Tribological Behavior of Atmospheric Plasma Sprayed HA-CNT Coatings of Biomaterials

Abstract

A MWCNT-induced hydroxyapatite, air plasma spray coatings were produced on SS-316 L, CoCrMo, and Ti6Al4V alloys at varied weight percent ratios. The shape, thickness, adhesion, structure, and content of the APS-treated samples, as well as particle distribution, were studied using an X-ray diffraction, scanning electron microscope, Fourier transform infrared, and 3D-profilometer. The study looks at dry and wet unidirectional sliding wear behaviors, as well as the influence of incorporating carbon nanotubes in different weight percent to hydroxyapatite by plasma spraying on its tribological properties in physiological condition. In-vitro investigation was conducted in controlled environment to model complicated interactions among cells growth factors. Both CNT and HA particles were absorbed into the APS layers, as evidenced by the results. Crystallinity and volume percent with open porosity were substantially higher in the APS-CNT imposed HA coating than in the control. When compared to pure HA coating, the enhanced hardness ranged from 2.4 to 5.6 GPa, the modulus of elasticity ranged from 105 to 172 GPa, and the fracture toughness from 0.6 to 2.4 MPa.m1/2, with a reduced wear rate of 50.2 × 10−5 mm3 to 4.2 × 10−5 mm3 N−1 m−1. CNT addition has no negative effect on osteoblast proliferation and cell viability.

1. Introduction

The fundamental requirement of bio implant should be its biocompatibility with the biological system, high mechanical properties to withstand various stresses induced, and excellent corrosion resistance in the body fluid [1]. Materials used for bio-implants are categorized as Metals and alloys, Polymers, ceramics, and reinforced composites. Advanced materials like Surgical stainless steel, CoCrMo alloys, and different Titanium grades are the most commonly used implant material due to their good biomechanical properties. But their main limitation is that these materials may not always be biocompatible with the body fluids, and tissues may not grow on these materials after they are implanted in the body [2]. The recurring failure of traditional materials used in orthopedic implant manufacturing was due to a lack of or inadequate integration of implant materials to the juxtaposed bone & stress–strain imbalance between the interface of tissue and implant material [3].

Ceramics are commonly used for bio implant applications due to their superior biocompatibility in the body environment. Due to their similarity with human bone natural tissues easily grow on their surfaces [4]. Few bioceramic materials are HA, tri-calcium phosphates (TCP), and bioactive glasses. But uses of ceramics are limited due to their poor mechanical properties.

Polymers are soft materials that can be easily formed into complex shapes; however, they have poor mechanical properties and cannot be used in heavy load-bearing applications such as knee and hip prostheses [5]. Composite materials are typically created by joining two or more distinct material phases, such as metallic-ceramic, polymer-ceramic, and metallic-polymer [6]. These materials have good mechanical properties and biocompatibility, but it is extremely difficult to make composite parts biocompatible [7].

Hydroxyapatite (HA) coating is an excellent aspirant for bio-implants in order to improve their biocompatibility in body fluid [8]. Because the chemical structure of HA is very similar to the structure of natural bone, it can form new tissues on it in body fluids, which is critical for the fixation of bio-implants inside the body and also protects the body from any harmful metal ion released by the metallic implant [9]. It has lower mechanical properties, such as low bending strength, fracture toughness, and impact strength, and thus cannot be used in pure form for load-bearing applications [10]. The coating of HAP alters the dispersion behavior of MWCNTC in the PP matrix, causing variations in the tensile and thermomechanical characteristics of MWCNTs/PP composites [11].

Hence it is a good idea to overcome these mechanical limitations to coat this material on the metallic substrates which have good mechanical properties. From the previous work [12] it has been found that pure HA coating results in poor adhesion, bonding strength, and other mechanical properties. Hence to improve these properties reinforced materials like zirconia, silica, titania, alumina, carbon, and boron nanotubes are added in bulk HA. Particularly there must be sufficient adhesion strength between HA coating and metallic implants to avoid spalling, cracking and wearing of the coating [13]. Many techniques like electrophoretic, sol–gel, and thermal spraying (plasma and HVOF), etc. are used for HA deposition on metallic substrates. Among these, the air plasma technique is widely used because of its excellent adhesion strength, better process control, crystal structure, and thickness of the coating as compared with other such coating techniques [14]. This technique is recommended and approved by Food and Drug Administration (FDA) in the USA for clinical trials [15].

