New Developments of Ti-Based Alloys for Biomedical Applications

2.1.1. Mechanical Properties

2.1.2. Biocompatibility

Ideally, biomedical implants are required to be highly innocuous without any inflammatory or allergic reactions in human body. Whether implant surgery is successful mainly depends on the reaction of human body to the implant, which evaluates the biocompatibility of a biomaterial [65]. The reactions deduced from the material and its deterioration in the body environment are the two main factors affecting biocompatibility. Understanding the direct effects of an individual alloying element is of prime importance since it can dissolve in the human body due to tough wear and corrosion, causing local and systemic toxicity, inflammation and immune response. When materials are implanted into human body, a series of reactions happen, and this describes the acceptability of these materials by our system. At present, issues concerning biocompatibility are thrombosis, including blood coagulation and adhesion of blood platelets to biomaterial surface, and the fibrous tissue encapsulation of biomaterials that are implanted in soft tissues. To examine this above-mentioned performance, various tests are introduced into the research extracting from the material, offering screens for genotoxicity, carcinogenicity, reproductive toxicity, cytotoxicity, irritation, sensitivity and sterilization agent residues [66]. Biological safety of metals including cytotoxicity of pure metals and the relationship between polarization resistance and biocompatibility of pure metals, Co-Cr alloy and stainless steels is shown in Figure 3.

It is concluded that Ti, B, Mg, Si, P, Ca, Sr, Zr, Nb, Mo, Pd, In, Sn, Ta, Pt, and Au are biocompatible elements [67], while harmful elements include Be, Al, V, Cr, Mn, Fe, Co, Ni, Cu, Zn, and Ag [67]. The toxicity of V has been widely reported. Al is a questionable element on certain disease. The cytotoxicity of pure metals and the relationship between biocompatibility and polarization resistance of typical pure metals and surgical implant materials have been reported by Steinemann [18]. The high cytotoxicity of V and tissue response of capsule (scar tissue) type due to Al has been demonstrated, while Ti, Nb, Ta and Zr exhibit excellent biocompatibility and belong to the loose connective vital group in the type of tissue reaction. Kawahara has also reported that metallic V and Fe are cytotoxic elements, while Ti, Nb, Ta, Zr and Sn are low cytotoxic elements [18]. These trends reported by Steimemann and Kawahara are shown in Figure 3. Ninomi et al. [61] have shown that the cell viability of Ti-29Nb-13Ta-4.6Zr is much superior to the Ti-6Al-4V alloy. The non-toxic elements like Nb, Ta, Zr, Mo and Sn were selected for designing new β-type Ti-based alloys with lower elastic modulus, greater strength and corrosion resistance.

Human body is a complicated electrochemical system composed of an aggressive corrosion environment for implants. Because body fluids contain different kinds of corrosive substances and implants are exposed to them. Corrosion resistance is one of the principal properties an implant should possess. The success of an implant depends on the serious inspection of this property which is directly related to its capacity in reacting with the severe body environment. Bad corrosion resistance in the body fluids will cause the release of incompatible metal ions from the implant, which is a main factor leading to allergic and toxic reactions [68]. The consequences of corrosion are the disintegration of the implant and the harmful effect on the surrounding tissues and organs is produced. Surface roughness will increase the functioning area of an implant and raise total amount of corrosion. Therefore, surface finishing is an important element in improving corrosion resistance and consequently biocompatibility of the implant [69]. Besides, wear resistance mainly determines the service period of an implant, for low wear resistance results in implant loosening and wear debris and causes reactions in the tissue where they are deposited [70]. Ti-based alloys with a high coefficient of friction can lead to formation of wear debris and thus result in inflammatory reaction and loosening of the implant [71]. In summary, development of implants with high corrosion and wear resistance is of great importance for the longevity of a biomaterial in our human system. The inability of an implant surface to integrate with the adjacent bone and other tissues due to micromotions can also result in implant loosening [72]. A fibrous tissue is formed between bone and implant, if the implant is not well integrated with the bone [72]. Hence, materials with an appropriate surface are highly essential for the implant to integrate well with the adjacent bone. Surface chemistry, surface roughness and surface morphology all act a critical role in good osseointegration.

Long-term performance of surgical implants is often restricted by their surface properties. Tribological property of titanium and its alloys can be enhanced to a large extent by suitable surface coatings. Surface modification techniques such as physical deposition methods like ion implantation and plasma spray coating, and thermo chemical surface treatments such as nitriding, carburization and boriding have been used to improve the surface hardness of Ti-based alloys [61]. TiN coated hip and knee implants have been found to possess increased wear resistance and good compatibility [73]. However, various coating techniques are under investigation to achieve good adhesion and other required properties [74].

Waviness and porosity of the implant also play a vital role in bone integration. Bone ingrowth into porous surface can cause strong interlocking of surrounding bone tissue with the implant, resulting in improved biomechanical compatibility and high resistance to fatigue loading [42,75]. The size of the pores should be in the range of 100–200 μm for better osseointegration [61]. In addition to porous coatings, development of porous biomaterials to enhance long-term fixation and bone growth have also been tried with great interest. The porous biomaterial is expected to lead to strong interface between the bone and the implant and also the modulus of such porous biomaterial is very low, and thus these materials are expected to overcome the stress shielding effect and loosening of the implants.

A nanosurface seems to be advantageous from the biocompatibility and biomechanical compatibility points of view [61]. Nanocrystalline titanium surface enhances cell growth and exhibits excellent wear resistance due to high hardness and strength. The grain size of metal implant also affects osteoblast adhesion. In vitro studies using ultrafine-grained CP Ti (grade 2) and Ti-6Al-4V alloy exhibited increased cell adhesion when compared to conventional coarse-grained (CG) materials. This increase in cell adhesion is attributed to the increase in surface energy at the grain boundaries.

Biomedical implants require a strength level greater than that of bone and an elastic modulus close to that of human bone. Therefore, biomedical Ti-based alloys must exhibit a low elastic modulus combined with enhanced strength, good fatigue resistance and good workability required in hard tissue replacement. The above discussions lead to a strong belief that the new β-type Ti-based alloys are more promising from the wear, corrosion and biocompatibility aspects for biomaterial applications.

2.2.1. Practice and Current Situation of Ti-Based Alloys in Biomedical Application

Different types of fracture repair mechanisms are known in medical practice. The surgical treatments of bone osteosynthesis are divided into external fracture fixation and internal fracture fixation. With external fracture fixation, the bone fragments are held in alignment by pins placed through the skin onto the skeleton, structurally supported by external bars. With internal fracture fixation, the bone fragments are held by wires, screws, plates, and/or intramedullary devices [79].

Surgical wires are used to reattach large fragments of bone provisionally or permanently to guide large screws during insertion. Screws are the most widely used devices for fixation of bone fragments. The bone immediately adjacent to the screw often undergoes necrosis initially, but if the screw is firmly fixed, permanent secure fixation may be achieved [80]. This is particularly true for Ti-based alloy screws or screws with a roughened thread surface, with which bone ingrowth results in an increase in removal torque [80]. Plates are intended to facilitate fixation of bone fragments. They range from the very rigid to the relatively flexible. The effect of the material on the rigidity of the plate is defined by the elastic modulus of the material for bending, and by the shear modulus for twisting [81]. Thus, given the same dimensions, a Ti-based alloy plate will be less rigid than a stainless steel one, since the elastic modulus of each alloy is 110 and 200 GPa, respectively. Intramedullary devices (nails or rods) are used as internal struts to stabilize long bone fractures. Nails are better positioned to resist multidirectional bending than a plate compared to plates, since they are located in the center of the bone. However, their torsional resistance is less than that of the plate [81].

The design of an implant for joint replacement should be based on the kinematics and dynamic load transfer characteristic of the joint. Overloading the implant-bone interface or shielding it from load transfer may result in bone resorption and subsequent loosening of the implant [82].

The prosthesis for total knee replacement (TKR) consists of femoral, tibial, and/or patellar components, Figure 4a. Compared to the hip joint, the knee joint has a more complicated geometry and movement mechanics, and it is not intrinsically stable. The eccentric movement of the knee helps distribute the load throughout the entire joint surface [83]. TKRs can be implanted with or without cement, the latter relying on porous coating for fixation. The femoral components are typically made of Co-Cr alloy and the monolithic tibial components are made of ultra-high molecular weight polyethylene (UHMWPE). In modular components, the tibial polyethylene component assembles onto a Ti-based alloy tibial tray. The patellar component is made of UHMWPE and a Ti-based alloy back is added to components designed for uncemented use.

The endoprosthesis for total hip replacement (THR) consists of a femoral component and an acetabular component, Figure 4b. The femoral stem is made of Ti-based alloy or Co-Cr alloy. The femoral head is made of Co-Cr alloy, aluminum, or zirconium. Although Ti-based alloy heads function well under clean articulating conditions, they have fallen into disuse because of their low wear resistance to bone or cement particles. The acetabular component is generally made of UHMWPE.

For maxillofacial osteosynthesis in the cranio-facial and mandibular areas, titanium plate and screw systems are preferred. In order to make them pliable, many of the plates are made from CP Ti sheet that is in the soft-recrystallized condition. The corresponding screws are either made from CP Ti or alloy and can be as small as 1 mm in diameter [84].

Traditionally, researchers have used already available materials that had been developed for aerospace or automotive applications, instead of developing new materials tailored specifically for biomedical needs. A typical example is THR, in which a dense metal is used that has a significantly higher density, stiffness and strength than natural bone which is a porous material. The typical lifetime of a THR is 7–12 years, and this lifetime has remained almost constant over the past 50 years, even though significant research and development have gone towards understanding the problem.

