Calcium Phosphate Coatings Deposited on 3D-Printed Ti–6Al–4V Alloy by Plasma Electrolytic Oxidation

Abstract

In this article, the process of creating calcium phosphate coatings through plasma electrolytic oxidation was investigated. Calcium phosphate coatings were deposited onto titanium substrates fabricated via the selective laser melting (SLM) method. The correlation between the characteristics of the coating and the applied voltage (200, 250, and 300 V) of PEO was studied. The surface morphology analysis indicates that an increase in applied voltage results in a larger pore size. It was discovered that, when a voltage of 300 V was applied, a layer of hydroxyapatite formed. However, at 300 V, the coating cracked, producing a significantly rough surface. Our analysis of the elemental composition of sample cross sections indicates the presence of TiO₂ layers that are enriched with calcium (Ca) and phosphorus (P). The coefficient of friction and wear rate are primarily influenced by the morphology, pore size, and density of the titanium dioxide layer. Furthermore, a rise in the quantity of the beta phase of the titanium on the surface can be noticed as the applied voltage increases. As a result, it also affects the mechanical and tribological characteristics of the coating. The sample treated to a voltage of 250 V demonstrates a higher resistance to wear and a lower elastic modulus in comparison to the other two coatings.

1. Introduction

At present, additive manufacturing offers new possibilities for producing personalized implants, prosthetics, and bone scaffolds with varied structures [1,2,3]. The selective laser melting (SLM) method is particularly well-suited for metallic materials, due to its ability to produce parts with a high density, complex geometries, and excellent mechanical properties. Metals (gold, titanium, tantalum, magnesium, zirconium, etc.), metal alloys (Ti–6Al–4V, 316L, Co–Cr, and Zr–2.5 Nb), polymers (polylactic acid, polyethylene, etc.), biocompatible ceramics (carbon and hydroxyapatite), and composite materials are frequently employed for biomedical applications [4,5]. Titanium (Ti) and its various alloys form the majority of implants. For hard tissue implant replacement, implants made from titanium and Ti alloys are preferred, because of their remarkable biocompatibility, low density, and high specific strength [6]. Despite their remarkable qualities, titanium and its alloys have limitations, such as poor biological activity, wear resistance, and corrosion resistance [7]. Thus, in order to solve these limitations, the current research mainly focuses on modifying surfaces through the application of a bioactive layer, predominantly based on calcium phosphates [8,9,10].

A variety of methods can be utilized for depositing bioactive calcium phosphate coatings onto a metal surface, such as sol–gel, plasma spraying, anodizing, physical vapor deposition (PVD), and chemical vapor deposition (CVD). The sol–gel method relies on the hydrolysis of a solution containing a precursor, leading to the formation of colloidal particles that are suspended in the solution. A gel-like substance is created after the condensation stage. The sol–gel method is a universal method, as it allows us to obtain a wide range of materials; however, its main disadvantages are low adhesion and the potential toxicity of precursors [11]. Plasma spraying is a process in which hot ionized gas is used to melt powder, and, as a result, the powder particles readily adhere to and agglomerate on the substrate surface. Plasma spraying can be used to coat many types of materials; however, there are limitations when coating complex-shaped implants, and the high temperatures during deposition can lead to phase changes that can affect the mechanical properties of the material [12]. Physical vapor deposition involves the deposition process occurring without any chemical reactions, while chemical vapor deposition involves film formation resulting from the chemical reaction of the precursors on the substrate. These technologies have comparatively low coating speeds, considerable process intricacy, and challenges in achieving uniform coating deposition control [13]. Anodizing is used for the deposition of an oxide layer on the surface of a titanium alloy (Ti–6Al–4V), in particular, on the implant’s surface, due to the necessity of protection from the toxic influence of Al and V and from corrosion and wear. However, the anodizing method does not provide sufficient protection against corrosion and wear [14].