Despite the many advantages of hydroxyapatite coating on metallic substrates, the brittle nature and low strength of hydroxyapatite delay clinical trials under varying loading conditions [16]. Because HA is brittle, secondary reinforcement materials such as Zirconia, Ni3Al, Alumina (Al₂O₃), carbon nanotubes (CNTs), boron nitride nanotubes (BNNTs), Silica, yttria-stabilized zirconia (YSZ), and Ti-grade-alloys are commonly used to improve mechanical properties such as fracture toughness and Young’s modulus [17]. CNT reinforced composites outperformed other composites in terms of biological and tribo mechanical properties. The addition of CNTs increases the crystallinity, shape, and biological characteristics of HAP [18]. The COF of a composite can be reduced by varying the rate of CNT. The inclusion of CNTs boosts cell growth and adhesion. Wear-related mass loss is reduced by 13.33%, 66.67%, and 83.33% [19]. To understand the functionalization of CNTs to increase their hydrophobic qualities, the intrinsic nature of CNTs, including their physical and chemical properties, is explained. CNTs are often functionalized with different functional groups (such as-OH and –COOH) using covalent and non-covalent approaches to increase dispersity in aqueous conditions and limit toxicity [20]. The specimens’ corrosion resistance reveals the biocompatibility of Ti–Co–Cr as being more unique due to the presence of titanium metal [21]. Many corrosion, wear, and abrasion resistant applications exist for plasma sprayed Al₂O₃-TiO₂ ceramics in industries such as aircraft, textile, and automotive [22]. Gell et al. [23] investigated plasma-sprayed nanostructured 13 wt% Al₂O₃-TiO₂ coatings and reported that these coatings had been approved for submarine and shipboard applications by the US Navy.

It is well understood that with HA coating wear indicates the possibility of orthopedic implant degradation failure. According to previous research, the main issue that affects the total joint replacement longevity is “worn particle-induced osteolysis, particularly adjacent to acetabular components” [24]. Wear particles are generated by the articulating surface and migrate along the implant interface, causing osteolysis, implant loosening, and, eventually, implant failure. However, there was no detailed information on the mechanism of microstructural failure. As a result, investigating the tribological behavior of HA-CNT coated specimens at the microstructure level requires considerable effort. The effect of CNTs on the human body and the environment is not well established, owing to a lack of standardization of toxicological tests, which leads to discrepancies in the results [25]. The current work will investigate the microstructural properties of applied coatings and investigate the potential of HA-CNT coating by plasma spray for bio implant applications. In this work, an attempt was made to deposit HA with 10% and 5% CNT on surgical grade SS 316L, CoCrMo, and Ti6Al4V substrates and to establish their suitability for in vivo applications.

2. Experimental procedure

2.1 Preparation of substrates and feedstock powder coating processes

The alloy samples of SS 316L, CoCrMo & Ti6Al4V were cut to form an approximate disc having 40x5 mm diameter size. The surface of coated samples is prepared using alumina (Al₂O₃) of size 65–90 μm in a grit blaster within a blasting pressure range of 0.50 MPa for better adhesion of coating powders. After grit blasting the samples were air blasted to clean the surface from alumina contamination before final coating. Table 1 shows the chemical composition of HA-CNT powder for coating substrate materials. Commercially available HA powder was used for coating deposition which was procured from Xi’an Prius Biological Engineering Co., Ltd. XS province, China, and MWCNT procured from Adnano Technologies Pvt. Ltd., Karnataka, India was used as reinforcing material. Reinforcing material was mixed in pure hydroxyapatite Powder in 10% and 5% by weight using laboratory ball mill up to 10 hours for uniform mixing of reinforced HA coating at Plasma Spray Coats, Bangalore, India. Particle size, composition (wt.%) with powder designation are shown in Table 1.