Presently, the materials used for these applications are 316L stainless steel, Co-Cr alloys, and Ti-based alloys. Unfortunately, these materials have exhibited tendencies to fail after long-term use due to various reasons depicted in Figure 5. Therefore, there is an increasing demand for improved implants which can perform for a longer lifetime in vivo.

The scenario has changed due to the advancements in medical technology. In addition, the prognosis is better for those who are physically traumatized due to sports or incorrect or over exertive exercise habits or due to road traffic and other accidents. Thus, the implants are now expected to serve for much longer period. The development of appropriate material with high longevity and excellent biocompatibility is highly essential.

2.2.2. Market and Application Prospects

Nowadays, THR and TKR surgeries are being carried out with a higher rate on younger and older patients. The revision surgeries of hip and knee implants have also increased. The data collected on total joint replacements surgery estimates that by the end of 2030, the number of THRs will rise by 174% (572,000 procedures) and the number of TKRs is projected to grow by 673% from the present rate (3.48 million procedures) [85]. The total number of hip revision surgery is expected to increase by 137% and knee revision surgery by 607% between 2005 and 2030 [85]. It is projected that approximately 272,000 THRs will be performed annually by 2030. Additionally, approximately 12.8% of the 152,000 THRs performed in 2000 involve revisions of previous hip replacements. Therefore, there is a tremendous demand for the new long-lasting implants.

There are several reasons motivating improvements in joint replacements. Firstly, continual aging of population has brought an ever-increasing need for materials specifically for human body. It has been estimated 90% of population over the age of 40 suffers from degenerative diseases and the aged people population has increased tremendously recently. The United States Census estimates that the total number of people of age 65 and above has increased from 4.9 to 39.7 million between 2002 and 2010; Second, the age range has been broadened to include older patients who have greater incidence of co-morbidities over the last decade; Finally, THRs are now routinely performed on younger patients whose implants would be exposed to greater mechanical stresses due to the more active lifestyle.

An acceptable reason for the increasing number of revision surgeries is due to the higher life expectancy of the implant vs. the ever-increasing life expectancy of the patient. Consistently, over 30% of those requiring THRs have been below the age of 65 and even those over the age of 65 now have a life expectancy of 17.9 years. With normal implant longevity of 12–15 years, the majority of those that receive hip implants at the age of 65 will require at least one revision surgery. So the fact that such a high percentage of joint replacements performed every year are revision surgeries, although troubling, is not surprising.

Human joints suffer from degenerative diseases such as osteoporosis (weakening of the bones), osteoarthritis (inflammation in the bone joints) and trauma. The degenerative diseases lead to degradation of bone due to excessive loading or absence of normal biological self-healing process. Musculoskeletal disorders are most widespread human health problem which costs around 254 billion dollars to the society [86]. Artificial biomaterials are the solutions for these problems, as surgical implantation of the biomaterials of appropriate shapes help in restoring the function of the otherwise functionally compromised structures. These revision surgeries are very expensive and also their success rate is rather small. Moreover, there is still a lack of bone replacement material that is appropriate for restoring lost structure and function, particularly for load-bearing applications. Thus, a very high boom in implant manufacturing is expected in coming years. Ever-increasing demand for implants makes it imperative that development efforts on biomaterials have been accelerated. Among several materials that are currently in use as biomaterials, Ti-based alloys are fast emerging as the first choice for biomedical applications.

Ti-based alloys exhibit low elastic modulus favorable for homogeneous stress transfer between implant and bone [87]. Presently, Ti-based alloys composed of biocompatible alloying elements are being developed mainly for biomedical implants to be used as implant devices replacing failed hard tissues and dental products, for example, artificial hip joints, artificial knee joints, bone plates, dental implants, crowns, dentures, artificial tooth roots, and screws, etc. Ti-based alloys are also expected to be used to fix soft tissue such as blood vessels [88].

5.3.1. Loose Sintering of Powder

The simplest fabrication technique for making porous titanium and its alloys is based on the partial sintering of metal powders. Compared with other technologies, this process is a simple, cost-effective and most common powder metallurgical route. Except die wall lubrication if necessary, the process does not require a polymeric binder and hence minimize possible contamination associated with binder removal [51].

Oh et al. [51] sintered spherical unalloyed titanium powders with and without applied pressure by controlling sintering conditions and titanium powder sizes in a vacuum of 1 × 10⁻³ Pa, as shown in Table 12. Neck formation between titanium powder particles is observed in the porous sample 13. Only contacts between titanium powder particles are observed in the sintered sample 14. The porous sample 13 with about 30 vol% porosity shows elastic modulus of 25 GPa, bending strength of 115 MPa and compressive yield strength of 61 MPa. Porous compacts with porosity greater than 40 vol% cannot be fabricated when their particle sizes are in the range of 65–374 μm for the strength becomes almost 0 MPa. Porous titanium and its alloys fabricated by partial sintering of metal powders can achieve a low elastic modulus, but the compressive strength is low compared with human cortical bone [51].

For loose sintered powders, the pore size, volume fraction, shape and distribution throughout the sample thickness and the inter particle neck size have a major impact on the mechanical properties of the resulting material [51]. The porosity decreases with decreasing initial powder size and with increasing sintering pressure. In contrast, sintering temperature affects very slightly the densification of titanium compacts, as can be seen in Table 12.

The porosity decrease by the decrease in powder size is attributable to the surface energy per unit volume. Smaller particles with high specific surface area have higher energy and thus they could be sintered faster in spite of lower sintering temperature. The elastic modulus, compressive yield strength and bending strength decrease linearly with increasing porosity. Load–displacement curves obtained by three-point bending tests show almost linear elastic deformation followed by plastic yielding and strain hardening up to a peak stress. The crack propagates predominantly along interparticle boundaries. This suggests that plastic deformation occurs at bonded contact regions of powder particles. No catastrophic fracture occurs, although the compacts are strained up to a plastic strain of about 16%. The limitation of the conventional powder sintering approach is that pore size and shape are determined by the powder size and shape, which is difficult to control. Porosity is associated with the degree of particle interconnectivity and particle size. For spherical powder particles, the porosity is limited to 50% and pore shape is highly non-spherical [232,233]. Sintered loose titanium powders are often very brittle with poor toughness and are prone to crack propagation at low stresses, cracks are likely to initiate at the sintered necks of individual powder particles or along the interparticle boundaries, especially under fatigue conditions.

The influence of pore size and shape on mechanical properties of loose sintered powder and powder sintered with space-holders is different [38,51]. For powder sintered with space-holders, as introduced in Section 5.3.2, porosity amount, pore shape and macropore-size can be adjusted in a wide range by varying the shape, size and volume fraction of space holder materials, which also exerts great influence on the mechanical properties of the resulting porous material [178,195,196].

Zhuravleva et al. [38] prepared porous β-type nontoxic Ti40Nb alloy by compaction of mechanically alloyed powder mixed with NaCl (125–300 μm) or Mg (350–500 μm) particles as space-holder materials. The size ranges selection of the space-holder particles were based on the anticipated pore sizes. The Ti40Nb powder was consolidated to a green compact under pressure of 700 MPa for 2 h. The green compact was then sintered at 1273 K for 2 h and finally water quenched. For the preparation of porous samples with NaCl and Mg space-holders, they were sealed into argon filled quartz tubes for subsequent sintering. For removing the NaCl particles, the samples were then immersed in distilled water kept at 353 K for 4 h. The Mg particles evaporated after sintering. The complete removal of the space-holder materials requires that the porosity of the sample is larger than 50%. Samples with addition of 50 vol%, 60 vol% and 70 vol% of NaCl powders were prepared. The macropores display the rectangular shape corresponding to the shape of the NaCl particles, since NaCl particles are not deformed by cold pressing. The final porosity of the sintered and washed compacts is about 62%, 70% and 80%, respectively. This means that there are about 10% micro-pores. The average pore size is around 208 μm for the macro-pores, which fits in the range of 200 to 500 μm recommended for biomedical applications [178]. For porous samples prepared with 50 vol% Mg powder. The macro-pores resulting from evaporation of Mg particles have a slightly ellipsoidal shape because of the deformation of the initially spherical Mg particles during uniaxial pressing. The micro-pores have irregular sizes and shapes and are hard to control. The porosity of the sample is ~62%. It can be suggested that the micro-porosity fraction due to incomplete sintering of the Ti40Nb particles was ~12%. The mean macropore size is ~410 μm which also fits to the range of 200 to 500 μm recommended for biomedical applications [178]. The distribution of macro-pore size can be further controlled by careful sieving of the starting Mg powders. The big macro-pores are interconnected via small micro-porosity channels and thus the total porosity is interconnected. The porous samples show almost spherical shape and homogenous distribution of pores throughout a cross-section. Single micro-pores are present in the compact structure of Ti40Nb sample without space-holders and the porosity is ~36%. Compression tests were carried out for samples produced with NaCl as a space-holder. The elastic modulus as well as the strength decreases with porosity increased. Stress–strain curves of these porous samples are alike the ones for cancellous bone. So it can be suggested that the samples deformation behavior follows the same three stages: at first some small number of walls, formed by sintered powder particles, fail; then slight hardening takes place; after that at the third stage a densification of broken walls occurs. The compression curves, where the maximum strain is only ~2%, suggest a brittle fracture of the samples. The compacts with porosity of 36%–80% demonstrated a very low elastic modulus of ~1.5–3 GPa and compression strength of ~10–35 MPa, which is suitable for potential implant material application. The very low elastic modulus is a result of both high porosity of the samples and almost pure β phase composition of the alloy.

