Processing and Development of Porous Titanium for Biomedical Applications: A Comprehensive Review

Abstract

Titanium (Ti) and its alloys are widely used in orthopedic applications, including total hip and knee replacements, bone plates, and dental implants, because of their superior biocompatibility, bioactivity, corrosion resistance, and mechanical robustness. These alloys effectively overcome several limitations of conventional metallic implants, such as 316L stainless steel and Co-Cr alloys, particularly with respect to corrosion, fatigue performance, and biological response. However, dense Ti alloys possess a relatively high elastic modulus, which can cause stress shielding in load-bearing applications. This challenge has motivated significant research toward engineered porous Ti structures that exhibit a reduced and bone-matched modulus while preserving adequate mechanical integrity. This review provides a comprehensive examination of powder metallurgy and additive manufacturing approaches used to fabricate porous Ti and Ti-alloy scaffolds, including additive manufacturing and different powder metallurgy techniques. Processing routes are compared in terms of achievable porosity, pore size distribution, microstructural evolution, mechanical properties, and biological outcomes, with emphasis on the relationship between processing parameters, pore architecture, and functional performance. The reported findings indicate that optimized powder-metallurgy techniques can generate interconnected pores in the 100–500 μm range suitable for osseointegration while maintaining compressive strengths of 50–300 MPa, whereas additive manufacturing enables the precise control of hierarchical architectures but requires careful post-processing to remove adhered powder, stabilize microstructures, and ensure corrosion and wear resistance. In addition, this review integrates fundamental aspects of bone biology and bone implant interaction to contextualize the functional requirements of porous Ti scaffolds.

1. Introduction

The demand for orthopedic implants is experiencing an unprecedented surge, primarily due to the increased prevalence of conditions such as osteoporosis, which weakens bones, and osteoarthritis, which causes joint inflammation [1]. Human bone joints are most likely to be affected by degenerative and inflammatory diseases [1]. Degenerative diseases impair the mechanical strength of bones either through excessive stress or due to a lack of the body’s natural healing mechanisms. Studies suggest that nearly 90% of individuals above the age of 40 are affected to some extent by degenerative joint disorders [2]. The replacement of diseased joint surfaces using metallic, ceramic, or polymeric material through arthroplasty surgery is the ultimate solution to this problem [3]. Arthroplasty is a surgical procedure where damaged natural joint surfaces are replaced with specialized implant materials to alleviate pain and improve mobility by creating a new prosthetic joint. Bones comprise four cell populations crucial for growth and resorption: osteoblasts, osteoclasts, osteogenic cells, and osteocytes, all originating from bone marrow. Although bones possess the capacity for physiological remodeling and self-repair, they are unable to effectively cope with severe injuries such as critical-sized defects (CSDs) [4]. CSDs are defects that will not heal on their own during the natural lifespan of the individual [5], and typically, defects in the range of 2 cm to 5 cm are considered critical in humans [6,7]. The standard approaches for treating CSDs include autograft (bone harvested from the same patients), allograft (bone harvested from other patients), and xenograft (bone sourced from other species) [8,9]. While these approaches often yield positive results, they come with specific limitations.

The application of autograft is not suitable for large-sized defects such as spinal arthrodesis [10], while allograft poses risks of immunologic rejection, donor dependency, poor osteogenesis, and a higher resorption rate [11]. Although xenografts present a promising solution to organ shortages, their clinical use is restricted by immune barriers, zoonotic infection risks, ethical concerns, and physiological challenges [12,13,14]. Such challenges and constraints drive researchers to actively explore and create dependable bone substitutes that closely mimic the characteristics of natural human bone. The field of biomaterials research began with the inaugural biomaterial conference held at Clemson University in South Carolina in 1969 [15]. Extensive research has led to the development of various metallic [16], ceramic [17,18], and polymeric [19] materials suitable for mimicking the properties of a healthy human bone. These materials belong to different classes, each possessing distinct physical, mechanical, chemical, and biological properties. For instance, metallic alloys, known for their robust mechanical properties, are commonly used in load-bearing joint prostheses [20,21,22,23]. Ceramic materials, renowned for their wear resistance and bioactive properties, are often used to coat metallic implants [24,25,26,27,28,29]. Polymers, being biodegradable and harmless to the body, are utilized in bone repair applications [30].

Scientists attribute the primary cause of failure in metallic implants to the ‘stress shielding effect’ [31,32,33]. This effect occurs due to the mismatch in mechanical properties like Young’s modulus, yield strength, ultimate tensile strength, fracture toughness, fatigue strength, and ductility (including elongation and toughness under cyclic loading) between natural bone and metallic alloys under in vivo conditions, leading to the resorption of bone tissues surrounding the implant material [34]. Moreover, metallic implants are limited by issues such as poor wear and corrosion resistance, as well as insufficient biocompatibility [35,36,37,38,39,40]. Traditional metallic implants, such as 316L stainless steel and Co-Cr alloys, particularly suffer from localized corrosion, ion release, and inadequate long-term biological compatibility. Although titanium (Ti) and its alloys are also metallic materials, they offer significantly improved corrosion resistance, superior biocompatibility, and better mechanical performance compared to these earlier metallic systems. Owing to their high durability and excellent biological response, Ti-based materials have become the focus of implant development; however, they still present challenges such as a relatively high elastic modulus in dense form, which must be addressed through alloy design and porosity engineering. The commercial production of titanium began with the invention of the Kroll process in 1946 [41], followed shortly by the introduction of titanium alloy–based implants within a year [42]. Since its inception, titanium and its alloys have found extensive use in biomedical applications such as spinal fusion, skeletal repair, and dental implants [15,43,44,45,46,47]. Titanium and its alloys have demonstrated superiority over medical-grade stainless steel, cobalt–chromium, and magnesium-based alloys concerning biocompatibility, mechanical strength, and corrosion resistance. However, their higher Young’s modulus (~110 GPa) initially limited their biomedical applications. The development of second-generation titanium alloys, specifically β alloys, aimed to reduce this modulus. Titanium alloys with a low elastic modulus, such as Ti-Nb-Zr-T, have been specifically developed to closely align with the modulus of different types of human bone [48,49,50,51,52].

The concept of porous biomaterials for osseointegration was introduced by Weber and White in 1972 [53], leading to extensive research on porous implant materials using various processing methods [54,55,56,57]. Porous scaffolds, characterized by a lower effective Young’s modulus than conventional titanium implants, offer several advantages. These materials enhance biological fixation by encouraging bone tissue to grow into the implant’s porous structure, enabling even stress distribution between the implant and surrounding bone. Moreover, porous scaffolds enhance colonization and subsequent substitution by biological cells [58,59]. Numerous manufacturing methods have been developed to produce porous structures for implant applications. Palka et al. [60] reviewed porous titanium implants, emphasizing various types of porosity in their structure, while Koju et al. [61] examined porous Ti6Al4V bone implants produced by additive manufacturing. However, a systematic review encompassing all production methods to fabricate porous implants is still not available. Hence, the primary aim of this review is to survey all existing methods to fabricate porous titanium through powder metallurgy and to provide a comprehensive overview of the research status of titanium and its alloys targeted for orthopedic applications. This article is divided into nine sections with several sub-sections, focusing primarily on the various techniques employed to develop porous titanium scaffolds for orthopedic applications using powder metallurgy.

2. Material Characteristics Critical to Implant Functionality

“As defined by Williams in 1987, biomaterials are non-living materials designed for use in medical devices that interact with biological systems”. Biomaterials are artificial or natural materials that restore the integrity and functionality of damaged or diseased biological structures. For an implant to achieve longevity and high compatibility with the body, the biomaterial must possess certain properties suitable for biomedical applications. These properties include biocompatibility, osseointegration, mechanical strength, wear resistance, and corrosion resistance. Figure 1 outlines the fundamental requirements that must be considered in the design and development of implants.

2.1. Mechanical Properties

The primary function of bone is to provide structural support under mechanical loading and protect vital organs. In cases of trauma or diseased biological structures, implants often replace the function of the affected organ. Currently, most implants used in orthopedic traumas and arthroplasty are of metallic composition due to their excellent mechanical strength [62]; therefore, the mechanical analysis of metallic implants is of prime concern. Essential properties include Young’s modulus, fracture toughness, hardness, yield strength, ultimate tensile strength, stiffness, ductility, time-dependent deformation, and creep. Wolff’s law of bone remodeling (1892) states that mechanical load can affect bone architecture [63] and hence promote bone remodeling and fracture healing [64,65]. Fatigue strength is also critical, as it affects material performance under cyclic loads, influencing the long-term success of implants [15]. Young’s modulus is crucial for the success of implants. A significant mismatch between the modulus of bone and the implant material can negatively impact load transfer from the implant to the bone and within the bone, leading to bone resorption and implant loosening. This biomechanical incompatibility is the stress-shielding effect [15,66,67].

Table 1. Physical and mechanical properties of metallic, ceramic, and polymeric biomaterials (ρ—density in g/cm3; E—Young’s modulus in GPa; YS—yield strength in MPa; UTS—ultimate tensile strength in MPa; UCS—Ultimate compressive strength in MPa; and FS—fatigue strength in MPa, 107 cycles).

Material	ρ	E	YS	UTS	UCS	FS	Ref.
Natural Bone
Cortical bone	1.8–2.0	7–30	-	164–240	100–230	27–35	[68]
Cancellous bone	1.0–1.4	0.01–3.0	-	-	2–12	-
Metals and Alloys
Ti-6Al-4V (cast)	4.43	114	760–880	895–930	-	600–700	[68]
Ti-6Al-4V (wrought)	4.43	114	827–1103	860–965	896–1172	500–800
Ti-6Al-7Nb	4.52	105	880	900	-	-	[15]
SS316L	8.0	193	170–310	540–1000	480–620	240–480	[68]
Fe20Mn	7.73	207	420	700	-	-	[69]
Zn-Al-Cu	5.79	90	171	210	-	-	[34]
Co-Cr-Mo	8.3	240	500–1500	900–1540	-	500–900
CoCr20Ni15Mo7	7.8	195–230	240–450	450–960	-	-
Pure Mg (cast)	1.74	41	21	87	40	-
Pure Mg (wrought)	1.74	41	100	180	100–140	-
AZ31 (Mg-based alloy)	1.78	45	185	263	-	-
AZ91 (Mg-based alloy)	1.81	45	160	150	-	-
Ceramics
Alumina Ceramics	4	260–410	-	400–580	-	-	[34]
Synthetic hydroxyapatite	3.15	6–102	-	-	0.22–4.1	-	[18,70]
Zirconia	3.98	210	-	800–1500	1990	-	[71]
Polymers
PLGA	1.2–1.3	1.69	3.8–26.6	13.9–16.7	-	-	[34]
PCL	1.15	281–686	8.37–14.66	68–103	-	-
PLA	1.8	3750	70	59	-	-

lists the mechanical properties of materials used for orthopedic implant applications and human bone.