Coating powder	Composition (wt. %)	Particle size	Designation
Pure HA	Ca10(PO4)6OH2	100–160 μm	Pure HA
Reinforced HA	Ca10(PO4)6OH2—MWCNT (10 wt% & 5 wt%) mixed in laboratory ball mill for 8 h.	Reinforcement 10–45 μm	Reinforced HA-C

Table 1.

Chemical composition and other details of HA-CNT coating powder.

Morphological studies of bare hydroxyapatite powder with carbon nanotube reinforcement on coated alloy substrates have been carried out using scanning electron micrographs. These powders were also characterized by the XRD technique using the Xta-LABminiTM Benchtop X-ray Diffraction/Crystallography system from Rigaku Americas Corporation in the angle range of 20°^–90°. Figures 1 and 2 depict the micrographs and EDX analysis of (10 wt% and 5 wt%) reinforced HA-C coating. SEM micrograph shows that reinforced material has crushed angular and ball shape.

Figure 1.

Scanning electron micrograph showing morphology of 10 wt% reinforced HA-CNT powder with element composition from EDX analysis.

Figure 2.

Scanning electron micrograph showing morphology of 5 wt% reinforced HA-CNT powder with element composition from EDX analysis.

2.2 Wear analysis

Ball/Pin on disc tribo tester (Ducom Instruments, model TR-20LE) was used for bare and HA-CNT coated samples under dry and wet (SBF) sliding conditions for analyzing the abrasive wear. The samples were prepared as per G95-99a. The effect of the counter body of Steel Ball SAE52100, H = 9.46 GPa was studied to analyze the counterparts. The roughness (Ra) values of testing samples were made in the range of 0.8 μm or less. A varying speed of 100 RPM with a 25 mm circular track radius with a total traveling distance of 90–200 m was used to study the wear at a macro level. A specific linear speed of the tribo probe was set to 10 mm/s. The counter body (probe) used was an 8 mm diametric steel ball. The inbuilt LVDT sensor gives the value of linear force in between the coated surface and steel probe under the depth of wear track. The coefficient of frictional data is acquired at the 17 Hz frequency level. Alicona 3D profilometer was used to obtain the profiles of wear track on tested samples. The value of wear volume is computed using the wear track depth profile.

Figure 3 shows the pictorial view of wear (tribological) study in physiological conditions. The tests were performed with samples immersed in externally supplied simulated body fluid (SBF), in the tribo tester with the same test parameters and conditions as was used during the dry wear test. The SBF is prepared and purchased from KET’s Scientific Research Centre Mumbai, India using Takadama Hiroaki [26] with the same ion concentration of chemicals as compared with human blood. The chloride ions are the cause of corrosion biomaterials, so these fluids are suitable for analyzing the corrosion resistance [27].

Figure 3.

Experimental set-ups for tribological wear test (a) wear test mechanism under SBF (b) wear test sample in physiological condition.

2.3 Coating characterization

The critical evaluation of SS 316 L, CoCrMo, and Ti6Al4V substrates coated by 10 wt% and 5 wt% CNT and worn-out surfaces are done. The metallographic study of as-sprayed coatings on these substrates is examined under ZEISS Gemini field emission scanning electron microscope (FE-SEM), operating at 18 kV. To have high-resolution imaging, gold is used as the sputter coating material before analyzing in SEM. It can be seen from the optical micrograph Figure 4 that the top surface of the coating is free from cracks and macro-level porosity. Most of the splats are well-formed without any sign of disintegration. Some melted grains are also visible on the surface of coatings in most of the micrographs. Further, it can be observed from the SEM/EDX analysis that a whitish appearance somewhere in the coating in SEM micrographs indicates Ca-rich HA particles. In reinforced HA coatings some streaks are detected whose EDX analysis confirms the presence of reinforcing content which is distributed in the matrix of the HA-C coatings [28, 29]. EDS spectroscopy indicated a Ca/p ratio of 1.62 in calcium phosphate deposits, which is similar to the ratio of 1.64 in natural bone [30].

Figure 4.

Optical micrograph of reinforced HA-C coating on Ti6Al4V alloy with element composition from EDX analysis.