5.3.2. Space Holder Method

Space holder method, namely powder metallurgy technique using space holder materials, comes into action with its advantages like adjustable porosity amount, pore shape, and pore size distribution [178,195,196], which depend on the shape, size and volume fraction of space holder used for synthesizing porous titanium and its alloys. The process consists of five major steps: powder selection, mixing, compaction, sintering and spacer removal. First, metal and spacer powders must be selected. After selection, they must be mixed completely to assure the homogeneity of the mixture. The next step is compaction of the powders into a mold under controlled pressure to form a green body. The next stage can be either sintering or removal of the space holder depending on what kind of space holder is used in the process [178]. Figure 18 shows the schematic process of space holder method. By choosing the size, shape and quantity of the space holder, the mechanical properties as well as the pore diameter, morphology and porosity level of the porous metal can be adjusted under relatively low temperatures with low reactivity [178,234–236]. It can produce porous metal samples with greater porosity (Figure 19).

Several materials have been attempted by various investigators as temporary space holders, including urea [178], ammonium hydrogen carbonate [48,196], tapioca starch [237], sodium chloride [238], polymer granules [239] and magnesium [195,240]. In some cases, water soluble spacers such as sodium chloride, potassium chloride, potassium sorbate, or a mixture thereof [241] have been utilized to produce porous titanium and its alloys. Temporary space holders are selected primarily on the basis of different criteria: (i) reaction with titanium; (ii) residue left; and (iii) easy processibility.

However, a general difficulty of this method is the removal of large quantities of space holder material and what kind of space holder material should be chosen. For biomedical applications, any residue left should be either bio-inert or bio-compatible. There is a possibility of absorption of residues left in case of space holders like urea, ammonium hydrogen carbonate and starch by the bio-implants which could import detrimental effect. In fact, most of the space-holders are evaporated at low temperatures or are removed by a dissolution process, generally in water. It is understood that sodium chloride could be a better choice because sodium chloride residue can be easily removed through dissolution in water. Additionally, even if any residue of sodium chloride is left, it would not be toxic for bio-implants. Patnaik [242] presented the advantages of using sodium chloride as space holder, such as low cost, fast dissolution in water, reduced etching of metal during dissolution, as well as much lower toxicity from residual content. Kohl et al. [243] showed good adhesion and proliferation of cells, indicating that any residual sodium chloride did not affect the in vivo performance of porous NiTi. Thus, there is an increasing interest of extensive research on the synthesis and characterization of porous titanium with sodium chloride as temporary space holder. The space holder method, which was first employed by Zhao and Sun [244], was used to fabricate porous Al by using sodium chloride as temporary space holder. Then porous titanium and its alloys with varying porosities (40%–80%) and macro-pore size (>200 μm) were prepared using sodium chloride as space holder through powder metallurgy [176]. Urea and ammonium hydrogen carbonate were utilized for porous titanium and magnesium production for the first time by Wen et al. [178]. Torres [245] focused on finding the optimum sodium chloride percentage used as space holder, and finding an improved procedure to eliminate the salt before the sintering step and concluded the optimum conditions for salt dissolution were: water temperature between 50 and 60 °C, maximum immersion time per cycle (4 h) and without agitation. These conditions ensure a faster route to salt dissolution, as well as to the best structural integrity of samples. The most important factors in controlling dissolution rate were: salt content, water temperature and immersion time per cycle.

In recent years, temporary space holders such as ammonium hydrogen carbonate or urea have been used to adjust porosity and pore structures [246]. However, the rapid decomposition of ammonium hydrogen carbonate at a lower temperature led to an imprecise control of porosity and pore structures, as well as an unsatisfactory repeatability of mechanical properties of porous alloys. In addition, the presence of sharp corners or notches of the pores in the alloys, which are due to the needle-like or flake ammonium hydrogen carbonate space holder particles, can obviously result in stress concentration and thus deteriorate mechanical properties [216]. Besides, the spacer particles, like urea and ammonium hydrogen carbonate, which have relatively low melting/dissociation temperature, are removed by thermal treatment and thus may cause oxidation resulting degradation of mechanical properties. Additionally, they may leave residue and cannot be removed completely, which makes them useless for many residue free requiring applications.

Porous Ti-10Nb-10Zr (wt%) alloy, with macropores of 300–800 μm and micropores of several micrometres, was successfully fabricated by space holder method [175]. The blended TiNbZr powder was firstly mixed with ammonium hydrogen carbonate (NH₄HCO₃), which was used as the space-holder material. The size of space-holder particles was 500–800 μm. The mixture of Ti, Nb, Zr powder and NH₄HCO₃ was then cold pressed into green compacts in a 50 ton hydraulic press. The green compacts were sintered in a vacuum of 10⁻⁴–10⁻⁵ Torr in two steps. The first step was carried out at 175 °C for 2 h to burn out the space-holder particles. In the second step the compacts were heated up to 1200 °C and held for 10 h. Porous TiNbZr alloys with porosities of 42%, 50%, 59%, 69% and 74% were fabricated by adding 20 wt%, 30 wt%, 40 wt%, 50 wt% and 60 wt% ammonium hydrogen carbonate to the powder mixture. Porous Ti-10Nb-10Zr alloy with 59% porosity exhibited an elastic modulus and yield strength of 5.6 GPa and 137 MPa, respectively, displaying significantly higher strength than unalloyed titanium scaffolds of the same porosity. Porous titanium was fabricated by powder metallurgy using urea as spacer material, heat treated at 180 °C/2 h in air to eliminate the spacer particles, sintered at 1200 °C/1 h under a vacuum (10⁻⁷ torr), exhibited three dimensionally interconnected pores. The average pore size was 480 (±210 μm) and total porosity was 36% (±2.4%) [247]. The microstructure showed interconnected pores and small isolated pores. Bram et al. [248] used urea powders as space holders, which could be removed at temperatures below 200 °C and had minimal contamination of titanium powders. Subsequent sintering at 1400 °C for 1 h resulted in porous titanium with porosities in the range of 60 vol%–70 vol% and pore diameters in the range of 0.1–0.24 mm, depending on the urea particle size. Wen et al. [178] used ammonium hydrogen carbonate as space holder with decomposition temperature of 200 °C and a sintering temperature of 1200 °C for 2 h. The resulting porous titanium had a porosity of 78% and exhibited a compressive strength of 35 MPa, whilst elastis modulus was 5.3 GPa. These values were a close match to those for cancellous bone. Porous Ti-10Mo alloy was successfully prepared from mechanical alloying particles and then consolidated by sintering using ammonium hydrogen carbonate as space holder [249]. Elemental metal powders of Ti were blended at Ti:Mo weight ratio of 9:1 in ball milling system and then milled for 5 h with a weight ratio of ball to powder of 20:1 and rotation rate of 300 rpm. The ball milling was carried out under a high pure argon atmosphere. The blended Ti-10Mo powder was mixed with ammonium hydrogen carbonate (NH₄HCO₃). The mixture powders were cold compressed under 500 MPa for 1 min to form a green compact. Then, the resultant green compacts were heated to 175 °C to burn out the spacer-holding particles in air. After that, they were sintered at 1200 °C for 5 h in a highvacuum furnace and then cooled in the furnace. The compressive strength and elastic modulus had been tailored in the range of 174–646 MPa and 20.2–43.7 GPa, respectively. There was a substantial increase in porosity, from 5.5% to 37.5%, in mean pore size, from 20 to 130 μm, with the addition of 0–25 wt% ammonium hydrogen carbonate.

Magnesium played an important role in preventing contamination due to thermodynamical conditions. Homogeneous porous TiNi alloys having porosities in the range 59%–81% have been produced using magnesium powder as pore forming agent [240]. Prealloyed TiNi and Mg powders (50%, 60%, 70%, and 80% Mg by volume, balance TiNi) were mixed for 30 min using 5 wt% polyvinyl alcohol (PVA) solution (2.5 wt% PVA + water) as the binder prior to compaction. Particle size of Mg powders was in the range 250–600 μm, averaging at 450 μm. Then the mixtures were compacted in a double ended steel die at 400 MPa using a hydraulic press. Subsequently, compacts 10 mm both in diameter and height were heated to 1100 °C in purified argon atmosphere at a rate of 10 °C/min which is low enough to allow debinding of PVA below 600 °C and simultaneous Mg evaporation and sintering at higher temperatures. Sintering time at 1100 °C was 1 h. Sintered porous specimenswere cooled in the cold zone of the furnace at a rate sufficient to prevent formation of intermetallics. The processed alloys exhibited interconnected open pores spherical in shape and with an average macro-pore size around 400 μm. Elastic modulus values observed to extend from 0.5 to 9 GPa, while yield stresses were in the range 2.2–31.0 MPa and compression strength values were in a broad range between 5.8 and 93.2 MPa depending on porosity content were comparable to that of cancellous bone. Porous TiNi alloys with porosities in the range of 38%–59% and homogeneous pore distribution, using 20 vol%–50 vol% spherical magnesium with a mean diameter around 450 μm as space formers, were also fabricated through the same fabrication procedure [250]. The porous TiNi alloys consist of fully spherical pores free from stress concentrations. Partial superelasticity has been observed in as-sintered and aged conditions. Linear and complete superelasticity up to 5% has been achieved after stress cycling.