2.2. Corrosion and Wear Resistance

Corrosion and wear are inevitable problems associated with orthopedic implants [72]. When a metallic implant is introduced into the body, it encounters biological fluid containing various cations and anions, leading to electrochemical corrosion on the implant surface. Four types of corrosion are typically observed in orthopedic implants: galvanic, pitting crevice, and fretting corrosion. Galvanic corrosion occurs due to the electrochemical potential difference between two different or the same metal surfaces when introduced in a biological fluid [73,74,75]. Studies showed that when Ti-based and Co-based alloys are coupled, the corrosion rate was observed to be as low as 0.02 μA/cm², and no instances of corrosion were found on the metallic interface [76]. Pitting corrosion is a localized form of corrosion that creates cavities in the material [77,78,79]. In the case of metallic implants, chloride ion breakdowns and the protective passive oxide films result in the formation of pits at the site [80,81]. Crevice corrosion is like pitting corrosion in terms of propagation mechanism but differs in the initiation mechanism [82,83,84]. Crevice corrosion generally occurs in confined spaces with low oxygen tension and high chloride concentration, destroying the passivation layer [85]. The mechanism of fretting corrosion differs from other types of corrosion. In fretting corrosion, the passivation layer is disrupted due to micro-motion between parts of an implant [86,87,88]. Continuous relative motion between two surfaces under load generates wear debris around the implant-bone interface. Research indicates that inadequate wear resistance can result in implant loosening, while the accumulation of these wear particles may provoke negative tissue responses [89]. In natural joints, synovial fluid functions as an effective lubricant, providing boundary and mixed lubrication regimes that significantly reduce wear [90]. However, the stability of this lubricant film depends strongly on surface chemistry, roughness, and pore geometry, which regulate protein adsorption and fluid retention. Poorly designed surface features or trapped debris can disrupt the synovial film and accelerate third-body abrasion. Therefore, implant materials and surface designs must ensure adequate corrosion and wear resistance while also supporting stable synovial lubrication to prevent premature mechanical failure.

Corrosion in metallic implants is an electrochemical degradation process that occurs when the metal interacts with physiological fluids. As described in the review article by Eliaz et al. [91], the human body behaves as an electrolyte containing chloride ions, proteins, and dissolved oxygen, which facilitates metal dissolution through anodic and cathodic reactions. The breakdown of the passive oxide film, especially under mechanical loading or fretting conditions, results in the release of metallic ions and particulate debris into surrounding tissues. This degradation can lead to local inflammatory responses, hypersensitivity, cytotoxic effects, and ultimately implant loosening or failure. The article emphasizes that corrosion is strongly influenced by alloy composition, microstructure, mechanical wear, and the stability of the protective oxide layer on the implant surface. Therefore, selecting materials with high corrosion resistance, such as titanium and its alloys, and optimizing surface treatments are essential for minimizing clinical complications associated with corrosion-induced degradation in biomedical implants [91].

The corrosion mechanism inside the human body can be easily understood by taking the example of Fe and Ti in biological environments. In physiological (aqueous) environments, the corrosion of metallic implants is propelled by coupled electrochemical half-reactions: metal dissolution (anodic) and reduction (cathodic) at the metal/electrolyte interface. For titanium alloys (e.g., Ti-6Al-4V), corrosion begins with the oxidation of titanium atoms, producing Ti⁴⁺ ions that hydrolyze to form a stable titanium dioxide (TiO₂) passive film. This passivation layer strongly inhibits further metal dissolution under steady-state conditions, but local disruption (due to mechanical stress, chloride, or reactive oxygen species) can lead to breakdown and renewed dissolution. Studies show that inflammation (e.g., H₂O₂) and cathodic activation can degrade the oxide film and drive selective dissolution of phases, particularly the β-phase in Ti-6Al-4V. The relevant electrochemical reactions can be expressed as follows:

For titanium alloys, we have the following:

Anodic Reactions (Oxidation)

$$\mathrm{T}\mathrm{i}\to {\mathrm{T}\mathrm{i}}^{4+}+4{\mathrm{e}}^{-}$$

(1)

Passive Film Formation

$${\mathrm{T}\mathrm{i}}^{4+}+2{\mathrm{H}}_2\mathrm{O}\to {\mathrm{T}\mathrm{i}\mathrm{O}}_2+4\mathrm{H}$$

(2)

This is consistent with the mechanism of oxide-film stabilization seen in studies of Ti-6Al-4V [92].

Cathodic Reaction (Reduction)

$${\mathrm{O}}_2+2{\mathrm{H}}_2\mathrm{O}+4{\mathrm{e}}^{-}\to 4{\mathrm{O}\mathrm{H}}^{-}$$

(3)

This is the dominant cathodic process in neutral, oxygenated physiological solutions [93].

For Fe-based alloys, corrosion typically proceeds via continuous dissolution due to the relatively less stable protective films in chloride-rich physiological environments.

The primary anodic reactions are as follows:

$$\mathrm{F}\mathrm{e}\to {\mathrm{F}\mathrm{e}}^{2+}+2{\mathrm{e}}^{-}$$

(4)

$${\mathrm{F}\mathrm{e}}^{2+}\to {\mathrm{F}\mathrm{e}}^{3+}+{\mathrm{e}}^{-}$$

(5)

Meanwhile, the cathodic reaction is generally the same oxygen-reduction process:

$${\mathrm{O}}_2+2{\mathrm{H}}_2\mathrm{O}+4{\mathrm{e}}^{-}\to 4{\mathrm{O}\mathrm{H}}^{-}$$

(6)

These equations and their relevance to in vivo degradation have been documented in the context of biodegradable metal implants [93,94].

2.3. Porosity Effect

Porosity is crucial in developing effective implants for bone tissue engineering. The key factors include pore interconnectivity, porosity volume, and pore sizes. The porous structure must have interconnected pores of appropriate size [95,96]. Interconnected pores are vital as they facilitate cell infiltration, in vivo vascular formation, osteogenesis, cell proliferation, oxygen and nutrient flow, and waste removal [97,98]. It is also reviewed that interconnectivity provides a way for the ingrowth of blood vessels to facilitate multiple vessel formation [99]. Earlier studies suggest that a minimum interconnection size of over 50 µm is suitable for bone ingrowth [100]. The porosity of the implant enhances bone integration because the porous surface provides an interlocking medium to the surrounding tissue and implant, resulting in good biomechanical compatibility and high resistance to fatigue loading [101,102]. It has been reported that pore sizes ranging between 100 µm and 200 µm are suitable for osseointegration [15].

It is important to note that any increase in the implant’s porosity content decreases the mechanical properties of the system [103]. Therefore, without compromising the mechanical properties, the porosity of the system should be optimized, and the pore size should be strictly controlled. Different techniques for adding space holder materials of defined size, densification of the green compacts, sintering conditions, etc., can be employed to incorporate pores in the final structure [104]. A detailed discussion of the production of porous structures using different techniques will be discussed in subsequent sections.

2.4. Surface Wettability and Its Role in Implant Performance

Surface wettability is a critical design parameter for metallic implants, directly influencing both their biological and tribological performance. Wettability governs the interaction of synovial fluid, proteins, and cells with the implant surface, thereby affecting lubrication, friction, wear, and early-stage osseointegration. Hydrophilic surfaces generally facilitate more favorable biological interactions compared to hydrophobic surfaces, largely because wettability affects the first biomolecular events at the implant interface. Surface wettability is known to regulate four major aspects of host response: (i) adsorption and conformation of proteins and macromolecules that form the initial conditioning layer; (ii) adhesion, spreading, and differentiation of hard and soft tissue cells; (iii) bacterial adhesion and biofilm formation; and (iv) the overall rate and quality of in vivo osseointegration [105]. These interactions begin within milliseconds of implantation and significantly depend on the hydrophilic or hydrophobic nature of Ti surfaces.

Recent works have demonstrated that micro- and nano-scale surface textures strongly influence the spreading characteristics and retention of lubricants and biological fluids on Ti alloys. Several studies [106,107] highlight that multiscale fractal roughness enhances dynamic wetting, fluid transport, and boundary lubrication, thereby reducing wear at articulating implant interfaces. These findings emphasize that wettability is not only important for osseointegration but also for the flow behavior of synovial fluid, which plays a major role in limiting wear during joint articulation. Insights from Gittens et al. [108,109,110] further confirm that hydrophilic surfaces support improved protein adsorption in favorable conformations, enhance osteoblast maturation, and promote faster bone–implant contact compared to hydrophobic surfaces. Hydrophilic and superhydrophilic surfaces also reduce hydrocarbon contamination and strengthen the biochemical environment for cell attachment, proliferation, and differentiation [111]. Conversely, hydrophobic surfaces may denature adsorbed proteins, inhibit cell adhesion, and alter the biological cascade leading to integration. Overall, incorporating appropriate surface wettability, often achieved through micro/nano-texturing, plasma modification, chemical treatments, or photocatalytic activation, provides a significant advantage for titanium implants [112]. It enhances lubrication and wear performance in load-bearing joints and concurrently improves osseointegration, offering a dual benefit highly relevant to porous titanium structures discussed in this review.

2.5. Biocompatibility

The biocompatibility of any material is the basis of understanding the host response to implants and is defined diversely by different researchers and regulatory bodies. According to the US Food and Drug Administration, biocompatibility is defined as the effect that the materials induce no measurable harm to the host [113]. It is observed that improper biocompatibility leads to dysesthesia (loss of sense), discomfort, pain, infection, resorption of bone [114], etc. The ions released from the metallic implant may induce hypersensitivity and lead to implant failure [115]. Williams (1987) [116] defines biocompatibility as a material’s capacity to generate a suitable response from the host in a given medical application. This implies that a biocompatible material must exhibit both bioactivity and bifunctionality.

The biocompatibility of any material can be defined as the ‘ability of a material to perform with an appropriate host response in a specific application’, which means that a biocompatible material should be bioactive and biofunctional. Similarly, Retner et al. [117] defined that no material can be biocompatible if it leaches cytotoxic substances when implanted. The success of biomaterials largely depends on the body’s reaction to the implant, which measures the material’s biocompatibility. In other words, an implant material is considered biocompatible if it elicits a positive response when exposed to a biological environment [118]. Biocompatibility is primarily determined by two key factors: the biological response the material triggers in the host and the extent to which the material degrades within the bodily environment [15]. Therefore, to achieve good biocompatibility, non-toxic alloying elements with excellent corrosion and wear resistance should be chosen during the design and development of metallic implants.

As discussed in Section 2.2, corrosion can directly affect the overall success rate of the implant, as biocompatibility and corrosion are intrinsically linked. As the degradation of metallic implants in vivo directly influences the host response. A biocompatible material must not trigger thrombosis, toxic or allergic reactions, adverse immune responses, or destruction of cellular components, and it should maintain its structural integrity without releasing harmful ions into surrounding tissues [91]. However, corrosion alters the implant surface chemistry and accelerates metal-ion release, which can destabilize local biological equilibrium and initiate inflammation or hypersensitivity. As corrosion progresses, ions such as Ni, Cr, Co, or V can accumulate and provoke cytotoxic or allergenic responses, sometimes even at low corrosion rates that do not compromise mechanical performance [91].

Body fluid composition further governs this relationship. Chloride ions promote passive-film breakdown, fluctuating pH during inflammation (as low as 4.0) increases localized corrosion, and reduced dissolved oxygen levels hinder repassivation, collectively intensifying degradation and influencing biocompatibility outcomes [119]. Proteins and macromolecules in extracellular fluids can bind metal ions and transport them away from the surface, destabilizing the electrical double layer, or alternatively form an adsorbed barrier that slows corrosion; thus, protein–metal interactions directly affect degradation rates and biological responses. Cells adhering to implant surfaces may either protect the metal by forming a physical barrier or enhance corrosion through secretion of reactive oxygen species such as O²⁻ and H₂O₂₉, which accelerate metal dissolution and negatively impact surrounding tissue [91].

