**1. Introduction**

Recently, there has been a high demand for implants for bone dysfunction caused by damages, diseases, and fractured bones by aging or accident, so numerous studies related to biomedical implant materials which can function on the damages or fractured bones by aging or accident have been actively conducted [1–3]. According to the distinct required properties for biomedical implants, limited metals or ceramics were selected. The biomedical implants must not react with a tissue of the human body because it causes biological instability due to corrosion or degradation of the implants [4]. Therefore, not only high strength and low stiffness with lightweight but also corrosion resistance in vivo should be satisfied to apply in biomedical applications. For metals, titanium (Ti) and its alloys, stainless steel, and cobalt–chromium alloys have been extensively reported [5,6].

Among the metals, Ti and its alloys are widely used as implant materials owing to excellent biocompatibility with tissues in the human body and high corrosion resistance and mechanical properties [4,7–9]. The Ti-6 aluminum-4 Vanadium (Ti6Al4V) alloy, the most extensively used Ti alloy, has been used successfully for dental and orthopedic implants, due to its high corrosion resistance and osseointegration by a thin oxide layer formed on the surface in a very short time [10]. The oxide layer exists on the surface of thickness in ~2 nm, which prevents elution of the metal ions and provides the corrosion resistance to withstand the intra-vital corrosion environment [11].

Hence, most biomaterials have used hydroxyapatite (Ca10(PO4)6(OH)2) or calcium pyrophosphate (Ca2P2O·2H2O) coated with Ti6Al4V alloys using several coating methods, such as plasma spraying, thermal spraying, and the hot dipping method to increase the osseointegration with the bones [12–15]. In order to give high osteoconductive property to Ti6Al4V alloy, calcium phosphate-based ceramics such as hydroxyapatite are also coated by plasma spraying [16].

Especially hydroxyapatite has high biocompatibility, bioactivity, osteoconductive property, and osteoinduction, since a lot of researchers have released results of biomaterials containing

hydroxyapatite [13,15,17–22]. Meanwhile, biomaterials containing calcium phosphate (Ca3(PO4)2) from biological wastes such as pig bones, fish bones, and horse bones have been used as raw materials for mass production of natural hydroxyapatite at low cost, according to recent reports [23–25]. Production of the hydroxyapatite from abundant natural sources rather than synthetic sources is much more economical and eco-friendly. In particular, horse bones are free from foot-and-mouth disease and have a positive e ffect on the extraction of natural calcium phosphate due to the high supply rate and low price due to the increase in the number of slaughter horses.

However, in the case of metal coated with the synthetically-produced hydroxyapatite, detachment of hydroxyapatite from a metallic substrate occurs, which is a fatal disadvantage [26–28]. For instance, biomedical implants coated with calcium phosphates from the natural bones using the spray method may cause necrosis of cells because, during long-term use, the di fferences in thermal expansion coe fficients and mechanical properties between the calcium phosphate and the metallic materials may cause peeling to occur. Therefore, a study on the Ti6Al4V matrix biocomposites containing hydroxyapatite produced by powder metallurgy has suggested that peeling problems can be avoided and presented high strength and low elastic modulus compared to microcrystalline Ti [29].

The aim of the present study is to sugges<sup>t</sup> the fabrication process of the Ti6Al4V/equine bone (EB) composite using powder metallurgy and overcome the peeling of the hydroxyapatite from the Ti6Al4V. Hence, Ti6Al4V/EB composite powders are produced using a low-energy ball-milling and the composite powders rapidly sintered by spark plasma sintering (SPS) have inhibited grain growth and decomposition of hydroxyapatite. The effect of the EB on the microstructure and mechanical properties has been evaluated as a function of EB weight fraction; a quantitative and qualitative analysis of the composites is also done.

#### **2. Experimental Procedure**

#### *2.1. Materials and Methods*

Ti6Al4V-EB composites were fabricated by a powder metallurgy route. Commercial powder of pure Ti6Al4V (AP&C, Quebec, Canada) with a mean particle size of 35 μm and a composition of 6.2% Al, 4.1% V, and 0.1% O in a Ti matrix (weight%), and EB powders (Jeju, Korea) with an average powder size of around 1 μm and a composition of 39.4% O, 41.4% Ca, 16.7% P, and 0.56% Mg were used. The Ti6Al4V-EB composite powders of 50 g were ball-milled and utilized by a planetary mill (Pulverisette 4, Fritsch, Idar-Oberstein, Germany) at a rotational speed of 200 RPM for 12 h, followed by a 40 min pause after every 20 min of milling to avoid overheating under air atmosphere. A stainless steel bowl (500 ml) was charged with the Ti6Al4V and EB powders and stainless steel balls without any process control agent. The diameter of the stainless steel balls was 15 mm, and the ball-to-powder weight ratio employed was 5:1. Owing to the shearing mode, EB powders were gradually embedded into the Ti6Al4V powder. The Ti6Al4V and ball-milled Ti6Al4V-EB composite powders were rapidly consolidated using SPS. The composite powders were poured into a 30 mm graphite die with 10 g, and the heating rate was 100 ◦C/min up to 1000 ◦C/min, and maintained for 15 min at an applied pressure of 50 MPa under a high vacuum atmosphere. After SPS, the pressure was removed, and the sintered specimen was cooled down to room temperature. SPSed Ti6Al4V and Ti6Al4V-EB composites have dimensions of 30 mm in diameter and 10 mm in height.
