1.  Introduction

    The organs and tissues in the human body have precise and complex functions. Our lives always depend on the performance of these organs versus their proper functioning. However, these organs and tissues are sometimes damaged and can fail due to diseases or accidents. There are many medical treatments available that can help treat damaged organs, but many of these treatments still lack the ability to repair the organ until it regains its full function. The field of regenerative medicine seeks to provide new tools that repair or replace damaged organs and tissues [1]. Regenerative medicine uses tissue engineering to develop new strategies for repairing damaged organs and tissues. Tissue engineering involves biology, chemistry, and engineering to make new biomaterials compatible with the human body and can be used to repair or replace damaged organs and tissues [2]. Biomaterials have improved significantly since their first development and are still changing as scientists try to better understand diseases and how these materials interact with the body. Biomaterials can take many forms and are produced from many different materials. Ideally, biomaterials should have a porous structure with small holes that allow air, fluids, and even cells to pass through, like the organs and tissues they are intended to treat [3]. Cells aiding in the healing process are loaded into these tiny holes that permeate the biomaterial [2]. In this way, a porous biomaterial can be used to transfer cells to damaged tissue. The biomaterial helps maintain new cells in tissues and, at the same time, is necessary to promote the healing process. Moreover, the porous structure of the biomaterial closely resembles the “exocellular matrix,” which resembles the hooks by which cells in the body are “grabbed” [2, 4].

    Titanium and its alloys are extensively used in biomedical implants because they offer a combination of high strength, lightness, resistance to corrosion, and compatibility with biological tissues. Currently, materials used in these applications include 316L stainless steel, cobalt-chromium (Co-Cr) alloys, and Ti-based alloys (specifically Ti-6Al-4V) [5, 6]. However, these materials sometimes fail after prolonged due to their stiffness relative to bone, inadequate wear and corrosion resistance, and insufficient compatibility with biological tissue. Achieving a critical stiffness matching between orthopedic implants and adjacent bone is crucial in avoiding stress shielding, bone resorption, and implant failure. Despite advances in biomaterials and tissue engineering, ongoing research focuses on developing durable metallic implants due to concerns about the long-term performance of current metallic biomaterials.

   Porous titanium structures have demonstrated immense potential to transform orthopedic and dental implant performance via tailorable interconnected pore networks facilitating bone infiltration and vascularization. It is reported that the elastic modulus of these porous titanium structures is significantly lower, ranging from 91% to 96%, compared to dense Ti alloys [7]. Moreover, porous titanium and its alloys have emerged as promising candidate materials for orthopedic and dental implants owing to their outstanding biocompatibility, corrosion resistance, and mechanical properties closer to natural bone [8-10]. Incorporating of controllable porosity in titanium implants facilitates bone ingrowth, transport of nutrients and/or biofluids, and vascularization, promoting osseointegration for device integration and stability. Subsequently, it can significantly enhance post-implantation healing outcomes and long-term implant lifespan performance. Accordingly, extensive research has focused on engineering Ti implant surfaces and bulk structures with optimized porous architectures. It includes the fabrication of porous coatings on conventionally dense implants as well as the printing of fully porous titanium components using additive manufacturing. Additionally, processing innovations, pore topography, interconnectivity, and gradients influence biological responses [11, 12].

    A prime determinant of in-vivo device integration is the effective pore volume fraction. Studies showed that ≥ 50% porosity enables sufficient bone ingrowth [13]. Wang et al. [14] found that in the context of early-stage osteonecrosis of the femoral head (ONFH) after core decompression, a biogenic trabecular porous Ti rod that is manufactured through the selective laser melting technique exhibited notable quantitative advantages over core decompression alone. The rod group demonstrated significantly higher ratios of bone volume to total volume (BV/TV) at both 3 months (890.0% increase) and 6 months (438.1% increase) compared to the core decompression (CD) group. The histological analysis supported these findings, showing substantial improvements in BV/TV in the rod group, with increases of 881.0% at 3 months and 413.3% at 6 months. Woodard et al. [15] found that hydroxyapatite (HA) scaffolds with multi-scale porosity, specifically microporous (MP) scaffolds with both microporosity (250 - 350 μm) and microporosity (2 - 8 μm), demonstrated superior osteoconductivity, drug-carrying efficacy, and mechanical properties compared to non-microporous (NMP) scaffolds. Thus, it is showcases the importance of scaffold microporosity for bone ingrowth and mechanical behavior in HA implant materials. The profound influence of porous titanium surface topographies on modulating osteoblast cell morphology, adhesion, differentiation, and mineralization also underlines the role of multiple hierarchical porosity. The presented study provides a short review of the current advances in porous titanium-based biomaterials for medical implants. 