2.6. Osseointegration

According to the American Academy of Implant Dentistry (1986), osseointegration refers to the direct contact between normal remodeled bone and an implant without any intervening non-bone tissue, ensuring sustained load transfer and distribution from the implant to and within the bone tissue. Clinically, osseointegration can be defined as the process by which a clinically asymptomatic, rigid fixation of alloplastic materials is achieved and maintained in bone under functional loading [120]. Factors such as design, surface chemistry, roughness, chemical composition, and loading conditions should be considered before developing an implant for optimal osseointegration [121]. Figure 2 shows a schematic diagram illustrating the timeline of osseointegration of a metallic implant with respect to cellular events for 28 days [122].

3. Bone Metabolism

Bone is a rigid, living tissue that provides support during mechanical loading and protects vital organs inside the body [123]. When a baby is born, their body contains approximately 300 soft bones. These bones are primarily made of cartilage, which allows for flexibility and growth. As the baby matures, bones gradually fuse and harden through ossification. By adulthood, this fusion reduces the total number of bones to 206, providing a more rigid and supportive skeletal structure. To understand bone metabolism, it is necessary to grasp bone physiology and mechanical properties. The different structural levels of bone are shown in Figure 3, illustrating that the bone tissue comprises collagen fibers at the nanoscale.

3.1. Bone Physiology

At the microscopic level, bone structure can be classified into two types based on the arrangement of fibers: woven bone and lamellar bone [124]. The following sections will explore these two types in detail.

3.1.1. Woven Bone

Woven bone, considered an immature type of bone tissue, typically has a mineral grain size ranging from 10 to 15 nanometers. It is commonly located in the metaphyseal region and at sites of fracture healing, such as the callus [125]. These are coarse-fibred and collagen fibers and are randomly oriented throughout the structure. It can also be inferred that the direction-independent mechanical behavior of woven bone results from nonuniform collagen [126].

Figure 3. (a) Macro-to-microscopic picture of cancellous and cortical bones. The endosteum structure lines the spaces of cancellous bone, where bone marrow is stored. The periosteum barrier protects the cortical tissue that is tightly packed within osteons. Haversian canals, which comprise blood vessels and nerve tissue, surround concentric lamellae with thicknesses of around 3 μm to create osteons. Osteocytes live in the osteon’s lacuna structures. (b) At the nanometric scale, bone tissue comprises collagen fibers with 67 nm periodic spacing and 40 nm gaps for mineral components. Reproduced with permissions from [127].

Figure 3. (a) Macro-to-microscopic picture of cancellous and cortical bones. The endosteum structure lines the spaces of cancellous bone, where bone marrow is stored. The periosteum barrier protects the cortical tissue that is tightly packed within osteons. Haversian canals, which comprise blood vessels and nerve tissue, surround concentric lamellae with thicknesses of around 3 μm to create osteons. Osteocytes live in the osteon’s lacuna structures. (b) At the nanometric scale, bone tissue comprises collagen fibers with 67 nm periodic spacing and 40 nm gaps for mineral components. Reproduced with permissions from [127].

3.1.2. Lamellar Bone

Lamellar bone is another structure of bone that can be distinguished at the microscopic level. Unlike woven bone, it consists of a mineralized matrix known as hydroxyapatite, with the chemical formula Ca₁₀(PO₄)₆(OH)₂. In the human femur, hydroxyapatite crystals typically measure between 20 and 80 nm in length and from 2 to 5 nm in thickness [125]. As implied by their name, lamellar bones are structured in layers and feature collagen fibers aligned with stress directions, contributing to their anisotropic mechanical behavior [126].

Further, based on structural organization, woven and lamellar bones are organized into trabecular and cortical bones [125,128]. Trabecular bone is highly porous, having 50 to 90% porosity and large pores up to several millimeters in diameter. Therefore, these bones are also known as spongy bones or cancellous bones. The trabecular bone bears compressive forces under physiological loading conditions [126]. These bones are typically located in the metaphyseal and epiphyseal regions of both long and cuboidal bones. The distal end radius is an example of a trabecular bone. Similarly, cortical bones are less porous and possess small pores of size up to 1 mm in diameter. Therefore, it is also known as compact bone, contributing to 80% of the weight of the human skeleton. It is harder, stronger, and stiffer than trabecular bone due to less porosity. The humerus and femur are examples of cortical bone.

3.2. Chemical Composition of Bone

Bone is primarily composed of approximately 70% inorganic material, 20% organic content, and 10% water. The inorganic portion mainly consists of crystalline hydroxyapatite (HAp) Ca₁₀(PO₄)₆(OH)₂, while around 90% of the organic component is made up of Type I collagen. The remaining 10% includes non-collagenous proteins, lipids, and various other macromolecules [125,126]. HAp is the principal inorganic component of natural bone and has been widely studied for orthopedic and dental applications because of its exceptional biocompatibility and bioactivity. Its crystal structure is hexagonal, and its Ca/P molar ratio of 1.67 closely matches that of biological apatite, enabling strong chemical affinity with native bone tissue [17,18]. Synthetic HAp exhibits the ability to directly bond with bone through the formation of a biologically active apatite layer, which stimulates osteoblast adhesion, proliferation, and mineralization. In addition to its osteoconductive behavior, HAp provides good compressive strength and structural stability, although it remains brittle and exhibits relatively low fracture toughness, which limits its role as a standalone load-bearing material. Consequently, HAp is frequently used as a coating material to enhance osseointegration and improve the biological interface. Its thermal stability, ionic substitution capability (e.g., carbonate, magnesium, zinc), and similarity to natural bone mineral make HAp an essential material for engineered bone scaffolds and composite implant systems. Boskey and Coleman [129] suggested that the mineral component of bone changes with age. Various changes in bone mineral composition, which occur with age, are listed below [123]:

• An increase in overall mineral content;
• Greater carbonate substitution within the mineral structure;
• A reduction in acid phosphate substitution;
• Higher hydroxyl content;
• An elevated calcium-to-phosphorus (Ca/P) molar ratio;
• Growth in crystal size and improved crystallinity.

3.3. Types of Cells in the Bone

Bones comprise four types of cells: osteoblasts, osteoclasts, osteocytes, and osteogenic cells. Figure 4 provides a schematic diagram of these cell types [130].

3.3.1. Osteoblasts

These cells are also known as bone-forming cells because they synthesize new bones. Osteoid (a protein mixture secreted by osteoblasts) plays an important role in this process, as it mineralizes to form bone [130].

3.3.2. Osteocytes

Osteocytes are fully differentiated bone cells derived from osteoblasts. The primary distinction between these two cell types lies in their position within the bone structure: osteoblasts reside along the periosteal and endosteal surfaces, while osteocytes are embedded within the bone matrix, organized in concentric rings around the central canal of an osteon and situated between lamellae [130].

3.3.3. Osteoclast

Osteoclast cells are responsible for bone tissue breakdown and essential for maintaining, remodeling, and repairing the vertebral skeleton’s bones. The osteoclast secretes acid and collagenase through bone resorption, which molecularly breaks down and digests the mixture of hydrated proteins and minerals. This mechanism similarly regulates the blood calcium level. A bone with a physiological grain size of less than 100 nm in diameter is generally termed a healthy bone. A healthy bone is continuously remodeled throughout the life span by a process that involves the formation of a bone modeling unit and activating bone cells. Osteoclasts perform the phenomenon of bone resorption. The osteoclast is derived from pluripotent cells of the bone marrow. Pluripotent cells tend to discriminate against different cells, including monocytes and macrophages. These pluripotent cells resorb bone by forming disordered cell membrane edges and thus increase the surface area of attachment on the bone surface. The carbonic anhydrase system produces hydrogen ions, which decreases the pH of the local environment, and due to this, the solubility of hydroxyapatite (a major inorganic component of bone) increases. After this process, proteolytic digestion takes place to remove an organic component, and the removal of organic and inorganic components results in the formation of resorption pits [131].

3.3.4. Osteogenic Cells

Osteoprogenitor cells, also known as osteogenic cells, are stem cells found in bone tissue that play a vital role in bone formation and regeneration. Located within the bone marrow, they serve as precursors to more specialized bone cells, including osteoblasts and osteocytes. From immature mesenchymal cells, osteoprogenitor cells develop into spindle cells on the surface of adult bones. They are more common in growing bones, triggering multifunctional phases that reshape them. As we age, our body’s capacity to produce or use more osteoprogenitor cells declines [132].

4. Materials Used for Orthopedic Implant Applications: Advantages and Disadvantages

Materials such as metals, polymers, and ceramics are commonly used for implants. This review article focuses on an in-depth study of the development of Ti-based porous materials. Still, it also includes a brief section on other materials like ceramics and polymers used in implant applications. The selection of material depends on the application area. Table 2 briefly describes the advantages, disadvantages, and applications of various materials used as implants. Ceramics are inorganic oxides of non-metals known for their excellent corrosion resistance, biocompatibility, and bioactivity. They are primarily composed of ionic bonds with some covalent character, making them brittle with low fracture toughness and a high elastic modulus. There are two main types of bioceramics: bio-inert ceramics and bioresorbable ceramics. Bio-inert ceramics, such as zirconia and alumina, exhibit high chemical inertness, making them suitable for biological applications [133,134,135]. Bioresorbable ceramics, like hydroxyapatite and β-tricalcium phosphate, are designed to gradually dissolve and be replaced by natural bone [136,137,138]. However, due to their poor mechanical properties, pure ceramics are not ideal for weight-bearing prostheses in long bones, as they tend to fail under load-bearing conditions. However, ceramics are successfully used in bone filling and dentistry applications [139]. Similarly, polymers are organic non-metal oxides; they can be used as an implant material in low-load-bearing fracture sites. Polymers have emerged as promising materials for orthopedic use due to their mechanical properties, which are comparable to those of trabecular bone, and their biodegradability has further heightened interest in their application [140]. Polymer-like ultra-high-molecular-weight polyethylene (UHMWPE) has properties like high impact strength, low friction coefficient, and low density. These properties have made UHMWPE a popular choice for joint replacement. However, the application of UHMWPE is limited due to long-term radicals in the bulk resulting from the ionizing radiation employed in the sterilization process [141]. Natural polysaccharide polymers like starch, cellulose, and alginate are also used in biomedical applications [139,140]. The main problem associated with the use of polymers in orthopedic applications is the overproduction of wear debris, which leads to inflammatory reactions between adjacent tissue and implant. This adverse tissue reaction causes osteolysis, bone resorption, and implant failure [142].

Metals and their alloys are preferred over ceramics and polymers for implant applications due to their excellent mechanical properties and biocompatibility [143]. Commonly used metals include titanium and its alloys, medical-grade stainless steel, cobalt–chromium, and magnesium alloys. These metals may contain elements such as Co, Cr, Al, Cu, V, and Ni, classified as allergic. Additionally, elements like Cd, Be, Pb, Ba, and Th are classified as toxic [144,145]. These considerations are crucial in ensuring the safety and efficacy of implants in medical applications. Elements like Ni, Co, and Cr are released from 316L stainless steel and Co-Cr alloy due to corrosion in the body environment [146,147]. The toxicity of Ni causes skin-related diseases like dermatitis, and the release of Co causes carcinogenicity [15]. Young’s modulus of 316L SS and Co-Cr alloy is much higher than that of natural bone, causing nonuniform load transfer between bone and implant and leading to bone resorption and implant loosening. Also, 316L SS steel has poor fatigue strength and wear resistance, limiting its application in orthopedics. 316L exhibits acceptable bulk corrosion resistance but is susceptible to localized pitting and crevice corrosion in chloride-rich physiological fluids and may release Ni/Cr ions under certain conditions, leading to allergic responses.