Figure 5 shows an insight on the FTIR analysis of HA-CNT reinforced coated samples in the range of 4000 cm⁻¹ to 500 cm⁻¹. OH– group at 3568 cm⁻¹ and 652 cm⁻¹ can be seen with the presence of PO₄³⁻ group at 960 cm⁻¹, 1033 cm⁻¹ and 558 cm⁻¹. CO₂ molecules which were not visible in the original powder are now can be observed at 2360 cm⁻¹. The intensity of OH⁻ group has become comparatively weaker after coatings in the case of reinforced HA-C coatings in all alloy substrates.

Figure 5.

FTIR spectroscopy for plasma sprayed reinforced HA-C coating powder (a) 10 wt% CNT.

The basic mechanical property of HA-C coating is microhardness, which may play an important role in bio implant application. At the coating interface, the microhardness of coatings on alloy substrates was measured. Profiles for microhardness with distance from the coating substrate interface are shown in Figure 6. The hardness of both (10 wt% & 5 wt%) reinforced HA-C coating is lowest (310 Hv, 340 Hv, and 402 Hv in case of uncoated SS 316 l, CoCrMo, and Ti6Al4V, whereas 322 Hv, 352 Hv, and 418 Hv in case of reinforced HA-C respectively) at the interface in case of all the three alloy substrates. Slight improvement in hardness was observed with the reinforcement content. Enhanced Vickers hardness resulted in improved hydrophobicity for a 5% HAP/MWCNT coating [31]. The impression of Vickers indent Figure 7 was observed under SEM for an accurate radial crack length measurement.

Figure 6.

Micro hardness profiles of reinforced HA-C coatings on SS 316L, CoCrMo, and Ti6Al4V alloy along the cross-section.

Figure 7.

Radial cracks on HA-CNT coating caused by microindentation.

2.4 Bio-compatibility test of the coating

A direct contact method was used to perform cytotoxicity test with test samples. Samples of Pure SS-316L, CoCrMo, and Ti6Al4V, 5 wt% coatings, and 10 wt% coatings as per ISO 10993-5 are prepared. The test was done on 6 well plates. A fresh medium is used to replace the culture medium from the 3T3-L1 monolayer. On the cells test samples, negative & positive controls in triplicate were placed. After incubation at 38°C for 2, 15, and 30 days, the cell monolayer was examined microscopically for the response around the test samples. The reactivity was based on the zone of lysis, vacuolization, detachment, and membrane disintegration as per norms. Hemocytometer was used for viable cell count as shown in with cell suspension >90% viability for each experiment. MTT experiments revealed that graphene-doped hydroxyapatite composite had a substantial influence on 293T cell growth and excellent biomimetic mineralization, as demonstrated by in vitro bioactivity studies.

3. Results and discussion

3.1 Characterization of uncoated and plasma sprayed HA-C reinforced coating

Figure 8a revealed the microstructure study of surgical stainless steel SS-316 L, as equiaxed austenite grains which are as per the grade of this steel [32]. Annealing twins are also present in the structure which reveals the resemblance of the structure discussed in the handbook. SS-316 L is austenitic stainless steel where ‘L’ denotes low carbon content which is limited to 0.03%. A carbide-free austenitic and strain-free microstructure is produced in this steel that retains in the soluble carbon at an annealing temperature of 1010–1065°C (1700–1850°F) due to rapid quenching. Figure 8b depicts that there exists a hexagonal close-packed and face-centered cubic crystalline structure with CoCrMo alloy. Typically, at normal temperature, the face-centered cubic phase is predominant causing fcc to hcp transformation. The microstructural grain consists of bigger Co dendrites having an hcp structure with the smaller surrounding zones correspond to the embedded carbides which are recognized as M23C6 type. The analysis suggests that there is uniformity in porosity with fine grain structure. The microstructure of Ti6Al4V Figure 8c is compared with the standard microstructures from Metals Handbook (1975) and ASM Handbook (1992, 2001). The microstructure consists of equiaxed α grains (light) and intergranular β (gray), which corresponds to the grade of this alloy. EDAX analysis confirms the CA/P ratio in between 1.67–2.4 which is the desired ratio for long-term applications of the implant [33]. CNT–HA composites can generate bone-like apatite on the surface of the sample due to greater calcium ion concentration in SBF and higher negative charge, as well as more accessible nucleation sites.