Porous Ti-6Al-4V alloy was synthesized by space holder method as a result of evaporation of magnesium to achieve desired porosity content [194]. Magnesium powder with a wide range of distribution (300–1500 μm) and an average particle size of 660 μm was used. In the space holder technique used for the production of porous Ti-6Al-4V, firstly, 5 wt% PVA solution was added as a binder to Ti-6Al-4V/magnesium powder mixtures containing 30 vol%, 40 vol%, 50 vol%, 60 vol% and 70 vol% magnesium and the powder mixture mixed thoroughly for 0.5 h to ensure a homogenous mixture. Then, the mixtures were cold pressed at around 500 MPa using a double-ended steel die to obtain green compacts 8 mm in height and 10 mm in diameter. Eventually, for debinding and removal of magnesium by evaporation, the compacts were slowly heated with a heating rate of 12 °C/min up to 1200 °C at which the samples were kept for 1 h under high purity flowing argon gas for sintering and complete removal of magnesium. Final products contained porosities in the range of 43%–64% with an average macro-pore size 485–572 μm, and calculated elastic modulus and yield strength were in the range 1.42–14.7 GPa and 28.2–150 MPa, respectively. A power law relation was obtained between compression yield strength, σ_y, and macro porosity fraction, P_macro, in the form of A × (1 − P_macro)ⁿ, similar to commonly used minimum solid area models. The proportionality constant “A” in relation was found to be 330 MPa and defined as the yield strength of porous cell walls. Similar to model of Nyce and Shaffer the strength of porous cell walls changes with the square of neck size ratio (X/D)², where X and D are the average neck and powder particle size, respectively, indicating a strong dependence of strength on neck size, which are actually the load-bearing areas of porous materials.

Porous titanium with an average macropore size of 200–400 μm and porosities in the range of 10%–65% was manufactured using polymethyl methacrylate (PMMA) powders with the size of 71–100 μm, 100–154 μm, 154–200 μm, 200–315 μm, 315–400 μm and 400–630 μm as spacer particles [209]. The mixtures containing 10 vol%–70 vol% PMMA and CP-Ti powders were compacted as a cylindrical rod (ϕ 10 mm × 10 mm) under the uniaxial pressure of 500 MPa to form green bodies. The green compacts were heated up to 250–450 °C for 5 h in a vacuum furnace under the pressure better than 7 × 10⁻⁴ Pa to remove the space holders, and then sintered at 1300 °C for 2 h. The compressive strength and elastic modulus of resultant porous titanium were observed in the range of 32–530 MPa and 0.7–23.3 GPa, respectively. With the increasing of the porosity and macropore size, the compressive strength and modulus decreased as described by Gibson-Ashby model. Porous Ti-24Nb-4Zr-8Sn alloy was fabricated by space holder method also using 30 vol% PMMA with an average pore size of 244–372 μm and porosities in the range of 35%–43% [62]. The mixers were cold rolled to form green compacts. Then the compacts were heated to 400 °C for 2 h to remove the PMMA space-holders and finally sintered at 1100–1300 °C for 4 h under vacuum condition and cooled in the furnace. The compressive strength and dynamic modulus of resultant porous Ti-24Nb-4Zr-8Sn alloy were observed in the range of 21–260 MPa and 1.54–6.66 GPa, respectively, which match those of cortical bone. Porous titanium with 64%–79% porosities and open cellular 100–300 μm pore sizes by using starch through powder metallurgy was successfully fabricated [237]. The yield strength and modulus were 23–41 MPa, and 1.6–3.7 GPa, respectively.

The space holder method is more attractive to make highly porous titanium and its alloys compared with loose powder sintering, slurry foaming [251] and hollow sphere sintering [229], which results in limited porosity (<50%) in the porous body. The space holder method can yield porous parts with desirable pore size, shape, volume and distribution using an appropriate space holder material, usually removed at low temperature without excessive contamination.

5.3.3. Spark Plasma Sintering

Conventional sintering of titanium and its alloy requires compacting metal powders in a die at room temperature, and subsequently a high sintering temperature (1200–1400 °C) in a high vacuum (4 × 10⁻⁴ Pa) for a long time (24–48 h) [252]. Not only are high pressures, high temperatures, and long times required, but in the case of reactive materials, such as titanium and its alloys, an inert atmosphere is also inevitably required. The high temperatures involved in these processes, however, result in detrimental changes in the microstructure and mechanical properties. The difficult sintering process limits the usage of sintered titanium and its alloys.

Recently, there has been a growing interest to create porous implants for orthopedic applications by using spark plasma sintering (SPS) [253,254], which involves the application of an external current to assist powder consolidation. This technique is known under different names, such as field assisted consolidation technique [255], electrical field activated sintering [256], plasma activated sintering [257], and electrical discharge compaction [258]. All techniques have in common the combination of an electrical discharge with rapid heating and pressure application to achieve fast sintering of powders. SPS is a high efficient and energy saving powder consolidation and sintering technology. The SPS studies on porous Ti-based alloys were mainly using low temperature and low pressure to decrease the relative density of samples [259–263]. The aforementioned methods are useful in that they can easily sinter titanium and its alloy powders, because ionization in the plasma created by the high current discharge can melt the local oxide surface film on the particles, bringing the particles into contact with each other, and allowing junctions to be formed. Electrical discharge compaction has been used to produce porous CP titanium implants [232,264].

Porous CP titanium with porosities of 30%–70% and pore sizes of 125–800 μm were prepared by combining SPS with sodium chloride dissolution methods [224], as shown in Figure 20. The porous titanium sintered at 700 °C for 8 min under 50 MPa showed pure α phase. All the samples exhibited highly interconnected structure and uniform pore distribution, as shown in Figure 21. The macroporous titanium with yield strength 27.2–94.2 MPa and elastic modulus 6.2–36.1 GPa had a potential to be used as bone implants. Porous titanium and Ti-5Mn alloy were successfully prepared by using SPS technique at pressureless conditions using ammonium hydrogen carbonate in size 250–245 μm as well as blowing agent TiH₂ in size <44 μm at different temperatures for 5 min [256]. Experimental results showed that pure Ti sample, which sintered under pressureless condition at sintering temperature 1000 °C the porosity was 53% and elastic modulus was 40 GPa. The Ti5Mn alloy indicated a good pore distribution, and the porosity decreased from 56% to 21% by increasing the sintering temperature from 950 to 1100 °C. Elastic modulus was increased from 35 to 51.83 GPa with increasing of the sintering temperatures from 950 to 1100 °C. The elastic modulus data obey the Gibson-Ashby mode well.

SPS is a combination of strong electric field, stress field and temperature field under the action of consolidation method [265], with advantages of low sintering temperature, fast cooling rate, integration of sintering and heat treatment, etc. SPS can produce field effect between powder particles during sintering [265] and effectively inhibit grain growth with controllable structure, which is a kind of interface clean nanostructured/ultrafine-grained crystal material forming technology [107]. It should be noted that due to high melting point features of alloy components, fabrication of Ti-based alloys by traditional sintering method might bring some shortcomings, such as too high sintering temperature, too long sintering temperature, and too high vacuum condition. As an advanced materials forming technology, SPS can provide relatively low sintering temperature, short holding time, and rapid cooling during the whole process. Especially, an outstanding advantage is that SPS can provide a high heating rate to several hundred K/min. Therefore, alloys with controllable microstructures can be obtained by SPS due to its special sintering mechanism on powder particles [107]. SPS provides a new way of fabricating nanostructured/ultrafine-grained porous biomedical titanium and its alloys with excellent properties and provide an alternative method for synthesizing new biomedical materials.

5.3.4. Microwave Sintering

Roy et al. [266–269] firstly obtained a fully sintered dense metal in a single-mode microwave cavity, after which microwave assisted powder metallurgy methods were developed for the sintering of metal alloys and their composites.

Tang et al. [270] fabricated porous Ti-6Al-4V alloy/TiC composites by the in situ formation of TiC from the reaction between multi-walled carbon nanotubes and titanium alloy powders, using the microwave sintering technique as shown in Figure 22. Ti-6Al-4V powders were mixed with 14.5 vol% of multiwalled carbon nanotubes (MWCNTs) and the particle size of Ti-6Al-4V was reduced to about 5 μm by ball milling. The MWCNTs acted as microwave susceptors as well as reactants for the rapid sintering of Ti-6Al-4V alloys. The green compact of the Ti-6Al-4V/MWCNTs powder was sintered in a 1.4 kW, 2.45 GHz bi-directional microwave emission furnace without using any protective gas environment. The microwave sintering was completed within 2 min. The surface temperature of the sample reached 1620 °C during sintering, as measured by a pyrometer. The processed porous composite had a porosity of approximately 25% and displayed compressive strength of 270.41 ± 24.97 MPa, yield strength of 145.48 ± 27.28 MPa, elastic modulus of 10.87 ± 2.46 GPa and compressive strain of 3.75% ± 1.11%. The Vickers hardness value in the TiC region (545.4 ± 13.9 HV) was significant higher than that in the titanium alloy region (435.9 ± 12.9 HV). An overall Rockwell hardness of 47 HRA was measured.

Traditional powder metallurgy processing of Ti-based alloys and their composites requires a long sintering time. Although a prolonged heating time can ensure complete sintering, it also increases the chance of impurities formation, such as oxides. In such cases, the sintering needs to be performed in either a protective gas atmosphere or in a vacuum [271]. Microwave heating has been used for the processing of materials such as ceramics, metals and composites, taking advantage of time and energy saving, reduced processing cost, decreased sintering temperature, better production quality, and lower environmental hazards [272]. The transfer of microwaves to the materials directly can result in a faster generation of heat, thereby achieving rapid and uniform heating.