However, plasma electrolytic oxidation (PEO), also referred to as micro-arc oxidation (MAO), is regarded as the most appropriate method to achieve this goal. The PEO process is based on electrochemical and physicochemical interactions that occur at the interface of an anode and an electrolyte when the applied voltage exceeds the threshold for dielectric oxide layer breakdown. The PEO process involves the generation of short, limited discharges of electricity and the formation of oxygen on the anode surface, with the release of significant heat, melting, and the subsequent solidification of the metal. The PEO process allows us to obtain a porous oxide layer with good adhesion to the substrate [8,15].

Titanium dioxide appears in anatase and rutile phases and can be formed using PEO. Anatase has been associated with superior cytocompatible behavior in terms of osteoblast proliferation and adhesion [16,17]. The rutile phase exhibits greater hardness compared to the anatase phase, along with reduced wear and friction properties [18,19,20]. By selecting the most optimal modes and combining components such as anatase and rutile, it is possible to improve the mechanical and tribological properties of the PEO coating. Moreover, during the PEO process, under the action of pulse current and local surface heating, it is possible to increase the beta phase on the titanium substrate surface. This leads to a change in the mechanical properties of the surface [21,22,23].

The PEO method provides an opportunity to incorporate bioactive elements such as Ca and P into the coating, which can improve osseointegration [24,25]. It is important to obtain a layer of hydroxyapatite (HA) on the surface, as it is an inorganic component of bone tissue that stimulates the fusion of bone with the implant, which can shorten the treatment time and reduce the risk of rejection of the implant by the body. Another important factor is the formation of micropores during the PEO process, which allows cells to adhere and proliferate [26,27]. Furthermore, the application of PEO is often related to simplicity, cost-effectiveness, and a quick process time. Furthermore, the resultant coatings exhibit much greater thickness than normal coatings, increased corrosion and wear resistance, and more favorable biological characteristics [28].

Recent developments in the field of additive manufacturing have opened up new possibilities for designing the next generation of metallic biomedical implants using three-dimensional porous scaffolds. The design and manufacture of porous titanium scaffolds can be tailored to meet specific needs and specific applications by controlling the architecture, porosity, and pore size of the scaffold [29,30,31]. This may be useful in optimizing scaffold performance in terms of its mechanical properties, its ability to support tissue growth, and its biocompatibility. The potential benefits of this new approach to implant design include improved patient outcomes, reduced surgical time and costs, and increased implant longevity. Despite the progress that has been made in additive manufacturing, more research and development is needed to improve these areas and fully realize the potential of this innovative approach to biomedical implant design [32].

There are numerous studies that investigate coatings formed using the PEO process on the commercial Ti–6Al–4V alloy, incorporating components such as magnesium, zinc, zirconium, and others [33,34,35,36]. However, few existing studies specifically focus on the titanium alloy Ti–6Al–4V produced using the SLM process, particularly in relation to the creation of bone scaffolds with varying structures [3,37].

Wu et al. [38] examined the impact of microstructure on the development of PEO coatings. The coatings formed on a substrate produced using the SLM technique exhibited a reduction in defects and pore size, and also an increase in coating thickness, in comparison to coatings formed on substrates made of the commercial titanium alloy Ti–6Al–4V. Nevertheless, this work did not include any tribological studies of the coatings. Due to the limited research on the tribological and physical–mechanical characteristics of TiO₂ coatings, namely those containing -P and -Ca ions produced using the PEO method on Ti–6Al–4V titanium alloy samples formed using the SLM process, it was determined that flat samples made from Ti–6Al–4V would be created. This study will enhance our comprehension of the mechanisms involved in PEO and contribute to the improvement of implants and bone scaffolds with a porous structure produced using the SLM method.

The main purpose of this research article is to examine the influence of PEO parameters on the structural phase states and mechanical properties of hydroxyapatite coatings deposited on Ti–6Al–4V substrates produced using selective laser melting. Moreover, the role of the PEO-process-induced β-Ti phase in changing the mechanical and tribological characteristics was revealed.