    The review explores the impact of pore characteristics on biological and mechanical aspects and summarizes recent progress in processing techniques for constructing optimized porous architecture.

2.Titanium Properties

   The widespread use of Ti in medical implants is grounded in its exceptional biocompatibility, with attributes inertness, a stable oxide layer, and the potential for osseointegration. Its strength-to-weight ratio and flexibility, resembling human bone's elastic modulus, result in robust yet lightweight implants, essential for load-bearing applications such as joint replacements. The adsorption of proteins onto Ti surfaces plays a crucial role in osseointegration, influencing cellular responses and promoting implant integration [16]. Beyond its neutral interaction with the body, Ti actively fuses with bone, ensuring the stability and durability of implants, particularly in dental and orthopedic applications. With minimal ion release, Ti exhibits corrosion resistance inside the human body, avoiding potential harm or allergic reactions. Its non-magnetic nature makes it ideal for patients requiring MRI scans, ensuring safety and compatibility.

    The innate oxide layer grants passive corrosion protection while permitting biofunctionalization that is essential for integration. However, elastic moduli still exceed natural bone, leading to stress-shielding long-term failures. Figure 1 shows the failure causes of implants leading to revision surgery.

    Thus, ongoing research scrutinizes the absolute biocompatibility of Ti alloys, leading to the development of new beta Ti (β-Ti) alloys with non-toxic elements (e.g., Tantalum (Ta), Niobium (Nb), and Zirconium (Zr)). These alloys show promise in improved biocompatibility, strength, and wear resistance, addressing concerns associated with traditional Ti alloys such as Ti-6Al-4V. The quest for materials devoid of cytotoxic elements and with a low elastic modulus continues, fueled by the need for enhanced biomaterials in medical applications. Hence, porous titanium structures better emulate the cellular architecture of bone through an interconnected, open-cell network with adjustable porosity and pore sizes while harnessing the advantages of dense titanium/alloy compositions.

                                                                                

Figure 1. The causes for failure of implants leading to revision surgery.

3.Types of Titanium-Based Alloys

    The biocompatibility and mechanical properties of titanium and its alloys, vital for medical implants, stem from their crystalline structure, transitioning between hexagonal close-packed (hcp) α-and body-centered cubic (bcc) β-phases at an allotropic phase transformation temperature [17-19].

    Titanium alloys are classified into α-, near-α, (α+β), and β-types based on their microstructure and alloying elements [20]. The α-type includes commercially pure titanium (CP-Ti) and Ti alloys, known for exceptional corrosion resistance but limited mechanical strength at room temperature [21, 22]. Near-α Ti alloys, featuring minor β phases, share similar characteristics with α-type alloys but have not been extensively utilized in biomedical applications. (α+β)-type Ti alloys, exemplified by Ti-6Al-4V, dominate biomedical applications due to their superior strength, corrosion resistance, and osteointegration capabilities [20, 23]. Despite their prevalence, concerns about toxic elements such as vanadium have led to the development of alternatives, for instance, Ti-6Al-7Nb and Ti-5Al-2.5Fe [24–26]. However, high moduli in (α+β)-type alloys may pose challenges (see Table. 1), prompting the exploration of β-type Ti alloys with non-toxic β-stabilizers such as molybdenum (Mo), tantalum (Ta), and Zr for improved biocompatibility and suitable elastic modulus [27, 28]. Figure 2 compares the elasticity modulus of different biomedical Ti alloys with human bone, stainless steel, and Co alloys. When comparing (α+β)- and β-Ti alloys to 316L stainless steels and Co alloys, their ultimate strength values are comparable with those of 316L stainless steels but lower than those of Co alloys. However, their yield strengths are also comparable with those of 316L stainless steels but closer to the lower side of the range for Co alloys, as shown in Table. 1.