However, due to the low cost of 316L SS compared to all other metallic alloys, 316L SS has maintained its demand in fixation devices like bone plates, bone screws, etc. Magnesium, zinc, and their alloys, due to their biodegradability and potential for preventing revision surgery, have increased the attention of the general orthopedic community for surgical fixation of injured musculoskeletal tissue [148,149,150,151]. It is believed that when a Mg-based alloy is introduced in a saline environment, it degrades to magnesium chloride, oxide, sulfate, or phosphate, and these ions do not cause any adverse effect on local tissues [152,153]. Despite many advantages, there are certain limitations to using Mg-based alloys. The high corrosion rate of Mg-based alloys is one of the major problems [154]. Corrosion causes the evolution of hydrogen gas, which creates the balloon effect in vivo [155]. Due to the high corrosion rate, the pH value of the surrounding surface also increases [156].

Table 2. Table shows a comprehensive list of biomaterials, along with their advantages, disadvantages, and applications [15,17,157].

Material	Advantages	Disadvantages	Applications
SS 316L	Widely available and cost-effective, excellent mechanical properties, biocompatible	High elastic modulus, inadequate resistance to corrosion, low wear resistance, potential to trigger allergic reactions in surrounding tissues, and stress shielding, which can lead to bone resorption	Bone plates, bone screws, pins, wires, etc.
Co-Cr alloys	Excellent resistance to corrosion, fatigue, and wear. High mechanical strength. Sustained biocompatibility over the long term	High cost, limited machinability, induction of stress shielding, potential biological toxicity from the release of cobalt (Co), chromium (Cr), and nickel (Ni) ions	Shorter-term implants, bone plates and wires, total hip replacements (THR), and stem or hard-on-hard bearing system
Mg alloy	Biocompatible, biodegradable, bioresorbable, similar density, Young’s modulus is that of natural bone, less stress-shielding effect, and lightweight	Hydrogen evolution during degradation and less corrosion resistance	Bone screws, bone plates, bone pins, etc.
Ti alloy	Excellent resistance, lower modulus, stronger than stainless steel, lightweight, and biocompatible	Poor wear resistance, poor bending ductility, and expensive	Fracture fixation devices such as plates, nails, rods, screws, fasteners, and wires; femoral hip stems; total koint replacement (TJR) systems; and arthroplasty procedures, particularly for hip and knee joints
Alumina (Al2O3)	Biocompatibility and bio-inert behavior, elevated hardness, strength, resistance to abrasion, minimal formation of fibrous tissue at the implant–tissue interface	Low fracture toughness, brittleness, limited ductility, and radiopacity	Porous coatings for femoral stems, femoral head, bone screws and plates, and knee prosthesis
Zirconia (Zr2O3)	Excellent fracture toughness; high flexural strength; low Young’s modulus; closely matching that of bone; bio-inert nature; good biocompatibility; and non-toxic behavior within the biological environment	Phase transformation, brittleness, and low toughness	Femoral head, artificial knee, bone screws, and plates
Bioglass	Biocompatibility, bioactivity, promoting integration with surrounding tissue, non-toxicity, and brittleness, which may limit load-bearing applications	Brittleness, low tensile strength, and poor fatigue resistance	Artificial bone and dental implants
Hydroxyapatite (HAp)	Bio-resorbable, bioactive, biocompatible, similar composition to bone, and good osteoconductive properties	Brittleness, low tensile strength, and poor fatigue resistance	Femoral knee, femoral hips, tibial components, and acetabular cup

Table 2. Table shows a comprehensive list of biomaterials, along with their advantages, disadvantages, and applications [15,17,157].

Material	Advantages	Disadvantages	Applications
SS 316L	Widely available and cost-effective, excellent mechanical properties, biocompatible	High elastic modulus, inadequate resistance to corrosion, low wear resistance, potential to trigger allergic reactions in surrounding tissues, and stress shielding, which can lead to bone resorption	Bone plates, bone screws, pins, wires, etc.
Co-Cr alloys	Excellent resistance to corrosion, fatigue, and wear. High mechanical strength. Sustained biocompatibility over the long term	High cost, limited machinability, induction of stress shielding, potential biological toxicity from the release of cobalt (Co), chromium (Cr), and nickel (Ni) ions	Shorter-term implants, bone plates and wires, total hip replacements (THR), and stem or hard-on-hard bearing system
Mg alloy	Biocompatible, biodegradable, bioresorbable, similar density, Young’s modulus is that of natural bone, less stress-shielding effect, and lightweight	Hydrogen evolution during degradation and less corrosion resistance	Bone screws, bone plates, bone pins, etc.
Ti alloy	Excellent resistance, lower modulus, stronger than stainless steel, lightweight, and biocompatible	Poor wear resistance, poor bending ductility, and expensive	Fracture fixation devices such as plates, nails, rods, screws, fasteners, and wires; femoral hip stems; total koint replacement (TJR) systems; and arthroplasty procedures, particularly for hip and knee joints
Alumina (Al2O3)	Biocompatibility and bio-inert behavior, elevated hardness, strength, resistance to abrasion, minimal formation of fibrous tissue at the implant–tissue interface	Low fracture toughness, brittleness, limited ductility, and radiopacity	Porous coatings for femoral stems, femoral head, bone screws and plates, and knee prosthesis
Zirconia (Zr2O3)	Excellent fracture toughness; high flexural strength; low Young’s modulus; closely matching that of bone; bio-inert nature; good biocompatibility; and non-toxic behavior within the biological environment	Phase transformation, brittleness, and low toughness	Femoral head, artificial knee, bone screws, and plates
Bioglass	Biocompatibility, bioactivity, promoting integration with surrounding tissue, non-toxicity, and brittleness, which may limit load-bearing applications	Brittleness, low tensile strength, and poor fatigue resistance	Artificial bone and dental implants
Hydroxyapatite (HAp)	Bio-resorbable, bioactive, biocompatible, similar composition to bone, and good osteoconductive properties	Brittleness, low tensile strength, and poor fatigue resistance	Femoral knee, femoral hips, tibial components, and acetabular cup

5. Titanium and Its Alloys: Material of Ultimate Choice for Implant Application

Titanium (Ti) is the ninth most abundant element in the lithosphere, as it is a constituent of practically all crystalline rock. Reverend William Gregor discovered it in 1798 [114]. Thorough explanations of titanium’s science, technology, and applications may be found in recent investigations. During the last 30 years, titanium and its alloy manufacturing techniques have advanced at a faster rate than any other structural material in metallurgical history. To enhance the performance and dependability of Ti and its alloys and prevent unanticipated failures, there has been an exceptionally high focus on removing manufacturing errors, from extraction and melting procedures to machining and joining. We observe that the primary impediment to an even wider variety of Ti and its alloy applications remains costly. The need to convert Ti oxides or chlorides to their metallic state accounts for a significant portion of the material’s cost. The industrial reduction technique is a batch operation that requires a lot of energy. It was developed many years ago by Kroll. Several initiatives are underway to co-reduce mixed chlorides using various reduction techniques to create metallic or alloyed Ti. Several of these techniques have proven workable in the lab, but scaling up to manufacturing volumes is proving more difficult than anticipated.

Ti is a transition metal with atomic number 22 with an incomplete valence shell. It can form a substitutional solid solution (SSS) with an element having a size factor of ±20%. The melting point of Ti is ~1678 °C, and the crystal structure of Ti is a hexagonally closed-packed (hcp) α-Ti structure up to the beta transus temperature (882.5 °C), transforming to a body-centered cubic β-Ti structure (bcc) above this temperature (Figure 5) [158,159]. The nature of the alloying element determines the microstructure (alpha (α) to beta (β) transformation temperature) of the Ti-based alloy. Elements like Al, O, N, etc., are commonly known as α-stabilizers because they tend to stabilize the α phase, as adding these elements increases the β transus temperature. Similarly, elements like V, Mo, Nb, Fe, Cr, etc., stabilize the β transus temperature and are known as β stabilizers. The addition of this element depresses the β transus temperature.

Titanium alloys are generally classified based on alpha (α) and beta (β) phases. They are classified as α, near-α, (α + β), and metastable β-alloy. The α alloy consists of only α phase (with or without the help of α stabilizers, i.e., Al, O, N, etc.). Near α alloys are a special class of α alloys comprising 1–2% of β stabilizers and about 5–10% of the β phase. Similarly, α + β alloys contain 10–30% of the β phase, and alloys with higher β stabilizers where the β phase is formed by fast cooling are known as metastable β alloys. Generally, α + β or metastable β alloys are employed in biomedical applications. Among all these alloys, α and (α + β) alloys are considered first-generation Ti alloys and possess a high Young’s modulus value (110 GPa). The development of a second-generation Ti-based alloy, i.e., β alloy, was established in 1990 [160]. Due to their ability to possess a lower Young’s modulus value (55–90 GPa), β alloys are the material of choice for orthopedic applications. The different Ti alloys and the mechanical properties used for biomedical applications are mentioned in

Table 3. List showing the mechanical properties of titanium and its alloys (E—modulus in GPa; UTS—ultimate tensile strength in MPa) [15].

Material	Standard	E	UTS	Alloy Composition
First-generation biomaterials (1950–1990)
Commercially pure Ti (CP grade 1–4)	ASTM F1341	100	240–550	α
Ti–6Al–4V ELI wrought	ASTM F136	110	860–965	α + β
Ti–6Al–4V ELI standard grade	ASTM F1472	112	895–930	α + β
Ti–6Al–7Nb wrought	ASTM F1295	110	900–1050	α + β
Ti–5Al–2.5Fe	-	110	1020	α + β
Second-generation biomaterials (1990–to date)
Ti–13Nb–13Zr wrought	ASTM F1713	79–84	973–1037	Metastable β
Ti–12Mo–6Zr–2Fe (TMZF)	ASTM F1813	74–85	1060–1100	β
Ti–35Nb–7Zr–5Ta (TNZT)	-	55	596	β
Ti–29Nb–13Ta–4.6Zr	-	65	911	β
Ti–35Nb–5Ta–7Zr–0.40 (TNZTO)	-	66	1010	β
Ti–15Mo–5Zr–3Al	-	22	-	β
Ti–Mo	ASTM F2066	-	-	β

. Figure 6 highlights the various biomedical applications where Ti and its alloys are used.

6. Fabrication of Porous Titanium Using Various Powder Metallurgical Techniques

Ti-based materials have low thermal conductivity and high reactivity with the surrounding environment. Due to this, their machining, melting, and casting are difficult. Therefore, Ti-based components are generally machined from forged Ti blanks at a low speed; in this procedure, ~95% of the raw materials are lost as scrap, and recycling this scrap is still a challenge [161]. Also, dense Ti-based alloys tend to have a stress-shielding effect due to their high modulus. Incorporating pores is a promising solution to reduce the stress-shielding effect in Ti-based implants, but manufacturing porous Ti-based structures is not technically easy. Various manufacturing methods have been established to produce porous titanium structures. Figure 7 presents a graphical comparison of porosity levels and pore size distributions achieved through these different fabrication techniques [160]. The use of porous material in artificial joint replacement is an attractive field of research as it includes different methods and materials that can be used to reduce stiffness mismatch with bone. In the present study, different methods for the synthesis of porous Ti scaffolds are described. Figure 8 shows the classification of various techniques that can be used to fabricate porous Ti structures. However, the following sections will focus exclusively on a few processes that can fabricate porous Ti-based structures.