2. Materials and Methods

The 20 × 30 × 2 mm³ substrates were fabricated using selective laser melting (SLM) using a MLab Cusing R additive manufacturing system (Concept Laser, Lichtenfels, Germany) using titanium alloy powder (Ti–6Al–4V) DIN EN ISO 2267 [39] Rematitan^® (Ispringen, Germany). The samples were treated in a vacuum oven at a temperature of 820 °C for a duration of 4 h, followed by cooling within the oven. Then, the samples were ground with SiC sandpaper of up to 100 grit, then cleaned in an ultrasonic bath, and washed in distilled water.

The PEO process was carried out in an aqueous medium with a total volume of 1000 mL containing 25 g/L sodium dibasic dodecahydrate sodium phosphate (Na₂HPO₄ 12H₂O) 25 g/L and 50 g/L calcium acetate (Ca(CH₃COO)₂ 2H₂O) (Krisanalyt, Kazakhstan). During the process, a stable temperature range of 20–25 °C was maintained inside of a titanium bath (Ti–6Al–4V) by circulating water in a jacketed beaker connected to a water-cooling unit. A titanium bath was used as the cathode, and a sample was used as the anode. Figure 1 shows an illustrative representation of a PEO process. During a PEO process on a titanium substrate, an amorphous dielectric film initially forms on the surface. As the voltage increases, this film transitions into a crystalline TiO₂ layer. However, further increasing the voltage leads to the breakdown of the TiO₂ dielectric. Additionally, Ca²⁺, ${\mathrm{H}\mathrm{P}\mathrm{O}}_4^{3-}$, or ${\mathrm{P}\mathrm{O}}_4^{3-}$ and OH^– ions from the electrolyte solution are incorporated into the coatings. More complicated ceramics, such as CaTiO₃, α-TCP, and Ca-deficient HA (Ca_10−x(HPO₄)x (PO4)_6−x(OH)_2−x, 0 ≤ x ≤ 1, e.g., d-HA), are generated [40]. For the application of coatings, an impulsive mode switching power supply, model “PV-500V/20kW” (Promgeotechnology Ltd., Tomsk, Russia), was used. The specific process parameters are detailed in Table 1, below. During the PEO process, the voltage and current were increased gradually, and were measured at the output of the power supply. The voltage from 300 V to a limited value and maximum current density values are given in Table 1.

The surface morphology and elemental composition were investigated with a JSM-6390LV scanning electron microscope (SEM) equipped with an INCA Energy Penta FET X3 energy dispersive microanalysis system (JEOL Ltd., Tokyo, Japan). An X-ray diffraction examination of the coatings was performed using a diffractometer with a wavelength of Cu Kα 1.54056 Å (PANalytical X’Pert PRO, Almelo, The Netherlands). Phase identification was accomplished by using the ICSD databases.

The tribological characteristics test was conducted utilizing a “ball–disk” scheme on a TRB³ tribometer (Anton Paar Srl, Buchs, Switzerland) using the following test parameters: radius: 2.50 mm; velocity: 5.00 cm/s; vertical load: 5.00 N. The track was characterized using “Surtronic S128” profilometry (Taylor Hobson Ltd., Leicester, UK). The microhardness and elastic modulus measurements were performed on the device “Fischerscope HM2000 S” (Helmut Fischer GmbH, Sindelfingen, Germany) with the following parameters: Load increase: F = 2000.000 mN/10 s; creep at final load: C = 5.0 s.

3. Results and Discussion

3.1. SEM–EDS Analysis of the Coatings

The surface morphology and cross-sectional SEM images of the oxidized layers of the titanium samples obtained using the PEO method at different voltages (200, 250, and 300 V) are shown in Figure 2.

Figure 3 shows the scanning electron microscope (SEM) images of the microstructure of the Ti–6Al–4V alloy produced using the selective laser melting (SLM) technique, following heat treatment. The images were captured at various magnifications. The Ti–6Al–4V alloy’s microstructure, upon heat treatment, exhibits a combination of α + β phases characterized by lamellar structures. The samples were etched with Kroll’s reagent.

It was found that the surface morphology changes significantly with an increase in the applied voltage. At 200 V, the PEO coating appeared thin and exhibited a furrowed morphology (Figure 2a). As the voltage increased to 250 V, the pore size grew, and the pores became more homogeneously distributed across the surface (Figure 2b).

