that the solution resistance (Rs) and film resistance Rf are higher for the DMLS produced Ti6Al4V alloy than the cast produced Ti6Al4V alloy. As a result, the DMLS produced Ti6Al4V alloy has a thicker film (as compared to the cast Ti6Al4V alloy) that leads to the formation of smaller pits. The SEM images (Fig. 3) reveal that the cast Ti6Al4V alloy specimen showed uniformly distributed α+ß grains, whereas the DMLS-produced sample has an acicular morphology generated during the manufacturing process [28][46]. The higher corrosion resistance of DMLS produced samples can be attributed to smaller grain morphology and the presence of phases α’ which results into the formation of relatively stable oxide layer on the surface as compared to cast one [36, 37]. The higher value of constant phase elements (Q) of the DMLS produced Ti6Al4V in all biofluids indicates a thicker oxide film on the surface. Moreover, higher semi-circle size and Rp value indicates a thicker passive film on the surface which results in higher corrosion resistance.

The ANOVA results from Table 7 also show that the manufacturing process has the most significant impact on the corrosion rate (63.9%), followed by biofluids (20.6%).

The corrosion behaviour of the DMLS-produced Ti6Al4V alloy was studied in various biofluids (PSS, PBS, and SBF). Further corrosion resistance was measured for DMLS-produced Ti6Al4V alloy using the OCP, PDP, and EIS tests and compared with the castTi6Al4V to evaluate its suitability for implant application. A dual layer Randle equivalent circuit was fitted to elucidate the mechanism of the corrosion process. The main findings from this investigation can be summarized as follows:

The morphology of DMLS-produced Ti6Al4V alloy consisted mainly of the acicular martensitic phase with a small amount of BCC Ti phase, whereas the cast Ti6Al4V sample exhibited a uniform distribution of HCP and BCC phases. This morphology affects the oxide film formation further, resulting into higher corrosion resistance for DMLS-produced Ti6Al4V alloy.

In SBF, the solution resistance and film resistance of the cast and DMLS produced Ti6Al4V alloys differed moderately; however, in PBS and SBF solution, the DMLS produced Ti6Al4V exhibited corrosion damage comparable to that of the cast Ti6Al4V alloy.

The OCP and PDP tests represent the presence of a passive layer of TiO₂ within the biofluids. For both types of samples, the passivation current showed the ascending trend in the sequence of PSS < PBS < SBF. Similarly, the pitting potential increased in the order: SBF < PBS < PSS.

The corrosion of DMLS-produced Ti6Al4V alloy in SBF has the lowest value of corrosion current density and the highest value of polarization resistance. This suggests a higher corrosion resistance rate in SBF solution as compared to the other media.

The PDP, EIS and EDS results showed the presence and dissolution of dense TiO₂ layers. PSS was found to be more aggressive than the PBS and SBF for both the cast and DMLS-produced Ti6Al4V alloy. Pitting corrosion occurred due to the dissolution of the oxide layer in PSS, with the more severe damage in the cast produced Ti6Al4V.

Contact angle was lower for DMLS-produced Ti6Al4V in SBF fluid which may contribute to bioactivity.

DMLS produced Ti6Al4V can be considered as a suitable candidate for orthopedic implant application.

Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

Author Contribution

Vipin Goyal: Conceptualization, methodology, investigation, data curation, formal analysis, validation, writing – original draft; Nisheeth Kumar Prasad: Visualization, writing – review & editing;Girish Verma: Supervision, visualization, funding, writing – review & editing.

Acknowledgement

The authors thank the Indian Institute of Technology Indore for providing characterization facilities and financial support. Mr. Vipin Goyal would like to acknowledge Dr. Vinod Kumar and Prof. I. A. Palani from IIT Indore for providing the corrosion and wettability testing facility.

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Experimental investigations into corrosion behaviour of DMLS manufactured Ti6Al4V alloy in different biofluids for orthopedic implants

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Abstract

Figures

1. Introduction

2. Material and methods

2.1. Sample preparation

2.2. Microhardness and density measurement

2.3. Microstructure and surface morphology characterization

2.4. Surface wettability measurement

2.5. Electrochemical corrosion testing

2.5.1 Open circuit potential (OCP)

2.5.2 Potentiodynamic polarization (PDP)

2.4.1 Electrochemical impedance spectroscopy (EIS)

2.6. Surface analysis of the corroded samples

2.7. ANOVA

3. Results and discussion

3.1 Analysis of phases and microstructure

3.2 Open circuit potential (OCP)

3.3 Potentiodynamic polarization (PDP)

3.4 Electrochemical Impedance spectroscopy (EIS)

3.5 Corrosion mechanism and morphology of corroded surface

3.6 XRD analysis

3.7 Surface wettability

3.8 ANOVA

4. Effect of parameters on corrosion rate

4.1 Effect of biofluids:

4.2 Effect of fabrication process

5. Conclusions

Declarations

Declaration of competing interest

Author Contribution

Acknowledgement

References

Additional Declarations

Status:

Version 1