    However, dense forms of these Ti biomaterials with a misfit Young's modulus can induce stress-shielding, which in turn causes instability at the bone-implant interface. Thus, it results in fibrous tissue ingrowth, disruption of osseointegration, implant mobility, and an inflammatory response necessitating revision surgery. Porous materials have been developed to overcome these issues associated with bulk materials such as Ti. These materials aim to decrease Young's modulus, approaching values similar to bone, thereby improving stress distribution patterns and creating favorable conditions for bone remodeling.

 

Table 1. A comparison of mechanical properties of metallic implants titanium alloys for biomedical applications with human bone, stainless steels, and Co-alloys.

Alloy	Elastic Modulus (GPa)	Yield Strength (MPa)	Ultimate strength (MPa)	Engineering strain (%)	Ref.
Cortical bone	5 − 23	30 − 70	194 − 195	-	[29]
SS 316L	200	200–700	500–1350	10–40	[30]
Co-alloys	240	500–1500	900–1800	10–50	[30]
Ti (unalloyed)	105	692	785	-	[31]
Ti-20Ag	~103	-	-	-	[32]
Ti-6Al-2Sn-4Zr-2Mo-0.1Si	114	990	1010	-	[33]
Ti-6Al-4V	113	999	1173	6	[34]
Porous- Ti-6Al-4V	5.1	-	171.86	-	[35]
Ti-6Al-7Nb	114	880-950	900-1050	8-15	[36]
Ti-5Al-2.5Fe	112	895	1020	15	[36]
Ti-6Al-2Sn-4Zr-6Mo	114	1000-1100	1100-1200	-	[33]
Ti-5Al-2Sn-2Zr-4Mo-4Cr	112	1050	1100-1250	-	[33]
Ti–5Al–1.5B	107	820–930	925–1080	15–17	[37]
Ti-10Cr-20Nb	17	1180	1580	-	[38]
Ti-20 Cr-10Nb	50	980	1590	-	[38]
Ti-20Cr-20Nb	28	1015	1700	-	[38]
Ti–12V–6Sn	57	897	1024	10	[39]
Ti-4.5Al-3V-2Mo-2Fe	110	900	960	-	[33]
Ti-35Nb-7Zr-4Cu	57	1062	1374	-	[40]
Ti-35Nb-7Zr-7Cu	63	1205	1602	-	[40]
Ti-35Nb-7Zr-10Cu	74	1469	1856	-	[40]
Ti-35Nb-7Zr-13Cu	79	1216	1571	-	[40]
Ti- 20Nb	74	-	-	-	[41]
Ti-45Nb	64.3	438	527	-	[42]
Ti-42Nb	60.5	674	683	11.7	[43]
Ti-42Nb	47.9	715	718	17.8	[44]
Ti-27.5Nb	70	800	820	10	[45]
Ti-15Mo	84	745	921	25	[34]
Ti-15Mo	78	544	874	21	[37]
Ti-35Nb-6Ta	50	-	820	~10-12	[46]
Ti-35Nb-15Zr	72.82	1185.18	1199.39	6.7	[47]
Ti–15Zr–10Cr	78	1038	-	-	[48]
Ti-7Mn-10Nb	87	842	1842	~34	[49]
Ti-16Nb-10Hf	81	730–740	740–850	10	[30]
Ti-25Zr-25Nb	60	1025	1588	32.9	[50]
Ti-24Nb-0.5O	54	665	810	-	[51]
Ti-24Nb-0.5N	43	665	665	-	[51]
Ti–35Nb–4Sn	55	-	~470	-	[52]
Ti-26Nb-8Mo	54.5	663	-	-	[53]
Ti–20Nb–1.0Ru	65	920	960	~22	[54]
Ti-15Mo–2.8Nb–0.2Si	83	945–987	979–999	16–18	[37]
Ti-15Mo-2.8Nb-3Al	82	771	812	-	[55]
Ti–15Mo–3Nb–0.3O	82	1020	1020	-	[55]
Ti-23Nb-0.7Ta-2Zr	55	280	400	-	[51]
Ti-25Zr-20Nb-5Ag	61	1544	2184	22.4	[50]
(50Ti-50Zr)77-15Nb-8Mo	96	545	-	-	[56]
Ti-25Nb-5Sn-2Cr	68	314	-	-	[57]
Ti–5Cr–3Au–1Cu	~ 11.2	~520	~670	31	[58]
Ti–25Nb–6Zr–1.7Fe	61	598	1256	15.7	[59]
Ti–35Nb–5Ta–7Zr–0.4O	66	976	1010	-	[55]
Ti-23Nb-0.7Ta-2Zr-1.2O	60	830	880		[51]
Ti-27.5Nb-8.5Ta-3.5Mo-2.5Zr–5Sn	70	826	846	-	[60]

 Figure 2. Elasticity modulus of various dense titanium alloys with a comparison of porous Ti-6Al- 4V.

