The recent focus of research is on biocompatible materials, which are artificial products that have been imposed by the needs of people, affected by disease or accident, to improve health, leading to increased life expectancy [1]. The main issue is the way the body accepts these materials, or biocompatibility [2]. From here, the main disadvantages of implants are determined by the nature, mode of synthesis and different physical and chemical properties of the materials used in relation to living tissues, such as: the ability to change its structure and properties according to the demands it supports (mechanical loading for bone tissue or blood flow for blood vessels) or self-healing ability [3,4]. Biomedical engineering, including the development of artificial implants such as knee and hip stents, has become an important international activity with a significant social impact [5]. Research contributes significantly to the understanding of the effects of mechanical forces on human recovery and locomotion and joint functionality. The boundless utilization of titanium (Ti) compounds in the area of implantology is principally because of the expanded consumption, obstruction and more elevated levels of biocompatibility. However, not all Ti combinations meet the prerequisites for biomedical applications. This has prompted the improvement of another collection of Ti amalgams with biocompatible components such as Mo, Zr, Ta, Nb, and so on, which as of late, have been introduced to the market to overcome the toxicity associated with Ti-V/Al-based alloys currently in use. In addition, because the surface of the biomaterial plays a key role in the cellular response, the improvement of biological and tribological properties by surface modification is often necessary [6,7,8].

Hydroxyapatite (Ca₁₀(PO₄)₆ OH)₂) is the major inorganic constituent in bone mass, weighing approximately 69% [9]. The most important property of hydroxyapatite as a biomaterial is its excellent biocompatibility, which is manifested in making direct connections with living bone tissue [10]. It is known that biological hydroxyapatite is a calcium phosphate that is found in the mineral structure of bone and tooth—in dentin and tooth enamel, but also in the case of pathological calcifications such as kidney stones, stones of the salivary gland, dental stones, etc. [11]. Hydroxyapatite can crystallize as a fine salt depending on the Ca/P ratio, the formation temperature, the presence of water or impurities and, depending on the preparation medium, in a humid environment at relatively low temperatures [12]. Hydroxyapatite is not only bioactive, but also osteoconductive and non-toxic [13]. The bonds they make with the bone are of a physico-chemical nature; the bone cells interact with the hydroxyapatite forming ionic, hydrogen and van der Waals bonds [13].

The notion of osteoconductivity has different meanings, depending on the field in which it is used [14]. Clinically, this means bone growth from the host bone tissue to the implant. Due to this meaning, any material (not only calcium phosphates, but also polymers) can be osteoconductive, based on the ability to regenerate the bone itself [15]. Current studies aim to obtain new materials for implants, as well as their biological and mechanical interfaces. The modification of surfaces, through chemical processes or the attachment of biomolecules, represents an important option for improving biocompatibility [16]. The properties of hydroxyapatite, including bioactivity, biocompatibility, solubility, mechanical and adsorption properties, can be adapted to a wide range of applications by controlling particle composition, size and morphology [17]. Hydroxyapatite in various forms such as powder, porous blocks or pearls can be used to fill bone defects and free spaces in the bone. These occur when portions of the bone have been removed due to disease (bone cancer) or when bone elongations are required (in dental applications). For medical applications, hydroxyapatite deposits are also obtained on solid supports (metallic or polymeric).

The surface modification techniques are listed and divided according to the governing processes and their purposes [18]. The main purpose of changing the metal surface is to improve the compatibility with hard tissue or to accelerate bone formation [19]. The development of biomimetic ceramic materials is of great interest in tissue and prosthetic engineering [20,21,22]. The most interesting property of hydroxyapatite is its ability to interact with living bone tissue, forming strong bonds with the bone. It is frequently used for orthopedic, dental and maxillofacial applications, either as a coating material for metal implants or as a bone-filling material [23,24]. Other studies report that various hip prostheses and component parts, based on metals or metal alloys, especially titanium alloys coated with a layer of hydroxyapatite, had the role of facilitating bone development on the implant. There are many ways to deposit hydroxyapatite on titanium alloys. In the field of coating techniques, several methods of deposition are mentioned, such as atmospheric plasma spraying, physical and chemical vapour deposition (PVD and CVD processes) [8]. In addition, in order to increase the bonding strength between coating and alloy substrate, thermal spraying technology is considered to be an effective coating preparation method, such as hot dip method [25,26], chemical vapor deposition method [27,28], slurry method [29,30], etc.

2. Microstructural Analysis

Figure 1 illustrates the microstructural analysis of the substrate alloys. Alloys present coarse grained, lamellar microstructures with large β-grains and a lamellar matrix of alternating α and β. This type of lamellar structure is specific to titanium alloys depending on the percentage of beta stabilizing elements. All alloys contain a large amount of beta stabilizing elements (Mo, Ta) that lead to this type of structure.

Figure 1. Microscopy images of substrate materials: (a) Ti15Mo7Zr5Ta; (b) Ti15Mo7Zr10Ta; (c) Ti15Mo7Zr15Ta.

Morphological aspects of coated samples with hydroxyapatite coating are presented in Figure 2.

Figure 2. Structural aspects of the surface of coated samples at 500×: (a) S1-HA; (b) S2-HA; (c) S3-HA; and 1000×: (d) S1-HA; (e) S2-HA; (f) S3-HA.