This study revealed that increasing the voltage leads to an exponential increase in coating thickness. Specifically, the coating thickness was approximately 1 µm at 200 V, around 3–3.5 µm at 250 V, and about 35–37 µm at 300 V. At 300 V, the PEO coating layer exhibited cracking, and the surface became notably rough, with craters ranging in size from 10 to 30 µm. These cracks are likely due to the rapid solidification of the molten materials during the PEO process. Moreover, the morphological changes observed at 300 V can be attributed to a rise in current density throughout the PEO process. This increased current density enhances the impact on the treated material’s surface, accelerating the coating process and increasing the oxide coating thickness. A thicker oxide layer demands more energy for breakdown, leading to the formation of large spark discharges and the creation of significant pores [41].

The thickness of the oxide coatings increased from 1 μm to 3.2 μm at voltages from 200 to 250 V (Figure 2d,e). At an applied voltage of 300 V, the thickness of the PEO coating was about 35 µm. (Figure 2f). The elemental analysis (mapping) of the cross section of the samples indicates that TiO₂ layers formed at a voltage of 200–250 V and were enriched with Ca and P (Figure 4a,b). The elemental mapping samples obtained at a voltage of 300 V show that layers containing Ca and P were located on the outer surface (Figure 4c). Hence, the examination of the surface morphology reveals the initiation of a notable titanium oxide layer at 250 V. Elevating the voltage to 300 V resulted in the emergence of a layer with an increased content of Ca and P, measuring roughly 35 μm in thickness, atop the pre-existing titanium oxide layer.

According to the energy-dispersive spectroscopy analysis (EDS), the weight ratio of Ca/P increased with the voltage range from 200 V to 250 V and then decreased at 300 V. The observed values ranged from 2.24 to 2.99 and 1.96. The values of 2.24 and 1.96 were close to the stoichiometric Ca/P hydroxyapatite ratio of 2.15, which corresponds to a molar ratio of 1.67 [42]. The Ca and P content increases with increasing voltage, and the Ti content decreases, respectively (Table 2). At voltages of 200–250, the Ca and P content in the coating compared to the sample obtained at 300 V was significantly lower. This indicates that effective formation of hydroxyapatite layer occurs at high voltage treatment.

3.2. The X-ray Diffraction and FTIR Analysis of the Coatings

The X-ray diffraction (XRD) patterns of the PEO coating acquired at different applied voltages are illustrated in Figure 5. The α-Ti and β-Ti phases were identified on the substrate (Ti–6Al–4V) manufactured using the SLM process (Figure 5a). The coating at a voltage 200 V and 250 V was composed of the different phases of titanium (α-Ti and β-Ti) and titanium oxide (TiO₂): anatase (ICSD 98-008-2084) and rutile (ICSD 98-003-3844) (Figure 5b,c). At a voltage of 300 V, Ti and TiO₂ peaks were substantially reduced, and hydroxyapatite became the dominant phase (Figure 5d), which exhibited typical diffraction peaks of (002), (121), (022), (123), and (014). This is well correlated with the established standard ICSD database for hydroxyapatite (HA) (ICSD 98-005-6309).

Furthermore, the XRD analysis shows the percentage of various phases. Increasing the applied voltage leads to an increase in the content of the β-Ti phase. However, at a voltage of 300 V, this increase was not observed, possibly due to the formation of a thick surface layer of hydroxyapatite that prevents the identification of the underlying phases. During the PEO process, there is a local heating of the sample surface, which possibly contributes to the increase in the β-phase of titanium, which may influence the mechanical properties of the coating [23,43,44]. In addition, when the voltage is increased, the sample is heated, and anatase is transformed into a more stable form—rutile. Some studies suggest that rutile is superior to anatase in terms of biocompatibility [45,46].