6.1. Space Holder Technique

The space holder technique is one of the important powder metallurgical techniques that can control the size and shape of the pores, amount of porosity, etc. Because these parameters depend on the size of space-holder particles. The basic requirement of this process is that the particle size of the metal powder should be less than the particle size of the space holder. This method involves the addition of space holder particles with the metal powder, followed by mixing them uniformly throughout. Then, the mixture is compacted uniaxially to form a green compact. The green sample is pre-sintered at an optimized temperature (low temperature) so that the complete removal of space holder particles can occur. This also leads to the initial sintering of metal powders. Finally, the pre-sintered samples are sintered at an elevated temperature in an inert atmosphere to avoid contamination (including metal oxidation). The schematic representation of the process for preparing porous Ti via the space holder technique is shown in Figure 9. Several materials have been used as space holders, including bio-wastes, metals like Mg granules, urea, ammonium hydro carbonate, water-soluble materials (like sucrose, potassium chloride, and sodium chloride), paraformaldehyde, etc. Before selecting the space holder, the factors to be considered are its affinity with Ti, the amount of left-out residue after burning, and ease of processing. Complete removal of space holder material from the substrate is the main problem associated with the space holder technique because the presence of any residue may impart detrimental effects, which may influence biocompatibility. Materials like NaCl and sucrose are suggested as they can be removed completely when treated with water [162].

Kim et al. [163] used sacrificial Mg granules as space holder particles since Mg ions are directly involved in numerous biological mechanisms in our body, such as the channelizing of ions, DNA stabilization, enzyme activation, and stimulation of cell growth and proliferation. So, the problems associated with space holders, like partial removal of space holder particles, are completely mitigated. Mg particles of size 20 mesh to 100 mesh with 0.5 wt.% ethanol as a binder were used. The mixture of Ti and Mg was compacted uniaxially, followed by removing space holder particles by dipping them in HCl and ethanol for 24 h. Sintering the green samples at an elevated temperature of 1300 °C for 2 h in high vacuum results in controlling the porosity within 50–71% and pore size within 132–262 µm. The mechanical testing of the samples reveals that the compressive and rupture strengths are 59–280 MPa and 85 MPa, respectively. Esen and Bor [164] processed Ti foam using Mg as a space holder particle; the porosity content was observed to vary between 45 and 70%; the pore size was found to be 525 µm; Young’s modulus ranged between 0.42 GPa and 8.8 GPa.

Bio-waste, like rice husk, is a rich source of silica, and its low-temperature combustion property can be utilized for space-holder applications. Due to the amorphous nature of silica and its high relative reactivity along with carbon for thermal reduction, rice husk can be a favorable candidate for synthesizing low-temperature porous material [165,166]. Apart from the above space holders, researchers have used many other materials to produce porous titanium scaffolds. Dabrowski et al. [3] employed paraformaldehyde, with an average particle size of 500 µm, as a space holder in the fabrication of porous titanium implants due to its ability to fully decompose at relatively low temperatures. The resulting porosity can vary between 60% and 70%, and Young’s modulus can vary between 1 GPa and 8 GPa, which is very close to that of cancellous bone. Xiao Jian et al. [167] studied the porous Ti structures produced with varied spacer contents (Figure 10).

Reports suggest that urea is an effective space-holder material that tends to decompose at low temperatures, resulting in the formation of a pore. Vasconcellos et al. [168] synthesized porous Ti with 3D interconnected pores with a size of ~480 µm and a total porosity of 36%. Wenjuan et al. [169] used urea of size 200–600 µm as a space holder and polyethylene glycol as a binder to synthesize a porous Ti scaffold with porosity varying between 55 and 75% and pore size varying between 200 µm and 500 µm. The porous scaffold reveals Young’s modulus in the range of 3–6.4 GPa, which is close to that of natural bone. It has also been noticed that the rapid decomposition of urea at low temperatures causes rough control over porosity. The needle-like shape of urea particles provides sharp corners and notches in the pore, which lead to stress concentration, deteriorating their mechanical properties. It is also found that the remaining urea in pores makes the implant unfit for medical use [170].

Xiang et al. [171] prepared porous Ti with a porosity between 44 and 77% by using ammonium acid carbonate as a space holder. The fabricated porous Ti scaffolds have pore sizes between 200 µm and 500 µm, with Young’s modulus and compressive strength values ranging between 2.1 GPa and 3.4 GPa and 60 MPa and 140 MPa, respectively. In addition to the space holders, other materials can serve as pore formers, meeting essential criteria such as complete residue removal through dissolution in water and being economically viable. In this class, sucrose and NaCl are the best-suited materials that can be used as a pore former. Torres et al. [172] reported excellent adhesion and proliferation of cells when NaCl is used as a space holder. Also, any left-out residue in the scaffold will not affect the in vivo performance of the Ti alloy.

Chen et al. [173] introduced a novel space holder material that offers excellent biocompatibility for fabricating porous titanium structures with open and interconnected pore networks. Spherical sugar pellets were used to synthesize porous Ti with 20 to 54% porosity and a pore size ranging from 212 µm to 500 µm. The Young’s modulus of the scaffold is found between 12.1 GPa and 18.5 GPa, which is very close to that of natural bone. Polymethyl methacrylate (PMMA) is also used as a space holder material to fabricate porous scaffolds. Li et al. [174] used PMMA to produce macro pores of size 200–400 µm and porosity 10–65%. The green compact was heated between 250 and 450 °C to completely remove the space holder particles. The compressive strength and elastic modulus were observed to be between 32 MPa and 530 MPa and 0.7 GPa and 23.3 GPa, respectively.

Table 4. Mechanical properties of titanium scaffolds prepared by different space holders (P—porosity in %; PS—pore size (µm); E—Young’s modulus in GPa; YS—Yield strength in MPa; UTS—ultimate tensile strength in MPa; UCS—ultimate compressive strength in MPa).

Space Holder Material	P	PS	E	YS	UTS	UCS	Ref.
Mo Wire	32–47	-	23–62	76–192	-	-	[175]
Mg	45–70	525	0.42–8.8	15–116	-	-	[164]
Mg	50–71	262–132	-	-	-	59–280	[163]
Mg	30–50	-	15–44	117–222	-	-	[176]
Ti Fibers	35–84	150–600	2–4	-	200–600	-	[177]
Rice Husk	50–60	100–550	-	-	-	17–70	[165]
Rice Husk	25–36	-	-	-	-	440–938	[166]
Rice Husk	15–34	-	6–15	-	-	116–396	[16]
Sucrose	20–54	212–500	12–50	-	-	-	[173]
Urea	36	480	-	-	-	-	[168]
Urea	55–75	200–500	3–6	-	10–35	-	[169]

summarizes the mechanical properties of different porous materials.

6.2. Replication Method

The synthesis of porous Ti with the help of replicating polymeric sponges followed by high-temperature sintering is a unique technique. This process offers the fabrication of scaffolds with a high degree of porosity and highly interconnected microspores with identical shapes and sizes. The porous structure produced via this method has a pore shape and size like cancellous bone [178,179,180,181]. In this method, the precise regulation of the rapid drying process of the coated slurry is essential for achieving the desired structural characteristics. Cachinho et al. [182] described a unique method of preparing a porous Ti scaffold in which the scaffold was prepared by replication of a sponge followed by reactive sintering. The main advantage of this method is the easy production of complex shapes at a low cost. The use of a sacrificial polymeric sponge results in scaffolds with interconnected pores, a critical feature for promoting bone ingrowth and vascularization in newly formed tissue [183]. Such porous architectures are commonly applied in dental implants, permanent osteosynthesis plates, and intervertebral disc replacements [184]. Cachinho et al. [182] reported using 45 vol.% TiH₂ powder with a mean particle size of 15.6 µm and a specific surface area of 0.5336 m²/g. The polymeric sponge blocks were dipped into the slurry and infiltrated. After the removal of excess slurry, the samples are dried at room temperature for a period of 24 h. The sintering of samples at a low heating rate of 1 °C/min with dwelling at 500 °C for 2 h and 1000 °C for 4 h results in the formation of a highly porous Ti scaffold with a porosity of 75% and a pore size ranging between 100 µm and 600 µm. The porosity range of 100 µm and 600 µm is appropriate for the growth of new bone tissues and the transport of body fluids. Further, to improve biological properties, the porous Ti is coated with hydroxyapatite and heat-treated at 700 °C. Li et al. [185] synthesized a porous Ti-6Al-4V alloy by replicating 70 wt.% of Ti-6Al-4V powder in water and ammonia solutions. The high-temperature sintering of samples results in open-cell porous Ti struts with a porosity of 88% and a compressive strength of 10 MPa. It should be noticed that the second deposition of powder slurry on the previously sintered scaffold, followed by re-sintering, increases density and compressive strength to 36 MPa. Wang et al. [186], in his study, proposed an improved sponge replication method. A novel solvent consisting of ethanol and water was used to maintain a fast-drying rate and appropriate viscosity of Ti slurry. This slurry was used for multiple Ti coatings, and the Ti scaffold prepared possesses a compressive strength of ~84 MPa with a porosity of ~66%.

Wang et al. [186] explored an enhanced polymeric sponge replication method to fabricate biomedical porous titanium scaffolds. By optimizing key processing parameters, they achieved scaffolds with open and interconnected pore architectures. In their approach, titanium slurries were prepared using a mixture of ethanol and water as the liquid phase. Figure 11a,b display the asymmetrical titanium particles and the open, interconnected structure of the polyurethane (PU) foam used as the sacrificial template. Figure 11c outlines the process flowchart for fabricating porous titanium via this improved replication technique. Notably, three modifications distinguished this method from conventional approaches: (1) Ethanol–water mixtures replaced water alone to improve slurry formulation; (2) excess slurry was removed through centrifugation rather than squeezing, preventing the formation of closed pores; (3) variations in slurry viscosity were introduced for more uniform coating. SEM images in Figure 11d,e illustrate the resulting microstructures, showing that scaffolds fabricated with ethanol exhibited a denser titanium coating compared to those produced using only water. This densification, attributed to ethanol’s anti-foaming properties, resulted in enhanced compressive strength. Furthermore, in vitro studies demonstrated that the scaffolds supported the adhesion, proliferation, and growth of mesenchymal stem cells (MSCs), confirming their suitability for biomedical applications.

6.3. Entangled Metal Wire Technique

Porous Ti implants fabricated via conventional methods exhibit low toughness and tensile strength. The major drawback of these conventional techniques is the difficulty in avoiding contamination and impurities in Ti that arise during processing. Sometimes, undesirable cracks and metallographic defects in sintered Ti struts make them brittle, and thus, they fail to bear tensile load [187]. Also, the porous Ti fabricated by powder metallurgical techniques using a space holder and plasma spray technique has low ductility that may break in the body’s environment when subjected to uncertain overloading and accidents [188]. The entangled metal wire technique (EMWT) is a novel technique that improves these mechanical properties. In this technique, a Ti wire of a diameter of ~0.08 mm to 0.27 mm is used as raw material, and this wire is coiled around a 1.5 mm-diameter rod to form a coiled spring-like structure. The coiled structure is stretched equably such that the distance between two spirals (screw pitch) reaches the external diameter of the coil. This stretched coil is now entangled around a 1 mm diameter rod to form a pre-compacted sample. Finally, this pre-compacted sample is compacted with the help of a piston in a cylindrical die [189].