5.3.5. Combustion Synthesis

A recently developed effective method of producing high purity porous alloy is a process known as combustion synthesis. Particle fusion is obtained through an extremely rapid self-sustaining exothermic reaction driven by the large heat released in the synthesis. The exothermic reaction can then be instigated under two different regimes: (a) the reactants are gradually heated until reactions take place simultaneously throughout the whole sample; and (b) self-propagating high-temperature synthesis (SHS), once ignited, a strong exothermic reaction propagates as a combustion wave through the entire mixture, without requiring additional energy. Various processing parameters such as the reactant particle size of powder, the use of a binder and the compaction pressure affect the final microstructure and porosity of the sample [273]. A schematic of the stages in this process is shown in Figure 23.

Several researchers have investigated the use of combustion synthesis in the fabrication of porous NiTi alloys [274–281]. Li et al. [282] successfully prepared porous NiTi alloys with anisotropic pore structure and porosity of 54 vol% by SHS. The mixed powder with a Ti: Ni atomic ratio 1:50 was cold pressed in a newly designed die. The green compact (ϕ 33 mm × 80 mm), whose porosity is 40 vol% after being cold pressed with a pressure of 75 MPa, was placed in a reaction chamber and ignited at a preheating temperature of 550 °C. The NiTi SMA synthesized has an open porous structure, also its pores and channels with various size and shape are three-dimensionally interconnected. 100% Ni-Ti intermetallic compounds are formed in the SHS synthesized Ni-Ti SMA. A porous structure formed along the propagation direction of the combustion wave was shown in Figure 24.

Porous NiTi alloys were also fabricated by combustion synthesis of powders with the help of ZrH₂ as foaming agent and TiB₂ as endothermic agent [283]. 32.5 vol% TiB₂ and 0.5 wt% ZrH₂ were added in the NiTi mixture. Cylindrical compacts (ϕ = 15 mm, h = 15 mm, relative density = 0.7) were produced by cold pressing of the powder blend with compacting pressure of 165 MPa. The compact was placed inside a crucible and ignited in a chamber filled with argon gas. A thermocouple was inserted in the powder compact to measure its temperature during the combustion reaction. The powder size of Ni and Ti affected the ignition temperature of the combustion reaction, cell morphology and microstructure of this alloy. Porosity and pore diameter of samples decrease by increasing the powder size of Ni and Ti. Porosity of combustion synthesized samples using different size Ni powder and Ti powder was 14.5%–60.0% and 42.6%–60.0%, respectively.

This process can reduce the amount of energy required to fabricate porous materials with high melting points by utilizing the heat that was produced during combustion reaction. Also, it was found that in addition to an endothermic agent (TiB₂), a foaming agent was necessary to make such porous materials [284]. The heat of NiTi formation reaction was so high that the viscosity of molten NiTi during the reaction was very low. An advantage of this fabrication process is the high purity of the resulting porous materials, which is largely due to the expulsion of volatile impurities under the extremely high temperatures. In vivo studies suggest strongly that there is natural bone ingrowth with no apparent immune response in the composite materials produced in this process [273].

5.3.6. Metal Injection Molding

Metal injection molding (MIM) could be an attractive process combining the technique of polymer injection molding with the advantages of powder metallurgy. The possibility of near-net-shape production combined with a high potential for automation makes the MIM process suitable for the industrial production of porous titanium and its alloys. The MIM process starts with the mixing of metal powder with one or more polymers to produce a feedstock, followed by molding of the homogenized feedstock into shape parts, removal of polymer and final sintering.

Chen et al. [285] fabricated porous titanium by using MIM method with space holder sodium chloride. Mold temperature was 30 °C, injection temperature was 155 °C, and injection pressure was 100 MPa. The compacts were sintered at 1150 °C for 2 h under a vacuum of 1.33 × 10⁻³ Pa. Porous titanium powders with porosities in the range of 42.4%–71.6% and pore size up to 300 μm were obtained. Mechanical test showed that with sodium chloride content increased, compressive strength decreased from 316.6 to 17.5 MPa and elastic modulus decreased from 3.03 to 0.28 GPa. These mechanical properties were similar to those of cancellous bone. The microstructures of porous titanium samples were shown in Figure 25. It could be seen there were interpenetrating macropores with size in the range of 50–300 μm which were formed by spacer holder removal. Macropores in sample with 42.4% porosity are mostly isolated (Figure 25a); those in sample with 71.6% porosity are well interconnected (Figure 25b). Simultaneously, the cell walls of the porous titanium are rough and honeycomb-like. The interconnected pore structure and rough wall-surface of macropores were appropriate for ingrowth of new bone tissues and transport of body fluids [286].

5.3.7. Capsule-Free Hot Isostatic Pressing

Porous NiTi alloys were initially fabricated by combustion synthesis [274,277,278,287]. However, the porous alloys fabricated by these this process showed inhomogeneous constituents and non-uniform pore distribution. These are not beneficial for tissue ingrowth. In addition, their mechanical properties are not as good as expected; in particular the shape memory effect (SME) and superelasticity (SE) deteriorate remarkably.

Recently, capsule-free hot isostatic pressing (HIP) process was developed to improve the quality of porous NiTi alloys [279,280]. This process can densify the dense part of the alloy and decrease sharp angles of open pores through extra high pressure argon gas [279]. Therefore, fabricated porous alloys showed nearly spherical pore shape with homogeneous pore distribution [279,280], and exhibited good SME and SE [279]. However, this process still has some problems, such as relatively low porosity and small pore size. For example, average pore size is 129 μm according to Weibull distribution calculation [280]. Although some research was carried out to improve the pore characteristics by using TiH₂ as the vesicant in fabrication of porous NiTi alloys [281], the pore size turned to be small after the addition of TiH₂ and use of TiH₂ is dangerous in fabrication process. More recently, a new process was reported by using sodium chloride as pore forming agent for fabricating near-net-shape porous NiTi parts [238]. The results showed that the porous alloys exhibited large pore size (355–500 μm) and sodium chloride did not deteriorate functional property of the alloys. However, the downside of this process was that removing sodium chloride was relatively complicated, and most of pores were not connected, while high connectivity of pores is highly required for use as tissue implants.

Zhang et al. [288] reported a new attempt for fabricating porous NiTi alloys with adjustable and high porosity as well as large pore size by use of ammonium hydrogen carbonate and an updated capsule-free HIP process. Firstly, elemental Ti and Ni powders were mixed with ammonium bicarbonate. Then the mixed powders were blended by ball milling for 4 h. After that, the blended powders were cold pressed into cylindrical compacts of ϕ 15 mm × 25 mm under a pressure of 100 MPa. The green samples were heated, under protection of high-purity argon, to a temperature of 200 °C for 30 min to decompose the pore forming agent. Then the treated green samples were sintered at 1050 °C (under protection of high-purity argon gas) in the furnace for 3 h under hot pressing pressure of 50 MPa. Finally, all samples were aged under 450 °C for 2 h, and then quenched into ice-water. The porosity of the fabricated porous alloys increased linearly from 30.4% to 48.4% with addition of 0–11.9 wt% ammonium hydrogen carbonate, and the alloys showed more uniform pore distribution with most pores larger than 300 μm (Figure 26).

5.3.8. Solid State Isothermal Foaming Technique

Nugroho et al. [289] have investigated the use of pressurized bubble entrapment in the manufacture of porous Ti-29Nb-13Ta-4.5Zr alloys, which were used instead of traditional prealloyed powders. Firstly, the elemental powders were blended for 30 min in a roller mixer. Following this, approximately 50 g of alloy powder was placed into a stainless steel can with an inner diameter and length of 28 and 30 mm, respectively. Each can was subsequently evacuated to approximately 10⁻² Pa and then backfilled with 0.48 MPa of argon and sealed. The pressurized cans were then densified using hot isostatic pressing (HIP) at 1000 °C for 2 h or 1100 °C for 4 h under 100 MPa of argon and then furnace cooled. The HIP-ed billet specimens were then rapidly heated in a graphite element vacuum furnace to 1150 °C (for the 1000 °C HIP-ed specimens) or 1350 °C (for the 1100 °C HIP-ed specimens) and held for 10 h (isothermal foaming process). During this time the pressurised pores were expected to expand due to creep of the surrounding metal. The resulting alloys were found to contain porosity and sizes ranging between 20%–40% and 20–200 μm, respectively, with some interconnected pores. Backscattered SEM and OM micrographs illustrating pore morphologies for the specimens foamed at 1150 and 1350 °C using a backfilled argon pressure of 0.48 MPa were investigated. The specimen foamed at 1150 °C was noted to contain approximately 30% porosity with uniform pore size and morphology. On the other hand, the pore size distribution in the specimen foamed at 1350 °C varied with the pores at the specimen center being discrete and mainly equiaxed with typical diameters of 20–50 μm; the porosity in this region was found to be approximately 14%. In contrast, the pores close to the specimen surface had noticeably expanded with the porosity level achieving approximately 45% and pore sizes of typically 100–200 μm.

However, complex-shaped porous implants cannot be fabricated using the solid state isothermal foaming technique and the properties of the samples fabricated are mechanically inadequate.

5.3.9. Freeze Casting and Reverse Freeze Casting

Among the prospective methods of fabricating porous metals, freeze casting (FC) has been widely studied with regard to mechanical performance of the scaffold. FC method is quite simple and offers the possibility to produce many parts. This method has great potential, as it can be employed for the fabrication of various kinds of porous materials (e.g., polymers, bioglass, ceramics and metals). In addition, a variety of structures can be produced for a fixed material composition by adjusting the process parameters, such as the freezing temperature and time, freezing direction, freezing vehicle, etc. [290]. By varying the processing conditions, wide variations in mechanical properties of freeze casting materials can be achieved [291–293]. The FC technique can lead to randomly distributed pores and enables a porous material with a compressive strength close to that of bone to be obtained, but its main drawback underlined by Yook et al. [293] is pore size does not exceed 300 μm for a long manufacturing time of 7 days. The control of the elastic properties anisotropy seems to be difficult. Jung et al. [294] used a new manufacturing technique called dynamic FC to create porous titanium scaffolds with a uniform porous structure and good ductility. Chino and Dunand [295] developed directional FC of aqueous titanium slurry, followed by ice sublimation and powder sintering. Porous titanium with aligned and elongated pores was created from this technique.