Figure 8.

Optical micrograph of HA-C coating on (a) SS 316L, (b) CoCrMo, (c) Ti6Al4V.

It has been reported in the earlier research work [34] that thicker coatings have been more durable when implanted in the body environment. But thickness more than the critical limit leads to self-disintegration of the coatings and this would lead to the problem of increased mechanical competence of coating [35]. High crystalline coating with higher thickness may lead to the formation of brittle material causing cracking under shear or bending forces. So it was aimed to produce coatings having thickness 0f 150–200 μm. Some randomly selected coated samples were measured along the cross-section for determining the coating thickness and reported in Table 2.

Sr. No	Coating (APS)	Substrate	Coating thickness (μm)
1	Reinforced HA-C (10 wt%)	SS 316 L	178
2	Reinforced HA-C (5 wt%)	SS 316 L	184
3	Reinforced HA-C (10 wt%)	CoCrMo	188
4	Reinforced HA-C (5 wt%)	CoCrMo	190
5	Reinforced HA-C (10 wt%)	Ti6Al4V	190
6	Reinforced HA-C (5 wt%)	Ti6Al4V	193

Table 2.

Average coating thickness of plasma sprayed HA-C coatings.

The porosity of the coating in bio implant application has an important role. Earlier literature proposed that the surface of the coating should be free from porosity and coating should act as a railing between the substrate and body fluid environment to avoid metal ion release in the body [36]. A slight decrease in porosity is observed with the addition of reinforcement of 10 wt% and 5 wt%. This decrease in porosity is due to the small size of reinforcement. The slight reduction is in agreement with the findings of Morks, [37, 38]. APS coated Ti samples have 0.40%–1.35% porosity, and in contrast to the higher side, the SS & CO samples exhibited marginally more porosity of the range 0.45%–1.40%. The surface of the bio implant must be rough enough for easy cell growth which improves the fixation of the bio implant in a physiological environment. Both Ra and Rz values of the reinforced HA-C coating on ally substrates are measured. Average values of surface roughness (i.e. Ra and Rz) of HA-C coatings on SS-316 L are 6.32 μm and 33.83 μm respectively, whereas for Ti6Al4V these are 6.06 and 39.41 μm respectively. It can be seen from the results that the average value of surface roughness (Ra) for uncoated samples is slightly lower than that of reinforced coatings on all the alloy substrates.

Figure 9 represents XRD spectra of HA-C coatings of samples tested under wet (SBF) conditions. All peaks in the case of coating correspond to HA. One small broad peak of CaO is observed in all reinforced HA coatings which confirms the presence of amorphous CaO in a very minute amount. An important observation shows the presence of tricalcium phosphate (α-TCP, β-TCP) and tetra calcium phosphate (TTCP) peaks between 29° to 32° angle range on all as-sprayed coatings. As the original powders do not have these phases. The severe failure and destruction of the implant were caused due to the rapid solubility of the amorphous phase in the human body’s physiological environment. The XRD analysis also depicts the sharp HA peaks which turn to be broader after coatings indicating the corrosion of crystalline materials to the amorphous phase during spray coating.

Figure 9.

XRD pattern for plasma sprayed reinforced HA-C coated wet samples. (a- 10 wt% & b- 5 wt% wet) HA, a-αTCP, b-βTCP, t-TTCP, c-CaO, g-graphene, #-Al₂O₃, $-TiO₂.

Coating Crystallinity was calculated at a 20° to 90° angle. Random HA-C coated samples were analyzed for average values and presented in Table 3. The fraction percentage of the amorphous (TCP & TTCP) phase was also shown.

Sr. no	Coating	Substrate	Crystallinity (%)	TCP, TTCP phases (%)
1	Reinforced HA-C	SS 316L	66.8	14.45
2	Reinforced HA-C	CoCrMo	66.9	14.24
3	Reinforced HA-C	Ti6Al4V	68.5	14.12

Table 3.

Crystallinity, TCP and TTCP phases in reinforced HA-C coatings.