The mechanism of fabrication and sample porous titanium scaffold prepared by the entangled metal wire technique is shown in Figure 12. Several researchers have fabricated porous titanium scaffolds using EWMT. Zou et al. in 2008 [177] prepared open-cell porous titanium with porosity of 35 to 84% by sintering titanium fibers of 200 µm diameter in a vacuum. The titanium fibers were curved into a helix with the help of a screw, and then, this helix was arranged in a cylindrical form, followed by compaction and vacuum sintering at 1250 °C for 2 h. The resulting porous scaffold has a pore size of 150–600 µm. Young’s modulus was in the range of 3.5–4.2 GPa, and compressive strength was 100–200 MPa. Liu et al. [190] fabricated entangled titanium wire material through different procedures: one with normal wire and another with coiled wire; its compressive and pseudo-elastic hysteresis behavior was investigated. The details of the properties obtained are discussed in

Table 5. Mechanical properties of porous titanium prepared by entangled metal wire technique (P—porosity in %; PS—pore size (µm); E—Young’s modulus in GPa; YS—yield strength in MPa; UTS—ultimate tensile strength in MPa; UCS—ultimate compressive strength in MPa; FS—flexural strength in MPa; NR— not reported).

Method	Material Used	P	PS	E	YS	UTS	UCS	FS	Ref.
EWMT	Entangled Mo Wire	32–47	0.4	23–62	76–192	-	-	-	[175]
EWMT	Ti Wire	35–84	150–600	2–4.2	-	200–600	-	-	[177]
EWMT	Entangled Ti Wires	44–81	NR	0.03–2.25	-	-	-	9–325	[193]
EWMT	Entangled Ti Wire	53–55	NR	0.03–1	3–3.5	-	-	-	[188]
EWMT	Entangled Ti Wire	37–54	NR	22–47	-	-	175–246	-	[191]
EWMT	Entangled Ti Wire	40–55	100–400	0.4–1.4	12.9–52.5	-	-	-	[192]
EWMT	Entangled Ti Wire	45–58	50–200	1.05–0.33	75–124	48–108	-	-	[187]
EMWT	Normally entangled Ti Wire	48–73	-	0.13–0.82	2–31	-	-	-	[190]
	Coiled entangled Ti Wire	48–78	-	0.04–0.62	1–19	-	-	-

. Jiang et al. [191] fabricated an entangled porous titanium composite filled with biodegradable magnesium melted at 700 °C under the protective environment of SF₆ and CO₂ to improve the fixation bonding between the implant and host bone. Bisphenol A glycidyl methacrylate (BisGMA) is suggested as a bonding material to provide strong bonding strength and help in fixing the free nodes of the entangled structure [192]. Wang et al. in 2017 [175] proposed a novel technique for the fabrication of a three-dimensional porous titanium scaffold. This method combined two methods: An entangled Molybdenum wire was used as a space holder, and titanium liquid was cast in a vacuum environment, followed by etching off SH particles in an aqua regia solution. The resulting porous scaffold has three-dimensional interconnected pores with porosity in the range of 32–47% and exhibits elastic modulus in the range of 23–62 GPa and yield strength in the range of 76–192 MPa, as shown in Table 5.

6.4. Spark Plasma Sintering (SPS) and Hot Pressing (HP)

The conventional sintering or pressureless sintering process involves heating Ti and its alloys at an elevated temperature of 1200 °C to 1400 °C (but below the melting point of the material) and a high vacuum of the order of 4 × 10⁻⁴ Pa to 6 × 10⁻⁶ Pa for a long time period of about 24 h to 48 h [194] for densification and homogenization [195,196]. Even after this lengthy procedure, achieving a pore-free homogeneous microstructure is challenging [196]. SPS is an advanced consolidation technique that uses pressure-assisted pulsed current to sinter and can produce porous samples. In this process, the powder is loaded in an electrically conducting die that acts as a heating source when subjected to a pulsed direct current. Thus, the powdered samples will be heated from both sides under uniaxial pressure [197,198,199,200], and due to this fast heating, enhanced mass transfer and rapid powder consolidation will occur [201]. Two theories explain the consolidation mechanism of commercially pure (CP) Ti. According to the first hypothesis, the surface of the powder particles is cleaned and activated by spark discharges generated between metallic powdered particles, thus promoting mass transport for sintering [202,203]. Another hypothesis suggests that the densification of powder is due to particle deformation because, as the temperature increases, the yield strength of the powder particles decreases [204]. Spark plasma sintering (SPS) is also referred to by several alternative names, including field-assisted consolidation, electrical field-activated sintering, plasma-activated sintering, and electrical discharge compaction [203,205,206,207]. SPS is also known by other names, such as field-assisted consolidation technique [205], electrical field-activated sintering [203], plasma-activated sintering [206], and electrical discharge compaction [207]. These methods enable the rapid sintering of metallic powders by applying electrical discharge in combination with swift heating and concurrent pressure. Similarly, in HP, the metallic powder is sintered with the help of electrical resistance in a closed die under uniaxial pressure. Different heating methods, such as induction heating and electric conduction/convection/radiation heating, can be used in HP [208,209]. Ibrahim et al. [210] synthesized porous Ti and its alloy using a cost-effective SPS technique. In this process, porous Ti with different porosities was successfully synthesized by the powder metallurgy technique using NH₄HCO₃ as a space holder and TiH₂ as a foaming agent. SPS is used to consolidate powder at 16 MPa under pressureless conditions. The experimental results showed that pure titanium samples achieved full relative density at a relatively low temperature of 750 °C and a pressure of 16 MPa. The porosity of 53% and Young’s modulus of 40 GPa were achieved in the case of pressureless sintering at a temperature of 1000 °C.

Kashimbetova et al. [211] investigated the fabrication of porous titanium structures using pressure-less spark plasma sintering (PL-SPS). Their study focused on examining how different sintering temperatures influence the microstructural features and mechanical performance of porous Ti. Figure 13a,d,g display the porous strands formed at 1400 °C, where bonding occurred primarily at contact points between adjacent Ti particles in the early sintering stage. At 1500 °C, more developed sintering necks between particles were observed, as shown in Figure 13b,e,h, indicating enhanced densification with limited grain growth. Further sintering at 1600 °C brought the structures into the mid-stage of sintering, marked by significant pore shrinkage, though overall porosity was maintained (Figure 13c,f,i). Increased densification at higher temperatures had a direct effect on compressive yield strength (Figure 13j). At 1400 °C, the weakly bonded particles resulted in a low yield strength of 4.7 MPa. This increased to 26.7 MPa at 1500 °C due to the growth of sintering necks and reached 52.6 MPa at 1600 °C because of further densification. The samples sintered at 1500 °C and 1600 °C exhibited elastic deformation followed by strain hardening during compression. In contrast, the structure sintered at 1400 °C reached peak stress shortly after yielding and then gradually fractured, as shown in Figure 13k. The compressed sample sintered at 1600 °C maintained its pore geometry, particularly in the printing plane (Figure 13l). The findings suggest that intra-strand porosity significantly influences mechanical strength when inter-strand porosity is kept constant. A comparison of the properties of porous titanium synthesized using SPS and hot pressing (HP) from various studies is presented in Table 6.

Figure 13. Images depicting porous titanium fabricated using pressureless spark plasma sintering as a function of varying sintering temperatures: (a,d,g) 1400 °C, (b,e,h) 1500 °C, and (c,f,i) 1600 °C. (j) Compressive stress–strain curves of porous titanium as a function of different sintering temperatures. Images of the (k) fractured sample sintered at 1400 °C and (l) compressed sample (sintered at 1600 °C). Reproduced with permission from [211].

Figure 13. Images depicting porous titanium fabricated using pressureless spark plasma sintering as a function of varying sintering temperatures: (a,d,g) 1400 °C, (b,e,h) 1500 °C, and (c,f,i) 1600 °C. (j) Compressive stress–strain curves of porous titanium as a function of different sintering temperatures. Images of the (k) fractured sample sintered at 1400 °C and (l) compressed sample (sintered at 1600 °C). Reproduced with permission from [211].

6.5. Microwave Sintering

Electromagnetic waves with a frequency in the range of 300 MHz to 300 GHz are referred to as microwaves. The most used microwaves for material processing have frequencies between 2.45 GHz and 915 MHz [212]. When materials interact with microwaves, they convert electromagnetic energy into heat energy within the material. The main advantage of microwave sintering is the complete sintering of the material without forming impurities (like oxides). This technique is applicable for sintering ceramics, metals, and composites, offering advantages such as reduced processing time and energy consumption, cost-effectiveness, and environmental sustainability [213]. In powder metallurgy-based manufacturing routes, particularly after pressing and sintering, incompletely bonded or loosely attached particles may remain on the surface of porous titanium structures, and their release into the body can trigger inflammatory responses, macrophage activation, and osteolysis, ultimately jeopardizing implant stability [214]. Preventing such complications requires ensuring adequate particle bonding through optimized sintering conditions and robust neck formation, followed by thorough post-processing steps such as ultrasonic cleaning, chemical/acid etching, abrasive or mechanical removal of weakly attached particles, and thermal treatments to consolidate surface layers. Complementary quality-control procedures, including micro-CT analysis, surface morphology inspection, and mechanical integrity assessments, are also essential to verify the absence of detachable particles. A comprehensive discussion of surface post-treatments for bio-implants is presented in Section 6.6 of this review article. Table 6 demonstrates the processing conditions and physical and mechanical properties obtained using SPS, HP, and microwave sintering.

Table 6. Processing conditions and mechanical properties of porous titanium prepared by spark plasma sintering, hot pressing, and microwave sintering (ST—sintering temperature in °C; TP—time and pressure; P—porosity in %; PS—pore size (µm); E—Young’s modulus (MPa); YS—yield strength in MPa; UCS—ultimate compressive strength in MPa; UTS—ultimate tensile strength in MPa).

p	ST	TP	P	PS	E	YS	UCS	UTS	Ref.
Spark plasma sintering
Pure Ti	750	16 MPa	Fully dense	-	~125	-	-	-	[210]
Pure Ti	1000	Pressureless	53	-	40	-	-	-
Ti5Mn alloy	950	Pressureless	56	-	35	-	-	-
Ti5Mn alloy	1100	Pressureless	21	-	52	-	-	-
Pure Ti	700	-	30–70	125–800	6–36	27–94	-	-	[215]
β-alloy Ti-45Nb (gas-atomized)	1000	10 min, 30 MPa	0.5 ± 0.1	-	72 ± 1	550	-	-	[216]
β-alloy Ti-45Nb (milled)	1000	10 min, 30 MPa	4.0 ± 0.2	-	72 ± 1	867	-	-	[216]
Ti-6Al-4V	700	3 min, 30 MPa	32 ± 0.2	-	-	-	125	-	[217]
Pure Ti	600	3 min, 30 MPa	32 ± 0.4	-	-	-	113	-	[217]
CP Ti (Grade 1) Powder	900	5 min, 60 MPa	-	-	-	340	-	445	[218,219]
Cryomilled nanocrystalline CP Ti (Grade 2) powder	850	-	-	-	-	770	-	840
CP Ti (Grade 3) powder	900	5 min, 60 MPa	-	-	-	595	-	720
Wrought titanium grade 4	-	3 min, 80 MPa	-	-	-	480–635	-	655–690
Hot pressing
Ti-45Nb (gas-atomized)	600	30 min, 700 MPa	0.7 ± 0.2	-	70 ± 1	447	-	-	[216]
Ti-45Nb (milled)	600	30 min, 700 MPa	3.7 ± 0.1	-	70 ± 1	940	-	-	[216]
Microwave sintering
Ti6Al4V/MWCNTi powder	1620	-	25	-	11 ± 3	145	270	-	[213]

Table 6. Processing conditions and mechanical properties of porous titanium prepared by spark plasma sintering, hot pressing, and microwave sintering (ST—sintering temperature in °C; TP—time and pressure; P—porosity in %; PS—pore size (µm); E—Young’s modulus (MPa); YS—yield strength in MPa; UCS—ultimate compressive strength in MPa; UTS—ultimate tensile strength in MPa).