However, there are some limitations to FC method. Firstly, the degree of alignment is a critical issue when camphene is used as the freezing vehicle. It is difficult in practice to increase the length of aligned camphene dendrites, as the freezing rate of camphene is slower than that of water [296,297]; Secondly, most metal powders cannot be employed in the FC method; Thirdly, the maximum pore size for the existing FC method was 400 μm, which is the limit of the diffusion distance of the partially remelted frozen camphene; Finally, the merging and diffusion phenomena of the remelted camphene stopped when the thickness of the camphene dendrites reached 300 μm [293,298]. Therefore, a radically new solution was necessary in order to overcome the limitations of FC method.

Yook et al. [299] proposed a new method, referred to as reverse freeze casting (RFC), a suggested way to produce highly porous titanium with aligned large pores, and it was suggested that the powder particle size should not exceed 50 μm.The major concept of this method is based on the migration of raw powder, rather than the growth of camphene dendrites. Camphene-based slurries of titanium with adequate powder content (20 vol%) were prepared by ball milling at 60 °C for 6 h. After sufficient unidirectional freezing of camphene, the slurry was poured onto the frozen camphene, followed by immediate quenching to −20 °C. Thereafter, the resulting double layered material, consisting of the frozen camphene layer and the camphene–titanium slurry on top of it, was kept at 45.5 °C for various periods of time (1–3 days). This allowed for sufficient migration of the raw powders down the channels within the directionally solidified camphene. The exposure temperature of 45.5 °C was determined empirically by studying the solidification behavior of the slurries by means of DSC. After demolding, the green bodies were freeze dried for more than 10 h to remove the frozen camphene and then sintered at 1300 °C in a vacuum furnace. After the freeze drying and sintering, all of the fabricated samples had long aligned pores up to 500 μm in size, as shown in Figure 27.

The average size of the aligned pores was 300 μm, and their maximum length was 3 cm. The average porosity of titanium with highly aligned pores was controlled by tuning the casting time. With casting time increased, more titanium powder particles migrated downwards and the resulting columnar struts started to link up with one another. As the casting time increased from 24 to 48 h, the porosity was changed from 69% to 51% and the average wall thickness increased from 179 to 310 μm; the compressive strength of the porous titanium scaffolds in the direction of pore alignment increased from 121 to 302 MPa and the elastic modulus of the samples was in the range of 2–5 GPa.

5.3.10. Slip Casting and Gel Casting

Researchers have been using ceramic techniques to fabricate near-net-shape porous titanium products and this has emerged as an attractive approach [300].

Slip casting is a conventional near-net-shape ceramic technique that is commonly used to manufacture dinnerware and sanitary ware. It is a filtration process in which solvent based powder slurry is poured into a plaster mold, creating capillary forces as a result of the porosity and these absorb the solvent from slurry. Over a period, the powder compact will be formed in the mold. Viscosity is a critical parameter in slip casting [301]. This process involves binder, plasticizer and dispersant for controlling the viscosity of slurry. Neirinck et al. [301] presented slip casting of a mixture consisting of slurry of titanium and titanium hydride powder to form a green compact through absorption of solvent by plaster mold, followed by sintering. However, safety is a major concern with this process due to the high reactivity of titanium hydride powder. Asthana et al. [302] indicated that too high or too low a viscosity slurry yielded poor quality sediment. A solvent is an important component to make a slip in the slip casting. This offers a media for powder interactions and the interaction between the powder and organics additives. Although an organic solvent is commonly used as a dispersion media in colloidal processing due to its excellent compatibility with organic additives, water is emerging as an alternative because it is environmentally friendly, low cost and safe [303]. Xu et al. [304] demonstrated that porous titanium compacts can be produced from water based titanium slurry using a slip casting approach. This opens up a simple and low cost manufacturing route for porous titanium production to shape titanium compacts with controlled levels of porosity. An important aim was the optimization of the water based titanium slurry formulation to minimize the content of organic additives. A lower organic additive content reduces the probability of carbon contamination in the sintered titanium products.

Gel casting process, also a near-net-shape ceramic processing technology [305], can be easily adjusted, and porosity and mechanical properties of porous titanium parts can be controlled by altering the parameters. Furthermore, the easy preparation of complex shaped products combined with the low costs of the molds make gel casting more appropriate for the fabrication of porous titanium implants, especially in personal medical treatment areas. In this process, the powders, organic monomer binder, dispersant and solvent are mixed together to first form slurry. Subsequently, a crosslinking reaction occurs when initiator and catalyst are added to form an in situ polymerization green body. After solvent evaporation, the polymer acts to bind the powders and retain the desired shape. By altering the solid loading of the slurry, the porosity and mechanical properties of porous titanium parts can be changed, therefore easily obtaining custom made performance. Erk et al. [306] demonstrated the feasibility of a gel casting process to fabricate porous titanium in complex shapes. A unique porous titanium screw for implant applications was fabricated with a 20% porous core and a higher porosity near the surface. Yang et al. [307] have used gel casting and subsequently sintering to prepare porous Ti-Mo and Ti-Nb alloys with three dimensionally interconnected porosity and good property. Schematic representation of the gel casting process is shown in Figure 28. Sintering temperatures of 900 to 1050 °C for 2 h were used to fabricate Ti-Mo and Ti-Nb alloys with varying porosity. The results showed that total porosity ranging from 39% to 50% and the pore size ranging from 5 to 120 μm can be achieved by changing the gel casting parameters. These porous Ti-Mo and Ti-Nb alloys have elastic modulus range from 5 to 18 GPa and compressive strengths in the range 141–286 MPa.The bending strengths are in the range 89–131 MPa for Ti-Mo alloys and 79–120 MPa for Ti-Nb alloys. Microstructure study indicated that the pores were irregular in shape with uniform pore size distributed in the two samples processed under the same processing parameters (Figure 29), promising for load-bearing implant. Also, Yang et al. [308–310] prepared porous Ti-Co and Ti-Mo alloys by gel casting. Sintering temperatures of 1000 and 1100 °C for 2 h in vacuum at 3.5 × 10−³ Pa were used and achieved open and closed three dimensional pore morphologies and total porosity ranging from 38.34% to 58.32%. In contrast to porous titanium fabricated by gel casting, the compression and bending strengths of porous Ti-based alloys were significantly increased by adding Co and Mo with elastic modulus ranging between 4 and 25 GPa and compressive strength from 82 to 333 MPa, which is close to that of human cortical bone. Moreover, the bending strengths are in the range 64–109 MPa for porous Ti-Co alloys and 89–204 MPa for porous Ti-Mo alloys.

5.3.11. Slurry Foaming and Plasma Sprayed Coating

Figure 30 presents processing steps for fabricating porous titanium by slurry forming method using H₂O₂ as forming agent [311]. The resulting porous titanium presented three dimensional structure (Figure 31) with porosity 58% and average pore size up to 460 μm. The compressive strength and elastic modulus were 190.7 MPa and 4.15 GPa, respectively. These values well matched those of human bone [312]. However, current techniques that use foaming agents, either in solid state sintering processes or in molten metal techniques, have inherent limitations, such as contamination, presence of impurity phases, limited and predetermined part geometries, and limited control over the size, shape and distribution of the porosity [51,178].

Plasma spraying method can be used to create rough solid surface textures, porous surface coatings on solid cores and also fully porous structures. Porous surface coatings on solid cores and fully porous structures whether they are open cell or closed cell, allow bone ingrowth into pores. A schematic description of the plasma spraying process is shown in Figure 32. Porous coatings with varying degrees of porosity can be created on the substrate by adjusting the spraying parameters [313].

Hahn and Palich [44] first described titanium plasma sprayed coating for fabricating porous coated implants. By choosing an appropriate gun-to-substrate distance, a thin coating (approximately 900 μm thick) with porosity that varied from 0 at the substrate interface to about 50% at the coating surface was formed. However, coatings prepared with this method result in irregular and rough surfaces with some macro and micro pores and the pore interconnectivity is quite low compared to other techniques. Furthermore, it is difficult to form bulk structures with large pores (>100 μm). Due to short reaction time, deposited material is usually a composite of the starting material forming the matrix and in situ synthesized phases. This is indispensable for plasma spraying of titanium, which is a sensitive material prone to oxygen and nitrogen absorption [313].

5.3.12. Entangled Titanium Wire Materials

A group of entangled titanium wire materials with nonwoven structure were fabricated by using 12–180 MPa forming pressure, which have porosity in the range of 48%–82% [57].CP Ti (Grade 1) wires with diameters of 0.08 and 0.15 mm, respectively, are selected as raw materials. The spatial architecture constructed with wires is controlled by the wire-twisting methods and intertwisting procedures. A general procedure for forming nonwoven titanium wire materials is as follows: firstly the wire is entangled by winding and twisting to form a precursory bundle of twisted wires. Then the bundle is put into a specially designed steel mold for compaction and shaping. A high pressure in a range of 12–180 MPa is necessary. The holding time under certain pressure and at high temperature (around 650 °C) is about 10 min for avoiding wire-rebound and shape-finalizing. After that, the sintering temperature about 1200 °C and holding time 60 min were used. After the sintering a step-cooling followed with annealing at 650 °C is carried out for plasticity recovering of the titanium wires. The mechanical behavior of the porous titanium wire materials strongly depends on raw material (titanium composition, treatment status, and wire diameter), entangled wire architecture, and porosity. The pores in the resulted materials were irregular in shape, which had a nearly half normal distribution in size range. The yield strength, ultimate tensile strength, and elastic modulus were 75 MPa, 108 MPa, and 1.05 GPa, respectively, when its porosity was 44.7%. The low elastic modulus was due to the structural flexibility of the entangled titanium wire materials. The average pore size depends on the porosity and the wire diameter. The larger porosity and the larger wire diameter lead to the larger average pore size. The quasi-ordered entangled wire materials are practically useful because they are more homogeneous in porous structure and easily mass manufactured [314]. Figure 33 showed some cylindrical samples of the quasi-ordered entangled wire materials with different porosities.