3.2 Corrosion and sliding wear performance

3.2.1 Corrosion (electrochemical polarization) test

Corrosion behavior of all coated samples are evaluated by electrochemical polarization test in simulated fluid as proposed by Kokubo and Takadama [39]. The primary reason to use Tafel polarization was to evaluate corrosion current density of all HA-CNT coatings in simulated body fluids whereas Potentiodynamic measurement were carried out to study the passive behavior of the coatings. Tafel polarization measurement were carried out starting from −250_OCP mV to +250_OCP mV with a scan rate of 1 mV/s, whereas Potentiodynamic polarization were carried out starting from -250_OCP mV to 1600_SCE mV. Table 4 represents the polarization test results. The corrosion current in the case of SS 316 L is more as compared to that of CoCrMo and Ti6Al4V, hence it can be inferred that corrosion resistance of Co and Ti-based alloy is more than that of SS. Also, the passivation range of Ti alloy obtained from potentiodynamic curves is comparatively larger than that of SS and Co alloy, which supports the results of the Tafel polarization method. The results obtained from Potentiodynamic polarizer suggests that all as sprayed HA-CNT coatings show passive behavior in simulated body fluids which support the low current density presented by Tafel polarization method.

Specimen	E corr. (mv)	I corr. (μA)	βa (mv)	βc (mv)
SS 316L	−497.814	1.432	102.432	40.870
CoCrMo	−420.122	1.477	102.490	40.015
Ti6Al4V	−247.365	1.598	105.415	133.322

Table 4.

Result of Tafel polarization test of reinforced HA-C coating on SS 316L, CoCrMo and TTTi6Al4V in SBF solution.

3.2.2 Wear (Abrasive wear) and friction mechanism of bare and APS-coated specimens against steel

Morphological structure and hardness are the two main parameters affecting the wear resistance of any coatings. Figure 10a and b represents the micrograph image of the wear mechanism of the Ti sample during uniform directional sliding of HA-C coated specimen. SEM micrograph after wear test exposure shows uniform and continuous wear. It can be observed that wear tracks are visible due to abrasive action on the both uncoated and coated surface. The micrograph also shows the small debris by abrasion. SEM images revealed the presence of areas of cracks and fractures due to plastic deformation, which is attributed to abrasive wear [40]. In the HA-CNT wear track, the image depicts smaller craters and a huge area of abraded surface. Fracture and chipping form the crater, while abrasive wear results in the rough surface.

Figure 10.

Wear scar micrograph of APS- sprayed HA-C coatings against steel ball (a- wear mechanism & b- wear path).

Average Weight loss (kg/m²) of uncoated samples was 0.14, 0.12 and 0.09 kg/m² respectively whereas for reinforced HA-C coating it was 0.24, 0.22 and 0.18 kg/m² respectively. It may be observed from the wear results that weight loss of metallic substrates is comparatively lower than that of all as-sprayed coatings on alloy substrates. It can be observed from the plotted results from Figure 11 that, the wear volume was reduced up to 75% with the improvement in the wear resistance of reinforced HA-C coatings with 10 wt% and 5 wt% reinforcement. This improvement is because of the increased fracture toughness and elastic modulus.

Figure 11.

Variation in wear loss (kg/m²) of uncoated, 10 wt% and 5 wt% HA-C coatings alloy in dry and wet (SBF) condition.

The average wear rate and friction coefficient fall into a near mild wear regime, with values ranging from 10 to 7 to 10–4 mm³/Nm [41]. The tribological test parameters were chosen with the conditions of medical implants working inside the human body, specifically the failure of the femoral head within the acetabular cup of the hip joint. Normally 0.8–2.5 MPa of stress is considered in a hip joint while walking [42]. The HA-C coating has to withstand a maximum frictional force for a minimum period of 20 years [43]. To compensate the entire working life a high value of 10 N load is kept during the entire test.