p	ST	TP	P	PS	E	YS	UCS	UTS	Ref.
Spark plasma sintering
Pure Ti	750	16 MPa	Fully dense	-	~125	-	-	-	[210]
Pure Ti	1000	Pressureless	53	-	40	-	-	-
Ti5Mn alloy	950	Pressureless	56	-	35	-	-	-
Ti5Mn alloy	1100	Pressureless	21	-	52	-	-	-
Pure Ti	700	-	30–70	125–800	6–36	27–94	-	-	[215]
β-alloy Ti-45Nb (gas-atomized)	1000	10 min, 30 MPa	0.5 ± 0.1	-	72 ± 1	550	-	-	[216]
β-alloy Ti-45Nb (milled)	1000	10 min, 30 MPa	4.0 ± 0.2	-	72 ± 1	867	-	-	[216]
Ti-6Al-4V	700	3 min, 30 MPa	32 ± 0.2	-	-	-	125	-	[217]
Pure Ti	600	3 min, 30 MPa	32 ± 0.4	-	-	-	113	-	[217]
CP Ti (Grade 1) Powder	900	5 min, 60 MPa	-	-	-	340	-	445	[218,219]
Cryomilled nanocrystalline CP Ti (Grade 2) powder	850	-	-	-	-	770	-	840
CP Ti (Grade 3) powder	900	5 min, 60 MPa	-	-	-	595	-	720
Wrought titanium grade 4	-	3 min, 80 MPa	-	-	-	480–635	-	655–690
Hot pressing
Ti-45Nb (gas-atomized)	600	30 min, 700 MPa	0.7 ± 0.2	-	70 ± 1	447	-	-	[216]
Ti-45Nb (milled)	600	30 min, 700 MPa	3.7 ± 0.1	-	70 ± 1	940	-	-	[216]
Microwave sintering
Ti6Al4V/MWCNTi powder	1620	-	25	-	11 ± 3	145	270	-	[213]

6.6. Additive Manufacturing (AM) or Rapid Prototyping (RP)

Additive manufacturing (AM), also known as rapid prototyping (RP), is a computer-assisted advanced fabrication technique that utilizes computer-aided design (CAD) and computer-aided manufacturing (CAM) models to construct predefined microstructures, macrostructures, and precisely controlled hierarchical architectures [220,221]. As the demand of the manufacturing sector is more focused on precision and specific design, the AM technique has emerged as a strong tool for mitigating problems associated with conventional powder metallurgy techniques like material loss during post-processing (like machining, drilling, etc.), large production time, difficulty in producing complex shapes, etc. [222,223,224,225]. AM has been extensively utilized in the medical field by creating patient-specific components based on patients’ medical imaging [226]. The major advantage of this technology is the control over scaffold pore structure, including pore size, shape, volume, and interconnectivity [227]. It is a layer-by-layer fabrication process in which the selected part is built in a CAD file [228,229,230]. The file is sliced along the Z-axis in a virtual environment, and a machine-specific tool path is generated for each slice.

AM offers specific advantages like site-specific deposition with higher cooling rates and can easily produce intricate shapes [231]. The interconnectivity of the porous structures is a key issue in porous metallic structures. As mentioned above, this issue cannot be perfectly controlled by other conventional manufacturing techniques. In contrast, the final structure of the AM-fabricated components relies on the initial design with interconnected structures through AM. The porous Ti can be easily fabricated without design restrictions. Additive manufacturing (AM) techniques can be categorized based on the energy–powder interaction and consolidation mechanism, including methods such as electron beam melting (EBM), selective laser melting (SLM)—also referred to as laser powder bed fusion (LPBF)—laser-engineered net shaping (LENS), and binder jetting (BJG) [232,233]. Though there are several AM processes, only a few techniques, including SLM/LPBF and EBM, are widely used and can fabricate a wide variety of materials. Other AM processes like directed energy deposition (DED) are widely used for higher deposition rates and large-scale manufacturing [234].

Apart from fabricating intricate shapes, the AM structures also exhibit improved properties because of hierarchical microstructures resulting from higher cooling rates observed during the process [235,236,237]. Moreover, AM produces a unique microstructure and texture compared to conventional processing, which helps to improve the strength by maintaining the same porosity [238]. As mentioned in Section 5, Ti shows allotropy. Moreover, Ti6Al4V exhibits the coexistence of α and β microstructures, providing versatility in the Ti-based alloys. Since SLM/LPBF offers high cooling rates, a fine α’ martensite or α’ martensite-based microstructure is observed in these samples [239,240,241]. The α’ martensite appears as an acicular or needle-shaped microstructure, which has an hcp crystal structure like the α phase. Like ferrous martensite, Ti-based martensite also helps improve its strength at the cost of ductility. Hence, AM-fabricated structures may require post-processing, such as heat treatment, to tune their properties according to the requirements [242,243]. Sabban et al. [244] introduced a novel heat treatment approach that successfully transformed martensitic laths into a bimodal globularized microstructure in the AM Ti64 alloy, resulting in an 80% enhancement of ductility and a 66% improvement in toughness. In subsequent research by the same group, Gupta et al. [245] fabricated a Ti64-based bone plate using the AM technique. They employed a specially designed process of repeated cyclic heating and cooling, as previously described, to transform the microstructure from acicular to bimodal without the need for plastic deformation before heat treatment. Moreover, the in situ dissolution of α’ is also possible by changing the processing parameters and providing double or triple laser scanning in selective laser manufacturing [246,247,248].

AM-fabricated porous Ti alloys exhibit relatively better mechanical properties due to interconnected structures [35,249]. The good interconnectivity of the porous implants enables uniform load distribution among the structures, while poorly connected porous structures typically fail prematurely at the joints. Ample evidence in the literature suggests premature failure for porous metallic implants [250,251]. Given the advantages, there are also some drawbacks, including surface defects, which usually tend to deteriorate the build’s properties. One can understand the importance of AM for biomedical applications, as illustrated in Figure 14, Figure 15, Figure 16, Figure 17 and Figure 18. Biomedical implant dimensions and requirements vary from patient to patient, making other routes challenging. Porous Ti implants are usually used for orthopedic implants owing to their outstanding mechanical and biological properties. The elastic modulus of bone ranges between 10 and 20 GPa and is mostly composed of inner cancellous and outer cortical bone [252]. Meanwhile, dense Ti alloys have an elastic modulus of around 110 GPa. However, it is worth mentioning that metallic alloys fabricated via AM tend to possess a lower elastic modulus than regular Ti-based alloys, making them comparable to bone [32]. In addition, porous Ti also possesses high rigidity and good fatigue properties, making it a suitable candidate for biomedical implants.

Figure 14. (a) Medical titanium alloy for anti-corrosive properties in femoral implant [253]; (b) 3D-printed titanium hip ball-and-socket implant: image credit: Jabil [254]; (c) first trabecular titanium hip cup produced by Lima Corporate, San Daniele del Friuli, Italy and Arcam Mölndal, Sweden (now a GE Additive company) in the year 2007 [255].

Figure 14. (a) Medical titanium alloy for anti-corrosive properties in femoral implant [253]; (b) 3D-printed titanium hip ball-and-socket implant: image credit: Jabil [254]; (c) first trabecular titanium hip cup produced by Lima Corporate, San Daniele del Friuli, Italy and Arcam Mölndal, Sweden (now a GE Additive company) in the year 2007 [255].

Figure 14 shows the Ti femoral implant and the hip ball-and-socket joints [253,254,255]. Pelvic injury is prevalent in trauma patients, caused by impact, rolling, steep falls, and other injuries. Recently, AM has often been employed for preoperative examination and simulated surgery treating pelvic fractures. Investigations demonstrate the possibility of employing AM-fabricated Ti alloys for pelvic implants [256,257]. Wong et al. [258] manufactured pelvic-specific implants from Ti alloy and evaluated their performance. Broekhuis et al. [259] used AM to customize and create metallic pelvic prostheses for acetabular repair following tumor resections.

The spinal column is composed of interconnecting bones called vertebrae, which begin at the base of the head and finish at the tailbone in the lower back. The vertebral column is formed by stacking one bony vertebra on top of the next. The intervertebral discs are disc cushions that sit between each of the vertebral bodies. These discs act as shock absorbers, connecting the vertebrae (Figure 15). These discs also allow for the spine’s bending and twisting action, assisted by facet joints. The spinal column protects the spinal cord, which links the brain’s nerves to other body parts. In 2017, the United States FDA (Food and Drug Administration) authorized two alloys of Ti spinal implants made by AM. One example of an additively manufactured vertebral implant is the HAWKEYE Ti, while another is the NEXXT MATRIXX 3D-printed spinal implant. Hollander et al. [260,261] employed AM technology to fabricate spinal implants with tailored porosity and demonstrated that their surface characteristics supported the adhesion and proliferation of human osteoblasts. Lin et al. [261] utilized the SLM/LPBF method to achieve a porosity of approximately 55%, resulting in an elastic modulus close to that of natural human bone (~3 GPa). This structure also exhibited superior bone regeneration capabilities compared to conventional PEEK cages.

Figure 15. Human spinal column with the insert in black illustrating ATEC’s IdentiTi posterior curved porous titanium interbody implants [262].

Figure 15. Human spinal column with the insert in black illustrating ATEC’s IdentiTi posterior curved porous titanium interbody implants [262].

In certain cases, implants can act as bone replacements without being subjected to considerable and constant bearing or additional stress. Figure 16 is the cranial implant developed by Novax DMA [263] for a patient requiring a large metallic implant. In the case of a skull implant, external variables, particularly the implant itself, should be promoted rather than hindering the healing process. Most significantly, the implant should fit as perfectly as feasible, as offered by AM. The manufacturing technique, which uses a laser for building up a material (Ti) layer by layer, allows for maximum customization in terms of form and size. Because of the large hole in the patient’s bone structure, integrating biological activities and minimizing heat loss into the brain tissue were given high priority. A Ti construction may also be impermeable to brain tissue fluid. Only permeable construction may satisfy the requirements.

Figure 16. (a) Skull implants developed by Novax DMA. (b) Magnified image of the porous titanium cranial implants with porosity levels of 95% [263].

Figure 16. (a) Skull implants developed by Novax DMA. (b) Magnified image of the porous titanium cranial implants with porosity levels of 95% [263].

A lattice-structured implant with skull-integrated, screw-in fasteners allows fluids to pass through and combine with the skull’s bone. Furthermore, such a construction would provide insulation, reducing heat transfer entering the cranial cavity. The pore spaces are around 1 mm, whereas the cell links are about 0.2 mm thick. Murr et al. [264] described a reticular skull implant developed through EBM. Mazzoli et al. [265] used CT imaging, computer modeling, and AM to produce a biocompatible Ti skull implant. Zhao et al. [266] observed that personalized Ti-based skull prosthetics made via AM offered higher impact resistance while efficiently repairing skull deformities and protecting intracranial brain tissue. Yan et al. [267] employed EBM to create a mandibular prosthetic implant that had a 3D mesh and a porosity of 81.3%, which fits the parameters for implantation in the human body (Figure 17). Figure 17a illustrates the 3D anatomic model of the human mandible developed using a CT imaging technique. Figure 17b shows the 3D structure developed in CAD software (Unigraphics NX 8.0, EDS) by providing internal 3D meshing for porous structures. Meanwhile, Figure 17c,d demonstrate the mandibular prosthetic implant produced by EBM. Moiduddin et al. [268] utilized a Ti-based alloy to fabricate cheekbones.