So far, several shapes such as cylinders, rectangular prisms, tubes, sheets, and flat dog-bones have been made [57]. This kind of material is very promising for implant applications because of good toughness, perfect flexibility, high strength, adequate elastic modulus, and low cost. It should be noticed that the entangled wire material is a kind of fiber or wire mesh, in which the entangled wire can be either stochastic (like nonwoven) [315] or ordered [314]. Originally, titanium wire or fiber materials were introduced for producing porous surfaced implants [316]. The relationship between the mesh structure and the mechanical properties has not been established. Furthermore, the basic aspects in terms of the fabrication techniques (especially for the three dimensionally porous titanium wire materials), the structure characterization, and the mechanical properties tailoring are still open for investigation.

5.3.13. Replication Method

The sponge replication method seems to be effective because of its flexibility in producing different structures, and its simplicity and low cost [297]. Additionally, this method can achieve high porosity and good interconnection between pores [297,317], which is beneficial to human bone ingrowth and the vascularization of newly formed tissue [37]. These porous structures have many applications ranging from spinal fixation to acetabular hip prostheses, dental implants, permanent osteosynthesis plates, and intervertebral discs [246].

Highly porous titanium was prepared through sponge replication method [223]. Schematic representation of the three-step replication process is shown in Figure 34. The samples were slowly heated to 400 °C at a rate of 2 °C/min and maintained for 2 h to burn out the polyurethane foams and sintered at three selected temperatures (800, 900 or 1000 °C) for a hold time of 1 h at a heating rate of 20 °C/min in air in an electric furnace. The total porosity of the sintered porous titanium is larger than 79%, and open porosity is larger than 74%. The open porosity ratio is in excess of 0.92, suggesting that most pores are interconnected. Moreover, the strength (6–10 MPa) and elastic modulus (0.4–0.6 GPa) of the sintered porous titanium conform to the basic mechanical property requirement of cancellous bone.

Li et al. [251] utilized this method to produce porous titanium and its alloy structures. Polyurethane porous materials were immersed in titanium slurry comprising Ti-6Al-4V powder (70 wt%), H₂O (20 wt%) and ammonia solution. The sample was subsequently dried and the process was repeated until all the struts of the polyurethane porous material were coated with Ti-6Al-4V powder. After thermal removal of the polyurethane scaffold and binder, samples were sintered at 1250 °C for 2 h in a vacaum furnace. A reticulated open cell porous material with hollow titanium struts remained, as shown in Figure 35. Open cellular pores are necessary for human bone ingrowth, and extensive body fluid transport through the porous scaffold matrix is possible, which can trigger bone growth if substantial pore interconnectivity is established [31,232]. The porous Ti-based alloy fabricated in this way was found to possess a porosity of 88% and compressive strength of about 10 MPa.

It was found that the rheological properties of the titanium slurry played an important role in the impregnation process, governed by the particle size and shape of the raw powder, the type and content of the binder, etc. It was also discovered that a rapid drying process was important in maintaining a positive replica shape. It was later shown that a second deposition of powder slurry on a previously sintered porous material followed by a second sintering resulted in increased density (80%) and improved compressive strength (36 MPa), as a result of the correction of flaws in the titanium struts created during the first cycle [318].

5.3.14. Multilayer Porous Titanium

Although open pore structures have the advantages of low elastic modulus and bone ingrowth, leading to better fixation with bone, they have the disadvantage of insufficient mechanical strength compared with that of bulk structural materials [319]. As a result, the majority of these porous metals are applied as coatings on fully dense substrates [320].

Imwinkelried [321] reported the first application of a porous titanium device for human lumbar spine. However, the yield strength of this porous titanium with porosities of 63%–50% was 68–140 MPa under compression. This low mechanical property limited its application to the load-bearing devices for human lumbar spine. Therefore, porous metal fabrication technology ensuring adequate strength and tailorable open pore structure is assumed to be very important for orthopaedic implants. To achieve these contradictory requirements of adequate strength and high porosity with adequately open pore structure, some researchers developed novel porous titanium by stacking porous material sheets of different porosities layer by layer to create multilayer porous titanium (Figure 36). The resulting multilayer porous titanium exhibited an elastic modulus of 11–12 GPa and yield strength of 150–240 MPa in compression tests with porosities of 80%–17%.

5.3.15. Rapid Prototyping

Rapid prototyping (RP) techniques, combining computer-aided design (CAD) with computer-aided manufacturing (CAM), are considered as a viable alternative for achieving extensive and detailed control over porous architecture [323,324], making it possible to build objects with predefined microstructure and macrostructure [325] and controlled hierarchical structures [326]. The imperfection of conventional techniques has encouraged the use of RP technologies [310]. Since 1980s, RP technologies have emerged as a revolutionary manufacturing process with inherent capability to rapidly make objects in virtually any shape. Until now, RP developments mainly focused on polymer and ceramic materials [309,327]. The transfer of RP technologies to metal materials possesses a significant challenge, because there are few investigators on making metal scaffold for orthopedic and tissue engineering application by RP techniques [328]. A key requirement for RP technologies is tight control over the scaffolds’ pore structure, including pore size, shape, volume and interconnectivity [329]. This enables an accurate assessment of the level of precision that the fabrication process can deliver.

Two methods were applied to make metal scaffold. One is indirect to make porous scaffolds by invest casting melt metal or metal powder slurry into a mold where the mold was made using RP [328]. Another is developing porous titanium scaffolds for tissue engineering using direct metal deposition [330]. Using investment casting technique, Lopez et al. [49] created a porous titanium scaffold with 100% interconnected structure and porosity around 60%. Hollander et al. [214] used a technique called direct laser forming to create porous Ti-6Al-4V alloy. Direct laser metal sintering proved to be an efficient way constructing dental implants with a functionally graded material [331]. Garrett et al. [332] adopted a multi-stage RP technique combining RP and powder metallurgy techniques, and successfully produced porous titanium scaffolds with fully interconnected pore networks and reproducible porosity and pore size. The scaffolds’ porous characteristics were governed by a sacrificial wax template, fabricated using a commercial three-dimensional (3D) printer. Powder metallurgy processes were employed to generate the titanium scaffolds by filling the wax template with titanium slurry. Scaffolds with porosities of 66.8%–3.6% revealed compression strength of 104.4–22.5 MPa in the axial direction and 23.5–9.6 MPa in the transverse direction, demonstrating their anisotropic nature and potential ability employed in orthopaedic applications. By altering the wax design template, pore sizes ranging from 200 to 400 μm were achieved. 3D reconstruction enabled the main architectural parameters such as pore size, interconnecting porosity, level of anisotropy and level of structural disorder to be determined.

3D fiber deposition (3DF), also a RP technology, was successfully directly applied to produce novel 3D porous Ti-6Al-4V scaffolds with fully interconnected porous networks and highly controllable porosity and pore size [53] and thus meet the requirements of tissue engineering. Through a series of experiments, a nozzle diameter of 0.4 mm, a powder concentration of 66%, a fiber depositing pressure of 2.5 bar, a feeding speed of 350 mm/min, and an initial height of 0.25 mm, were found to be optimal parameters for fabrication of Ti-6Al-4V scaffolds. The scaffolds have a good attachment between layers, and a fully interconnected porous structure.

Selective laser sintering (SLS), as one of the RP technologies, can individually design and rapidly manufacture components with any complex geometry from powder materials based on CAD model [333]. In recent years, SLS technique has been explored to fabricate tissue engineering scaffolds and porous implants consisted mainly of biocompatible polymers and polymer/ceramic composites [334,335]. However, there is little information available on preparation of porous Ti-based alloys by SLS. Shishkovsky et al. [336] reported an original method to obtain porous NiTi alloy by a combination of SLS and SHS processes. Porous Ti-10Mo alloy was prepared from a mixture of elemental metal powders and polymer binders by indirect SLS [210]. When the sintering temperature increased to 1100 °C in a vacuum condition, the porosity and pore size were 63% and 178 μm, respectively. The alloy comprised major α and minor β phases and meanwhile the alloy exhibited compressive yield strength of 95.59 MPa and an elastic modulus of 4.89 GPa at room temperature. SLM process was chosen to produce porous titanium with a particular periodic internal architecture. A porosity of 53% and pore size in the range of 860–1500 μm were finally adopted.