Surface hardness, toughness, spraying technique, and contact pressure are considered as the significant effecting factors on the coefficient of friction value. Figure 12 represents the accumulated mean CoF from 0.35 to 0.42 due to the varying contact pressure of the steel ball. The effect of the spray technique is the drop in mean CoF up to 0.37 for the Ti coated samples in wet (SBF) conditions. The total mass loss study demonstrated that the presence of PBS in the electrolyte solution tended to exacerbate the deterioration of the alloy under the circumstances examined, which was also verified by the 3D analysis of the wear track geometry [44]. As hardness of counter body governs the tribological failure, so it primarily influences the specific wear rate. Ti6Al4V alloy has shown maximum wear resistance among all uncoated and coated specimens. Improvement in wear resistance may be observed with the incorporation of reinforcement in the case of HA coatings. There is a slight decrease in CoF from 0.8 to 0. 6 with the addition of reinforced CNT into the HA matrix. As the graphene layer is peeled off from the CNT surface it offers lubrication causing the decrement in CoF value.

Figure 12.

Average CoF curves for APS coating against steel ball SAE 52100 at contact pressure of 2000 MPa. (a- sliding distance vs CoF & b- time vs CoF curve).

A tensile force ≥11 GPa is required to remove the single graphite layer from multiwalled CNT along its axial direction [39]. The evaluated tensile stress in the wear track in the current study was ~12 GPa, which was found to be sufficient for the removal of the graphene layer from CNT. Similar results were seen when the test was carried out under SBF immersed solution. As the body fluid would offer additional lubrication on the implant surface it would decrease the amount of wear debris thus, expecting better performance of coating inside the human body. Tribo Mechanical wear is the most common wear mechanism, as seen by abrasion, adhesion, and cracking [45].

Figure 13 represents the wear depth profile and 3D topography of HA-C coating against SAE 52100 steel. The wear volume against the track was calculated (using average value of cross-section area from three variable experimental sets with track diameter) from the profile and found to be 0.25 mm³ for HA-CNT coating. With CNT addition it decreases up to 75% resulting in a “reduced probability of disturbance in the biological environment around the implant”. The 15.70 μm of wear depth is recorded which is maximum for both types of coatings versus a steel ball.

Figure 13.

(a) Wear depth profile and (b) 3D topography for APS coated sample.

3.3 Bio-compatibility evaluation of HA-CNT coatings

MTT assay was used to determine the mitochondrial activity and cellular viability of 3T3-L1 viable cells on uncoated and reinforced HA-CNT coatings. When incubated with viable cells, the reduction reaction of MTT reagent causes it to be reduced into purple formazan crystal. The cell viability is indirectly reflected by the absorbed formazan crystal. As a positive control, cells seeded on tissue culture plates are used.

3.3.1 Cell viability study

Morphology of cell-seeded on bare SS-316 L, CoCrMo, and Ti6Al4V substrates after 2 days were analyzed. Cell viability (%) is shown in the histogram presented in Figure 14, after 2, 15, and 30 days of incubation. It may be observed that cell viability on Ti6Al4V substrate was 72%, 81%, and 88% respectively whereas it was 68%, 78%, and 85% on CoCrMo and 66%, 78%, and 84% respectively on SS-316 L. Further, cells were comparatively less viable initially on all the three alloy substrates than that of the grown cells on the tissue-cultured plate (p,0.05), but viability increased with exposure time duration.

Figure 14.

Viability of 3T3-L1 cells on Ti6Al4V, CoCrMo and SS-316L alloy from 2 to 30 days.

The morphology of 3T3-L1 cells seeded on reinforced HA-CNT with 10 wt% and 5 wt% coatings (as-sprayed) on all Ti6Al4V alloy samples is shown in micrographs presented in Figure 15a–d, respectively. According to the viability results, cell viability of 5 wt% reinforced HA-CNT coatings is lower than that of 10 wt% coatings after 2 and 15 days of exposure, respectively. However, after 30 days of exposure, the viability of all as-sprayed coatings was comparable.

Figure 15.

SEM (BSE) image of 3T3-L1 cells seeded on (a) control (b) bare Ti6Al4V (c) 5 wt% Ti6Al4V (d) 10 wt% Ti6Al4V.