Figure 17. (a) Three-dimensional anatomic model of a human mandible prosthetic. (b) Design of mandibular scaffold 3D mesh internal microstructure. (c,d) Titanium alloy scaffolds produced by EBM with porosity levels of 81.3% and a weight of 107 g. Reproduced with permissions from [267].

Figure 17. (a) Three-dimensional anatomic model of a human mandible prosthetic. (b) Design of mandibular scaffold 3D mesh internal microstructure. (c,d) Titanium alloy scaffolds produced by EBM with porosity levels of 81.3% and a weight of 107 g. Reproduced with permissions from [267].

Yanez et al. [269] studied the gyroid porous Ti structures, where triply periodic minimum surfaces (TPMS) have been recognized as an effective approach for developing porous biomaterials. They developed a variety of porous Ti structures with varying degrees of porosities using EBM to investigate mechanical characteristics under compression and torsion loads, as illustrated in Figure 18. Compression testing indicated that deformed gyroid structures possess significant strength and stiffness when subjected to axial loading, particularly when a reinforcing shell is incorporated. In torsional loading, however, conventional gyroids outperformed deformed gyroids with high CAD-defined porosity (90%) in terms of both torsional strength and stiffness. Notably, deformed gyroids exhibited superior mechanical behavior under compressive loads compared to their regular counterparts. Conversely, regular gyroids demonstrated better adaptability to non-axial and torsional loading conditions [269].

Figure 18. (a) Six variations of gyroid-structured porous titanium: normal gyroid 75, normal gyroid 90, deformed gyroid 75, deformed gyroid 90, deformed gyroid with shell reinforcement, and double deformed gyroid. (b) Visual depiction of compressive testing performed on the gyroid porous titanium structures. (c) Image capturing the torsional testing of one of the porous titanium configurations. Reproduced with permissions from [269].

Figure 18. (a) Six variations of gyroid-structured porous titanium: normal gyroid 75, normal gyroid 90, deformed gyroid 75, deformed gyroid 90, deformed gyroid with shell reinforcement, and double deformed gyroid. (b) Visual depiction of compressive testing performed on the gyroid porous titanium structures. (c) Image capturing the torsional testing of one of the porous titanium configurations. Reproduced with permissions from [269].

Figure 19 depicts the many surface flaws that arise during bio-implant manufacturing. AM-processed bio-implants are often not ready for use due to surface defects such as powder adhesions, semi-welds, surface casing, porosity, balling effect, etc., highlighting the need for post-processing. The nature of the designs, prior processing software, and processing equipment all influence post-processing requirements. Post-processing includes removing unwanted material, improving the texture of the surface, aesthetic enhancements, component separation, rebinding and sintering, machining, surface finishing, non-thermal performance enhancement, heat treatment, and quality assurance. It is critical to enhance the surface finish by employing suitable surface quality techniques for AM-built biomedical implants to compete alongside widely recognized traditionally fabricated bio-implants [270,271,272].

Figure 19. Surface defects occurring during AM-fabricated implants (reproduced with permission from [272]).

Figure 19. Surface defects occurring during AM-fabricated implants (reproduced with permission from [272]).

Conventional processes involve employing multipoint or single-point instruments for cutting directly over the surface that is being finished or polished. These include robot-based finishing, rigid tool-based finishing, and computer numerical control-based finishing. Conventional processes are more difficult and time-consuming as they tend to induce residual stress on the surface. Cutting speed, feed, and other factors impact the machining characteristics and efficacy of difficult-to-cut minerals or metals, particularly Ti. The right choice of cutting and finishing parameters is critical for increased production and widespread use in industries. Because of the increased use of implant finishing and existing issues with traditional machining, innovative methods are necessary to improve the machining properties of metallic implants [273]. The finishing procedure comprises polishing, but only for the outside cylindrical sections; it cannot complete the inside cylinders or intricate geometrical forms of the implants. Similarly, the grinding method of finishing does not allow access to the implants’ deep holes and may leave markings on the final surface. Lapping is a laborious finishing method that is not suitable for intricate geometrical designs of implants. The burnishing procedure causes the completed surface to harden, making it unsuitable for thin-walled surfaces [274]. All limitations of conventional procedures have necessitated the development of improved finishing methods for bioimplants. Figure 20 illustrates the categories of surface post-treatments that may improve the surface characteristics of traditional or additively made bio-implants.

Figure 20. Classification of surface post-treatment for titanium implant materials for enhanced surface properties. Reproduced with permission from [270,271,272].

Figure 20. Classification of surface post-treatment for titanium implant materials for enhanced surface properties. Reproduced with permission from [270,271,272].

Srijan et al. [275] utilized surface mechanical attrition treatment (SMAT), a severe plastic deformation technique, to enhance the surface-dependent and functional properties of β-Ti alloy. The increased defect density resulting from SMAT led to a reduced corrosion rate through the formation of a stronger oxide layer and an improvement in wear resistance. Srijan et al. reviewed and classified recent advancements in surface modification, focusing on the induction of nano-crystallization on metallic biomaterials through various surface modification techniques [276].

6.7. Recent Trends in the Development of Porous Ti Scaffold: Titanium-Based Interpenetrating Phase Composites

As discussed in the previous sections, the major causes of implant failure, like stress shielding, implant loosening, etc., can be mitigated by reducing Young’s modulus by making them porous using different fabrication techniques. However, these modifications invariably reduce the strength and fatigue resistance as compared to their dense counterpart [31,277,278,279,280]. To enhance the mechanical and biological properties of Ti-based scaffolds, researchers explore various techniques to integrate a secondary bioactive phase into the 3D-printed porous framework [281]. Magnesium, zinc, calcium, and iron [44,282,283,284,285], or their alloys, are commonly utilized metallic biomaterials. At the same time, hydroxyapatite (HAp), wollastonite (CaSiO₃) [28], etc., are common ceramic materials that can serve as filler material. The logic behind adding all these materials inside the metallic porous network is to provide mechanical and biological support during the initial implantation period. Since these filler materials are either biodegradable or bioresorbable, they naturally degrade within the human body over time [286]. Zhang et al. [287] developed a Mg-Ti composite via the pressureless infiltration of Mg into an AM-fabricated Ti6Al4V scaffold. Since the constituents of the composite were continuous and mutually interpenetrated in the 3-D space, the composite was termed an interpenetrating phase composite. These porous structures were designed to mimic the bioinspired architecture (Figure 21). The composite shows effective stress transfer, delocalized damage, and arrest cracking. Similarly, Dou et al. [281] fabricated partially degradable interpenetrating phase composites by first creating a pure titanium-based dodecahedral cellular structure with pore sizes ranging from 400 to 500 μm using selective laser melting (SLM), followed by pressureless infiltration of pure magnesium into the porous framework. The developed composite possesses a higher strength (yield strength of ~64 MPa and ultimate compressive strength of ~275 MPa) and lower modulus (~47.3 GPa) than cast pure Mg and Ti. Also, the degradation of the Mg supports the ingrowth of bone tissue and biological fixation between the host and implant. In vitro and in vivo studies of these composites revealed that the composite exhibits accelerated corrosion compared to pure Mg. Still, it remains non-cytotoxic and does not induce obvious adverse reactions following implantation.

In a similar approach, Rahmani et al. [28] reported the addition of three different types of wollastonite (CaSiO₃)-based bio-ceramic via SPS in an AM Ti6Al4V lattice, as shown in Figure 22. The produced composite exhibited higher wear resistance, damage tolerance, and mechanical properties, which can improve the durability of the bones. In addition to Ti6Al4V, TiNi, and Ti22Al25Nb, porous samples infiltrated with wollastonite showed that TiNi has a higher damping capacity, while Ti22Al25Nb induces a higher portion of elastic deformation. In another study, Rahmani et al. [28] reported improved impact and osteoinductivity by manufacturing hybrid wollastonite and Ti6Al4V for craniofacial implants.

7. Current Challenges for Porous Titanium

Porous Ti has the potential to be suited for biomedical applications due to its high strength, lower elastic modulus, longer fatigue life, and good wear and corrosion resistance. In addition to the abovementioned mechanical properties, porous Ti exhibits excellent biocompatibility and helps pass fluids through the porous structures. Nevertheless, challenges are associated with the fabrication and processing of porous Ti implants:

(1)

Porous Ti implants are difficult to manufacture, and optimally controlling pores and maintaining the uniformity of pores are challenging tasks.

(2)

The interconnectivity of the pores is essential in determining the mechanical properties of the implants. It is paramount to maintain a tradeoff between the strength and porosity of porous Ti implants.

(3)

As implants undergo repeated cycles of loading and unloading during daily activities, the porous structures generally exhibit lower fatigue resistance compared to their dense counterpart. The pores serve as potential initiation points for fatigue cracks, which can gradually propagate and ultimately result in premature implant failure.

(4)

There must be a balance between patient-specific implants and large-scale production of implants, i.e., customization and production, since developing porous structures and meshing is time-consuming.

(5)

The implant cost should also be considered, as porous Ti implants are usually costlier than fully dense implants. Higher costs of implants reduce the demand in the market.

(6)

Implementing thorough checking to minimize defect concentration is important to obtain high-quality implants with better mechanical properties and biocompatibility.

(7)

Especially in the case of AM, thermal gradients can lead to variations in pore sizes and porosity levels.

(8)

The post-processing of the porous structures is also very difficult and time-consuming. It must be carried out carefully, as small disruptions can damage the interconnected structures, leading to defective implants.

8. Future Scope of Titanium-Based Porous Implants

Porous Ti has been proven to be the best-suited candidate, especially for orthopedic implants and prosthetics. It has taken the field of biomaterials, specifically implants, by storm and is expected to be the future of biomedical implants, especially orthopedic implants. However, in short, a few porous Ti challenges must be resolved to make it irreplaceable in the market:

(1)

AM has been gaining momentum in producing porous Ti structures. However, more efforts must be taken to fabricate porous structures with good pore interconnectivity, which will, in turn, pave the way for better mechanical properties and biocompatibility.

(2)

A biocompatible coating can be applied to enhance the osseointegration of the implant with body fluids.

(3)

Porous Ti structures can be used as multifunctional implants by integrating them with drug delivery systems or sensors.

(4)

Efforts should be made to make the porous Ti implants more accessible at an affordable price, which is possible through process optimization to achieve higher productivity at a lower cost.

(5)

Investigations of biocompatible joining strategies should be carried out for integrating porous Ti into various biomaterials or metallic implant structures.

(6)

The development of interpenetrating phase composites is limited to the Ti-Mg system since the developed composite has the potential to be utilized as an implant; therefore, a composite system comprising Ti-Zn/Ca/Fe and SS-Mg/Zn/Ca/Fe should be explored.

9. Conclusions

The present review focused on exploring the development of Ti-based porous biomaterial for orthopedic applications. Different characteristics and requirements of biomaterials, like mechanical, physical, and biological properties, are comprehensively reviewed. A section of the review was dedicated to the structure of bone and its physiology for the basic understanding of bone metabolism, which should be considered while designing and synthesizing next-generation orthopedic and dental implants. Considering the importance of porous Ti-based implants for orthopedic applications, this review presents different processing techniques, particularly based on powder metallurgy processes, and the properties of the developed porous implants are compared. Given the importance of the powder metallurgical route for synthesizing porous implants, a detailed study of various powder processing techniques was carried out. Finally, this review concludes by discussing the current challenges and future potential of porous Ti-based alloys for biomaterial implant application.