Selective electron beam melting (SEBM) was successfully used to fabricate novel porous Ti-6Al-4V structures with well-defined cellular structures for orthopaedic applications [55]. Firstly, a 3D CAD model is sliced into layers with constant thicknesses to provide layer information. The SEBM process starts with homogeneous application of a layer of metal powder on a process platform. After a preheating step, an electron beam scans the powder layer and creates a cross section of the part by fusing the loosely joined powder particles. Subsequently, the process platform is lowered by the thickness of one layer, and a new powder layer is applied and the process is repeated until the whole part has been built (Figure 37). The absorption of atmospheric gases should be effectively prevented because this might lead to a reduction in the ductility of titanium. Due to the layer-by-layer generation process, structures exhibited an interconnected porosity. Measurements gave a porosity of 80.5% and a mean pore size of 1.23 mm for the diamond structure, and 61.3% and 0.45 mm for the hatched structure, respectively, which is one of the most important requirements for tissue ingrowth. Thus, the structures exhibited porosities comparable to that of trabecular bone.

Laser engineered net shaping (LENS) is a solid freeform fabrication process that is primarily used to build near-net-shape metallic parts. Being a CAD and layer-based manufacturing process, LENS gives a significant advantage over conventional methods in terms of controlling the shape, size and internal architecture, particularly of porous structures. The schematic representation of the LENS process is shown in Figure 38a. First, a 3D model of a component is generated using CAD. Subsequently, a computer program slices the model into a number of horizontal cross sections or layers. These cross sections are sequentially created on a substrate, producing a 3D object. More detailed description of the process is provided in Reference [221]. Previous work involving LENS has been focused on net shape manufacturing [337], alloy development [338], gradient structures [339,340] and coatings [341]. Recently, this process has also been used to fabricate net shape porous Ti-based alloys [221] with superior mechanical, physical, and biological properties.

As a CAD driven process, LENS opens the way for mass customization of metallic parts, such as bone implants. LENS produced titanium parts have a natural rough surface with a macroporous structure, which also eliminates the need for applying a thermally sprayed titanium coating on implants. Moreover, this technique allows us to control porosity, pore size and distribution to meet a variety of different application needs. Bandyopadhyay et al. [342] have fabricated porous Ti-6Al-4V alloy structures (Figure 38 b), and demonstrated that the advanced manufacturing technique can be used to fabricate low modulus and tailored porosity implants with a wide variety of alloys. The effective elastic modulus of Ti-6Al-4V alloy has been tailored between 7 and 60 GPa, and porous Ti-based alloys containing 23%–32% porosity showed elastic modulus equivalent to human cortical bone. Also, Bandyopadhyay [212] used this technology to successfully fabricate 3D interconnected porous titanium structures with controlled porosity and pore size for hard tissue replacement. Meanwhile, the porous titanium exhibited a high mechanical strength between 24 and 463 MPa and an elastic modulus between 2.6 and 44 GPa, which are better matched properties with bone tissue compared with dense titanium.

Notwithstanding the advancements have been attained in porous titanium and its alloy scaffolds fabrication, the control over scaffold architecture using conventional sintering is highly process dependent. Over the past few years, novel direct fabrication processes of metallic components such as direct laser metal sintering [331], using metal powders to create functional parts with gradient porosity and cellular structures, can overcome the shortcomings of conventional methods. However, there are disadvantages such as high cost and low efficiency.

5.3.16. Advantages and Disadvantages of Various Preparation Methods

Many techniques have been applied to produce porous titanium and its alloy implants in recent years as above mentioned, but almost all the fabrication routes suffer from some problems stated to a certain extent.

Formation of secondary intermetallic phases, irregular pore shapes, inappropriate pore size, inhomogeneous pore distribution and the contamination during conventional powder metallurgy processing techniques, such as SHS, HIP, conventional sintering, SPS, MIM and space holder technique, limit their use. For example, when elemental powders are used in any of the powder metallurgy methods, undesirable nonequiatomic Ti-Ni phases leading to brittleness and inferior superelasticity often remain in the produced porous material. All these powder metallurgy processing techniques, especially MIM and space holder method, cannot avoid contamination during processing due to decomposition of binders and space filling materials which are mainly polymeric based or metallic salts. Although the combustion synthesis process can produce high purity porous materials, the complex shaped porous implants cannot be fabricated by the above-mentioned traditional methods. In addition, titanium and its alloys require very strictly controlled furnace atmospheres and they are extremely prone to react to interstitial elements such as O, C, N and H. As a result, contamination and more or less decrease in mechanical features are almost inevitable for all of the production routes. For example, through space holder method, homogenous pore structure can be produced with greater porosity, but nonmetallic inclusion tends to be introduced into the powders because of large quantities of space holder additions. Although higher pore sizes can be reached with the space holder technique, mechanical properties obtained may be lower than those of human bone [234]. In comparison, porous materials fabricated by powder sintering show high purity, but pore sizes and shapes are mainly predetermined by the raw powder characteristics, which makes pore structure difficult to control, while sintering of compacted loose powder normally results in low porosities. Meanwhile, fabrication technique of porous titanium and its alloys based on prealloyed starting powder generally contains lower levels of biocompatible elements and is expensive when compared to the use of elemental powders. Moreover, a common limitation of traditional processes is the inability of individually designing and manufacturing customized prostheses with complex and 3D framework. FC process mostly leads to randomly distributed pores whereas RP and laser processing are adopted when a controlled architecture is required. In both cases, the control of elastic properties anisotropy seems to be difficult.

A variety of powder metallurgy approaches have been developed to fabricate titanium products with tailored porosity. Wen et al. [178] used space holder method to create products with a large portion of open pores in an interconnected structure, which are created by the space holders. However, it is limited by high cost of molding dies, the limited complexity of compact shape and a variable density gradient. Chen et al. [285] employed metal injection molding, mixing titanium powders with space holders, to produce porous titanium compacts. MIM can be used to fabricate complex shapes, but the cost of molds and the problem of variable density gradient are important issues. In addition, a relatively large amount of binder is required for MIM, which raises concerns about binder residuals after the debinding stage, especially in the case of titanium.

A modified powder metallurgy technique called sponge replication [52] or impregnating method [343] has been used in fabrication of porous titanium and its alloys, by which pore size, shape, and distribution can be well adjusted in a large scale. Some interconnected porous structures that perfectly mimic spatial morphology of cancellous bone can be easily reproduced by this method. However, the major drawback is the difficulty to avoid contamination and impurity phases in the titanium materials. In addition, sintered titanium struts likely contain cracks and other metallographic defects, and thus porous titanium materials cannot bear tensile stress and even exhibit brittle nature. This limitation implies a risk of the orthopedic applications of the porous titanium in vivo where a tensile or impact load may be applied on the implant.

In order to improve toughness and tensile strength of porous titanium, a novel porous titanium material called entangled wire material [314] has been developed. Its prominent advantage is that strong metal wires acting as struts construct spatial porous structure, so that the drawbacks of as-sintered/as-solidified struts (e.g., metallurgical defects, contaminations, etc.) can be completely overcome. Compared with brittle nature of conventional porous titanium, entangled titanium wire porous material is very tough and exhibits excellent superelasticity. Its superiority in mechanical properties supposes the potentials for orthopedic applications, such as hip joint implant and dental implant, etc. However, elastic modulus is too low due to structural flexibility of entangled titanium wire materials and is not suitable for load-bearing biomedical applications. The relationship between mesh structure and mechanical properties has not been established. Furthermore, basic aspects in terms of fabrication techniques (especially for the 3D porous titanium wire materials), structure characterization, and mechanical properties tailoring, are still open for investigation.

In recent years, RP technique has been utilized to fabricate porous titanium and its alloys by using laser power to sinter/remelt titanium powders [214]. It is of great benefit to accurately control pore shape, size and distribution by beforehand design and programming, so that an ideal porous architecture can be constructed. With alternative fabrication approaches, this technique has been developed into direct laser sintering [331] and selective laser melting [221,344]. The RP technique provides good feasibility and excellent flexibility for production of custom built metallic implants with complicated structures. Another advantage of RP is the case sensitive manufacturing according to the needs of specific patient. However, porous titanium production also suffers from impurities, inclusions and other defects associated with laser sintering or melting processing, which significantly deteriorate mechanical properties such as tensile stress resistance and toughness. In the view of physical metallurgy, the struts in porous titanium with as-sintered (laser sintering) or as-solidified (laser melting) microstructure usually exhibits much lower ductility than that of the wrought or plastically deformed counterparts. It is unlikely that laser sintered/melted porous titanium implants can stand against tensile and impact stress, though there are very few data published [214,221,331,344]. Apart from RP, a similar 3D deposition approach was also used to fabricate porous titanium, in which deposited 3D porous structure was strengthened by drying in air and sintering in high vacuum [53]. This approach does not need laser source and control system involved. Despite the fact that a good 3D mesh structure was constructed, its low tensile strength and low ductility can be expected due to the similar above reasons. Besides, this method is often time consuming because of the need to use pre-designed wax templates, and hence it is difficult to employ RP in mass production.

Therefore, there are still problems to be solved in current techniques for fabrication of porous titanium and its alloys for biomedical applications: the difficulty to create controlled pore characteristics such as porosity, pore size and distribution, etc.; contamination, presence of impurity phases, limited and predetermined part geometries; the insufficient knowledge of porous structure-property relationships; the requirements of new sintering techniques with rapid energy transfer, and less energy consumption and so on. The growing demand for individualized treatment and excellent properties of a porous metallic implant requires the use of novel methods to fabricate near-net-shape implants, which likely could be achieved by combining the above-mentioned approaches. Fabrication methods for porous titanium and its alloys that can ensure uniform pore size, shape and distribution, mechanical properties and high levels of purity are in high demand [232]. Conventionally sintered porous titanium and its alloys are often very brittle and are prone to crack propagation at low stresses and usually suffer from loss of physical properties due to stress concentrations at porous interface, microstructures changes, and surface contamination from high temperature sintering process [51,178,232]. Influences of microstructural configurations on performance of porous titanium and its alloys will be discussed in the next section.