3.3.2 Cell proliferation study

Following the Cellular viability study viable cell count of 3T3-L1 cells on bare and coated substrates, the alloy was determined by staining them with 0.4% trypan blue at 1:5 ratios and counting them using a hemocytometer for cell proliferation study. Each experiment used a cell suspension with greater than 95% viability, as discussed in the experimental procedure. The difference in proliferation rates was found to be significant for alloy substrates and tissue culture plates (p < 0.05). Cell count on tissue culture plate (i.e. 110, 500, and 1500) was remarkably greater than that of alloy substrates (i.e. 48, 288, 365 for Ti6Al4V, 40, 218, 320 for CoCrMo and 32, 216, 310 for SS-316 L) after 2, 15 and 30 days respectively. A similar pattern with a slight increase in cell count was also seen in reinforced HA-CNT coatings with different wt% ratios. The in vivo bone creation of CNT-coated Nanofillers was dramatically accelerated by greater bone mineral density and up-regulated osteogenic signs (ODN, OCN, BMP2) of bone-forming cells [46].

The reason for higher proliferation on HA-CNT coating as compared to the uncoated metallic substrate is justified by previous researchers that “hydroxyapatite once implanted into the body can easily react with physiological fluids and form a tenacious bond to hard and soft tissues through cellular activity”. Another cause that gives an idea about the superior proliferation rate of coatings is the difference in surface roughness between porous HA-CNT coating and smooth bare metallic alloy substrates. Cytocompatibility and Osseointegration of the bone cement can be controlled by adjusting the MWCNT coating [47]. Furthermore, it may be observed from the results of this study that the proliferation rate on a tissue culture plate was comparatively higher than on all HA coatings due to their special design to enhance cell growth.

4. Conclusions

The HA-CNT coatings with different wt.% was processed by air plasma spray on alloy samples. The characterization and tribological properties comparatively were investigated:

• HA-CNT composites could be successfully synthesized using the air plasma spraying method. Reinforced CNT has not shown any adverse effect on densification. The uniform and improved densification of hydroxyapatite microstructure are because of the high electrical and thermal conductivity of nanotubes.

• The computed tensile stress in the wear track was ~12 GPa which, was achieved because of the well retained peeled off graphene layers. Both the 10 wt% and 5 wt% HA-C coatings form dense morphological structures. Due to high interstitial bonding of HA with CNT matrix the cracks were properly bridged and fracture energy was fully absorbed causing a better hardening mechanism.

• The wear mechanisms revealed by APS-sprayed HA-C coating on SS 316 L and CoCrMo that, due to higher contact stress a small amount of carbon adhesion with minor plastic deformation of austenitic grains took place. However, in the case of Ti6Al4V coatings exhibited brittle cracking, spalling and abrasive grooves.

• An improved wear resistance by 75% with the lesser volume of wear debris generation was seen as the 10 wt% and 5 wt% CNT reinforcement was added to plasma spray HA coating. The lubrication provided by peeled off graphene layer from CNT decreases the CoF on the coated surface. CNTs content reduces the wear particle size (HA-CNT: 0.15–3.4 μm).

• Contact pressure, hardness of counter body, spraying technique, and coating reinforcement with Hydroxyapatite are the major influencing parameters from the investigation.

• Biocompatibility of all HA-CNT coatings was improved and is relatively higher than that of bare alloys.

Control > as sprayed reinforced 10 wt% HA-CNT coatings > as sprayed reinforced 5 wt% HA-CNT coatings > Bare Ti6Al4V > CoCrMo > SS-316L.

Therefore, the outcome of the present research work could be used to improve the efficacy with enhanced service life of medical implants and other such applied components. A very attractive solution to existing problem of HA with different reinforcement material could be established by HA-CNT composites and coatings. The process of HA-CNT coatings with different wt% compositions are user friendly and capable to provide alternative solutions to other high end coating techniques and reinforcement materials.

Acknowledgments

The authors would like to express their gratitude to Plasma Spray Coats in Bangalore, India, for providing the coating facility. The authors would like to thank VJTI, Mumbai, as well as IIT Bombay, India, for providing testing facilities such as SEM, EDX, XRD, Nano-indentation, and Non-contact profilometer. A special extended thanks to Sree Chitra Tirunal Institute of Medical Science and Technology, Kerala for helping in performing the in-vitro biocompatibility test. This research is not supported by any financial grant from a public or private funding agency.

Disclosure statement

No potential conflict of interest was reported by the author(s).