The FTIR spectra of the specimen oxidized at a voltage of 300 V are given in Figure 6. Most of the bands corresponding to the hydroxyapatite, including bending vibrations of PO₄^3– absorption bands at 567, 580 cm⁻¹, and 964 and 1065 cm⁻¹, can be attributed to the asymmetric stretching mode (ν3) of the P–O bond. The vibrational bands of carbonate groups (1540, 1420, and 881 cm⁻¹) emerge as a result of presence of calcium acetate in the electrolyte. These bands indicate that carbonate groups partially replace the PO₄ and OH groups in the HA coating, leading to the formation of carbonate-substituted HA. The vibrational bands of carbonate groups CO₃^2- occur at about 1420 and 1384 cm⁻¹, and the adsorbed water bands (OH) at about 3420 and 1648 cm⁻¹ [47]. The peaks detected in the regions at 2855 and 2925 cm⁻¹ were attributed to the stretching vibrations of the CH bonds [48], which might be explained by presence of calcium acetate in the electrolyte. These results confirm the formation of the hydroxyapatite phase during the PEO process at 300 V.

3.3. The Tribological and Mechanical Properties of the Coatings

Figure 7 illustrates the results of the tribological investigations conducted on the PEO coating. In the control sample, the coefficient of friction increased immediately after the start of the sliding and varied around approximately 0.4 (Figure 7a), which corresponds with the literature data [37,49]. For samples treated with a voltage of 200 V, the coefficient of friction progressively increased, and, after approximately 5 m, the average was around 0.5, which is a bit more than that of the control sample (Figure 7b). At a voltage of 250 V, the coefficient of friction was similar to that of 200 V (Figure 7c), but a fluctuation was observed during sliding friction, which was largely due to the abrasive impact of hard TiO₂ phases and the presence of a porous structure. The friction coefficient initially began at a high value and quickly decreased to around approximately six at 300 V. Additionally, high fluctuations for coating obtained at 300 V were observed (Figure 7d). This result could be attributed to the formation of a substantial porous layer of hydroxyapatite.

The wear rate was assessed with the profilometer “Surtronic S128” (Table 3), examining the track depth obtained after the tribological properties test, and evaluated according to ASTM G99-17 [50]. The wear resistance is significantly influenced by the density of the coatings, which is associated with an increase in pore size. The wear resistance notably improved in the samples treated with 250 V, likely due to the thicker oxide film compared to those treated with 200 V. However, at 300 V, the wear resistance decreased, apparently due to the formation of a more porous hydroxyapatite layer.

In this way, the coefficient of friction and the wear rate largely depended on the morphology, pores size, and density of the layer of titanium dioxide. Therefore, the sample at 250 V exhibited better wear resistance compared to the other two coated samples.

Figure 8 demonstrates the microhardness and elastic modulus of the samples obtained. The measurements were taken with a Fischer scope HM 2000 S measurement instrument. The results show that the microhardness and elastic modulus have a direct correlation. The coatings’ microhardness and elastic modulus reduce during PEO as a porous structure forms on the surface. Moreover, it was previously observed that the morphology and surface porosity significantly change with rising voltage and increasing content of β-Ti, which has more flexibility than α-Ti [51]. The mechanical characteristics of the modified surface suggest that the obtained coating positively impacts the implants’ ability to maintain high strength properties. This is expected to significantly support the patients’ activity levels [52].

4. Conclusions

This paper investigates the coatings obtained using the PEO process on Ti–6Al–4V substrates fabricated using SLM technology. The PEO process occurs in sodium phosphate dibasic dodecahydrate and calcium acetate containing a base electrolyte at a voltage of 200, 250, and 300 V, respectively. The hydroxyapatite coating was successfully achieved using plasma electrolyte oxidation at a voltage of 300 V, which is confirmed by the results of XRD and FTIR analysis.

The analyses of the surface structure of the coating reveal that there is a rise in the size of pores and the thickness of the coating as the applied voltage increases. Also, the elemental analysis shows an increase in calcium and phosphorus content with increasing applied voltage. At a voltage of 300 V, the calcium to phosphorus ratio (1.96) is close to the stoichiometric ratio of hydroxyapatite. Furthermore, it shows that bioactive components like calcium and phosphorus enrich the coating. The XRD analysis indicates the presence of titanium dioxide phases (anatase and rutile) in the coating, and an increase in β-Ti content with increasing applied voltage is observed.