 

4. Importance of Porosity in Medical Implants

    Porosity is a critical property of biomaterials used for medical implants that enables tissue integration, vascularization, and diffusion of nutrients and waste products [11, 61-63]. Implants made from metals, ceramics, or polymers often aim to mimic the porous architecture of natural bone through engineered surface modifications that create microscopic pores and channels. The size, density, interconnectivity, and orientation of pores within an implanted biomaterial dramatically impact its performance and determine outcomes in osseointegration, infection resistance, and long-term viability.

    For bone implants such as joint replacement prosthetics, dental implants, and fracture fixation devices, porosity facilitates bone ingrowth during osseointegration. Pore sizes between 100-400 microns allow osteoblasts and mesenchymal stem cells to penetrate the implant surface, differentiate, and begin secreting new bone matrix [64 - 66]. Moreover, it has been noted that for an implant to effectively encourage the growth of bone, having an optimal porosity level ranging from 20% to 50% is essential [67]. Highly porous surfaces with three-dimensional (3D) interconnected networks of pores enable more rapid bone integration across the entire implant rather than isolated regions. The degree of porosity can be tuned during manufacturing through extrusion, injection molding, or 3D printing of the bulk biomaterial. Beyond osseointegration, porosity imparts critical advantages in infection prevention and antibiotic delivery for orthopedic implants. Studies found that porous surfaces help protect bones from competing bacterial colonization while allowing the migration of macrophages, lymphocytes, and nutrients to fight infection [68, 69]. Local antibiotic elution is also enhanced by porous channels and reservoirs that increase drug loading capacity [69]. For hip and knee arthroplasties, the built-in porosity serves as a safeguard, where chronic infection can necessitate implant removal.

     The porous architecture that assists short-term bone ingrowth can also determine the long-term stability of an implant by allowing continued diffusion of nutrients and waste transport. Densely calcified, avascular interfaces between bone and implants often lead to fibrous encapsulation, isolating the implant over time. Such starvation causes cell death, loosening, and a potential fracture around the affected region. Implant designs and surfaces that facilitate highly vascular integration through interconnected porosity help prevent this scenario.

5. Titanium Scaffolds Design

    Porous titanium scaffolds for bone implants and tissue engineering applications require careful design considerations across multiple vital parameters. The pore geometry, including overall topology (e.g., spherical, cubical, etc.) and pore size distributions, needs optimization based on mechanical requirements and intended bone and vascular ingrowth behavior. According to Van Bael et al. [70], Ti-6Al-4V scaffolds featuring hexagonal pores exhibited the most significant cell growth, which decreased in scaffolds with rectangular pores and further diminished in those with triangular pores, as illustrated in Figure 3 Hence, such discrepancy is attributed to the higher number of corners and the shorter distance between the two arches in the corners, especially noticeable in hexagonal pores. Consequently, cell bridging occurs more rapidly in hexagonal pores than in rectangular and triangular pores, where the struts are more widely spaced. Despite this, it was observed that the regulation of osteogenic differentiation in the cells was independent of their proliferation, and alkaline phosphatase (ALP) activity increased in triangular pores [70].

     Manufacturing technique constraints, including solid-state foaming, powder sintering, electrodeposition, or additive manufacturing, inform achievable pore geometries [71]. For a given production method, fine-tuning processing factors specifically applied stresses, sintering profiles, laser scanning patterns, and post-treatments enable tailoring final porous structure metrics such as density, surface-area-to-volume ratios, interconnectivity, and anisotropy based on application needs. Robust characterization using scanning electron microscopy (SEM), micro-computed tomography, and related image analysis quantifies the resulting pore morphology down to the micron scale. Then, it feeds back into subsequent design revisions and process parameter improvements for the Ti scaffold architecture. Following these steps in scaffold design, optimized application-specific porous titanium implants can be constructed to promote bone ingrowth.

    (a)                                                                                                                                   (b)