The images confirmed that the hydroxyapatite coatings on the titanium surface were obtained, and the coatings were uniform.

After modifying the surface, the latest research suggests that the application of various biomimetic coatings can be beneficial for implant therapy success [31,32,33,34].

Figure 3 shows the hydroxyapatite layer deposited on the surface of titanium is uniform, with a thickness of the order of microns (Table 1) and contains crystals of nanometric dimensions as shown by natural bone. As shown in Table 1, the thickness of the samples decreased from 35 to 29 µm. Figure 4 shows a representative image of an HA layer on titanium alloys.

Figure 3. EDS elemental mapping of the metallic surface: (a) S1-HA; (b) S2-HA; (c) S3-HA.

Figure 4. Representative aspects layer of HA on titanium alloys.

Table 1. The thickness of the deposited layers.

Sample	Layer Thickness (µm)
S1-HA	35 ± 3
S2-HA	29 ± 4
S3-HA	31 ± 2

Hydroxyapatite coatings are often used for titanium implants in order to modify the surface properties. Most applications of HA coatings are for endosseous and subperiosteal dental implants and for orthopedic devices.

By coating the implant with a layer of hydroxyapatite, the implant benefits from biocompatibility, the ability to form chemical bonds with living bone and the mechanical properties of the titanium substrate. Due to the osteophilic surface of hydroxyapatite, the mechanical load acting on the implant will be transferred to the skeletal bone helping to combat bone atrophy [35,36,37].

The bone filling will form a skeleton and will facilitate the rapid filling of the pores by growing natural bone tissue. Hydroxyapatite as a filler is an alternative to bone grafts, becoming part of the bone structure and reducing the time required to heal diseased tissue.

Research in bone engineering shows that the structural properties of hydroxyapatite give it a better degree of resorbability and better osteoconductivity than dense hydroxyapatite. In addition, researchers recommend it as a good bone substitute for orthopedic implant surgery [38,39].

Figure 5 presents the X-Ray Diffraction (XRD) patterns of the coated alloys. It can be observed that the highest peaks observed in all the samples belonged to substrate metals S1, S2 and S3. This is because XRD beams can penetrate the thin coatings of hydroxyapatite and are diffracted from the underlying metals.

Figure 5. X-ray diffraction (XRD) graph of the coated alloys.

Figure 5 shows the result of X-ray diffraction obtained on the substrate samples of TMZT and were determined by some characteristic peaks of the substrate, which show the phases α and β, at an angle of 2θ: 35.40; 38.50; 39.45; 40.45; 53.30; 57.00; 63.55; 71.00; 74.90; 76.80; 78.10 with d (nm), respectively: 0.2536; 0.2338; 0.2284; 0.2230; 0.1719; 0.1614; 0.1464; 0.1328; 0.1268; 0.1241 and 0.1244. The β phase is clearly characterized by (110) and (200) and by the characteristic distortion reflection of the above-mentioned structure 0.2284 and 0.1614 nm respectively according to [40,41,42].

HA has a hexagonal crystal structure with the main one’s diffraction peaks at 2θ = 15.9; 29.0; 31.8; 32.2; 32.9; 34.0; 39.8; 46.7; 49.5; 50.5 and 83.1 corresponding to the crystal orientation planes (002), (210), (211), (112), (300), (202), (310), (222), (213), (321) and (004), as shown in Figure 5. The figure shows that the XRD model of the HA target is in accordance with the international standard, JCPDS 9-0432 [43].

Alloys show a crystalline structure HA with calcium phosphate (TCP) and Ca₃(PO₄)₃ (PO4)₂ that is formed during the sintering process.

3. Micro-Indentation Analysis

Table 2 presents the micro-indentation analysis of coated samples’ results. Results are provided by the indentation test with the help of the VIEWER program which records them with the UMT 2 software (CETR, Campbell, CA, USA). The Young modulus obtained shows values between 55.35 ± 0.3 and 56.45.27 ± 0.3. This type of hydroxyapatite coating reduces the Young modulus of the Ti alloy (103–120 GPa) by 50% and makes them approach close to the cortical bone value (10–32 GPa). Figure 6 shows the response of the alloys during an indentation test, the force–displacement curve.

Figure 6. Micro-indentation graphic: (a) S1-HA, (b) S2-HA, (c) S3-HA.

Table 2. Micro-indentation results *.

Sample	Loading Deformation (N)	Release Deformation (μm)	Young Modulus (GPa)	Stiffness (N/μm)	Specimen Poisson Ration
S1-HA	13.35 ± 0.4	12.75 ± 0.3	55.35 ± 0.3	6.35 ± 0.1	0.27
S2-HA	13.25 ± 0.5	12.10 ± 0.1	56.25 ± 0.2	6.25 ± 0.2	0.27
S3-HA	13.75 ± 0.4	12.35 ± 0.2	56.45 ± 0.3	5.75 ± 0.2	0.27

The more the coefficient of elasticity of the implant resembles the biological contingent tissue, the smaller the movement relative to the tissue-implant interface. Cortical bone is at least five times more flexible than titanium. As the size of the tension increases, the difference between the relative stiffness of the bone and the titanium increases. As the size of the tension decreases, the difference in relative stiffness between bone and titanium becomes much smaller. In this aspect, titanium alloys for orthopedic applications must have a modulus of elasticity as small as possible for long-term success [44].

4. Contact Angle Analysis