The PEO coatings show a high value of the coefficient of friction compared to the substrate. At a voltage of 250 V, better wear resistance is observed compared to that of the other modes. Also, a decrease in the elastic modulus and microhardness of the PEO coating is observed. The decrease in the elastic modulus is related to the increase in β-Ti during PEO processing. The metal implant should have mechanical properties close to those of organic human bone, as complications may arise due to the shielding effect; therefore, a reduction in the elastic modulus and microhardness may have a positive effect.

Author Contributions

A.S., writing—original draft and conceptualization; D.D., investigation and methodology; A.T., supervision and validation; D.K., data curation and visualization. All authors have read and agreed to the published version of the manuscript.

Institutional Review Board Statement

Not applicable.

Informed Consent Statement

Not applicable.

Data Availability Statement

The data and results presented in this study are available in the article.

Conflicts of Interest

Amanzhol Turlybekuly is employed at the Institute of Materials Engineering LLC. The funders had no role in the design of the study; in the collection, analyses, or interpretation of data; in the writing of the manuscript; or in the decision to publish the results.

Footnotes

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Figures and Tables

Figure 1. Schematic diagram of the PEO process.

Figure 2. Surface SEM micrographs of PEO samples formed at (a) 200 V, (b) 250 V, and (c) 300 V; and coatings’ cross sections of (d) 200 V, (e) 250 V, and (f) 300 V.

Figure 3. Microstructure of SLM Ti–6Al–4V alloy (a) ×500 and (b) ×2000.

Figure 4. EDS elemental mappings of the PEO samples’ cross section: (a) 200 V, (b) 250 V, and (c) 300 V.

Figure 5. XRD patterns of the PEO samples: (a) control sample, (b) 200 V, (c) 250 V, and (d) 300 V ([Image omitted. Please see PDF.] α Ti (ICSD 98-007-6265), [Image omitted. Please see PDF.] β Ti (ICSD 98-004-1503), [Image omitted. Please see PDF.] HA (ICSD 98-005-6309), [Image omitted. Please see PDF.] Anatase (ICSD 98-008-2084), and [Image omitted. Please see PDF.] Rutile (ICSD 98-003-3844)).

Figure 6. FTIR spectra of the sample obtained at a voltage of 300 V.

Figure 7. Coefficient of friction of the PEO samples: (a) control sample, (b) 200 V, (c) 250 V, and (d) 300 V.

Figure 8. Microhardness and elastic modulus of PEO coatings.

Table 1

PEO process parameters.

Pulse Voltage	Current Density	Duty Cycle	Pulse Frequency	Processing Time
200 V	0.057 A/cm2	10%	50 Hz	5 min
250 V	0.11 A/cm2	10%	50 Hz	5 min
300 V	0.2 A/cm2	10%	50 Hz	5 min

Table 2

EDS analysis of the surfaces of the PEO samples.

Element (wt.%)	200 V	250 V	300 V
O	44.15 ± 3.31	41.88 ± 5.72	39.63 ± 9.15
P	1.81 ± 0.19	2.15 ± 0.27	20.03 ± 2.00
Ca	4.06 ± 0.69	6.44 ± 0.92	39.34 ± 7.30
Ti	45.24 ± 3.16	45.36 ± 5.22	1.01 ± 0.13
Al	2.75 ± 0.11	2.24 ± 0.18	-
V	1.99 ± 0.24	1.93 ± 0.33	-
Ca/P	2.24	2.99	1.96

Table 3

Coefficient of friction and wear rate of the PEO samples.

Sample	Coefficient of Friction	Wear Rate (mm³/N/m)
Control sample (Ti–6Al–4V)	0.396	3.087 × 10−4
200 V	0.540	3.447 × 10−4
250 V	0.546	2.765 × 10−4
300 V	0.600	3.270 × 10−4

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