The purpose of tissue engineering scaffolds is to produce customized scaffolds with pre-engineered shapes, structures, and functionalities for improved tissue regeneration [29]. Because 3D printing yields structures that morphologically resemble the multiscale architecture of human tissue, it has been recognized as a useful method for creating tissue engineering scaffolds. Bone tissue engineering, utilizing 3D printed scaffolds, has become a promising approach for bone reconstruction and treatment, overcoming the limitations of traditional bone grafts [3]. Unlike subtractive manufacturing techniques, 3D printing, also known as additive manufacturing, solid freeform fabrication, or rapid prototyping, builds objects from 3D model data, usually by adding material layer by layer [29]. The most commonly used 3D printing techniques that employ solid polymers for product creation include fused deposition modeling (FDM), stereolithography (SLA), digital light processing (DLP), and selective laser sintering (SLS) (Fig. 3). FDM, one of the earliest 3D printing methods, fabricates 3D scaffolds using a thermoplastic polymer filament, which is extruded through a heated nozzle onto a platform where it hardens and sets [30]. Biocompatible and biodegradable polymers such as polylactic acid (PLA), polyglycolic acid (PGA), poly(lactic-co-glycolic acid) (PLGA), and polycaprolactone (PCL) are frequently used for implants. Additionally, many pharmaceutical materials, including cellulose derivatives, polyvinyl alcohol, Eudragit, ethylene vinyl acetate, polyethylene oxide, and thermoplastic polyurethane have been successfully applied in FDM [31]. Laser-based 3D printing systems use the energy of a laser beam or ultraviolet (UV) light to fuse materials. Charles Hull invented the first 3D printing method, SLA, and filed a patent for it in the mid-1980s. SLA, a widely used technology based on vat photopolymerization, selectively polymerizes exposed photocurable resins through various chemical reactions when exposed to light [32]. However, the harshness of UV-based crosslinking, extensive post-processing, inadequate mechanical properties, the potential for liquid resin to become trapped within the finished product, and the lack of biocompatible and biodegradable materials have limited its biomedical applications [33]. The DLP technique, which uses a mask projection method instead of SLA’s approach, offers the advantage of faster production times, essential for treating severe bone injuries. DLP is a light-based technology that accurately produces 3D structures, facilitating the simulation of complex internal bone topologies akin to porous cancellous bone [34]. However, DLP also relies on photocuring, sharing the same drawbacks as SLA, including the harshness of UV-based crosslinking, extensive post-processing, inadequate mechanical properties, and a shortage of suitable biocompatible or biodegradable materials. In the SLS process, a focused laser selectively scans polymer powder material just below its melting point [35]. The primary advantage of SLS for bone tissue engineering is the 3D printing of metals like titanium and ceramics like HA. However, the downsides include the energy consumption and high costs associated with the sintering phase of SLS manufacturing equipment [36]. Moreover, increasing the laser's energy can enhance mechanical strength by melting the material rather than just sintering it.

Bone organoids constructed from bioactive materials are 3D self-regenerating and self-organizing microbone tissues with biomimetic spatial characteristics [37,38]. In order to stimulate bone regeneration, porous bone tissue-engineered scaffolds with the right shape, pore size, porosity, degradability, biocompatibility, mechanical properties, and cellular responses are needed [29]. A material with outstanding biological and mechanical properties is required for use in 3D printing to fabricate scaffolds for bone tissue regeneration. Various types of materials are employed in scaffold construction, including synthetic materials, natural materials, ceramics, and composite materials (Table 1) [39-50]. Due to the limitations of HA and calcium sulfate as bioceramic materials, β-tricalcium phosphate (β-TCP) has become the most widely used [51]. TCP is a bone replacement biomaterial with the chemical formula Ca3(PO4)2, which may be converted at certain temperatures into 2 different α-TCP or β-TCP phases. The crystallinity and compressive strength of α-TCP are stronger than those of β-TCP due to its higher sintering temperature; however, its biological activity is ineffective compared to β-TCP. In addition to having good osteoconductivity and biocompatibility, β-TCP degrades in the body and is 10 to 15 times more soluble than HA under typical conditions [52]. Fairag et al. [40] reported that scaffolds made of pure polymer (poly[lactic acid], 100 M, Lactoprene; Poly-Med Inc., Anderson, SC, USA) and β-TCP, in particular, demonstrated exceptional mechanical properties, biological characterization, and supported bone matrix formation. However, β-TCP has disadvantages such as slow disintegration and an inability to prevent bone resorption [53]. Incorporating trace elements is crucial for bone bonding in the human body. Consequently, a significant area of potential research involves integrating trace metals such as strontium, magnesium, zinc, silver, and silicon into scaffolds to enhance bone tissue regeneration. Liu et al. [41] found that strontium-containing hydroxyapatite (SrHA) could increase osteoblast activity, promote proliferation, and enhance ALP activity. This suggests that 3D printing technology can efficiently produce PCL/SrHA composite scaffolds, which show great potential as implantable materials for bone tissue engineering applications. Additionally, it was observed that a micro/nano-HAp surface layer promoted mouse bone mesenchymal stem cell adhesion, osteogenic differentiation, and the expression of angiogenic genes in human umbilical vein endothelial cells [42]. Zinc (Zn) and its alloys are promising biomaterials with favorable mechanical strength, degradation rates, and biocompatibility [54]. The authors noted that the role of Zn in osteogenesis and osteoclastogenesis is associated with the Wnt/β-catenin and NF-κB signaling pathways, respectively. They also indicated that PCL/Zn scaffolds progressively enhanced osteoclastogenesis with increasing Zn concentration [43].

Polymers are commonly used as synthetic materials for creating scaffolds, with linear aliphatic polyesters such as PGA, PLA, and PLGA being the most widely used [39]. Additionally, biocompatible polymers like polyether ether ketone, PCL, and PLA are utilized in tissue engineering to replace bone tissue due to their biodegradable, non-toxic, non-immunogenic, and non-inflammatory properties [55]. Ledda et al. [50] proposed the use of 3D-printed PLA scaffolds with a microstructure inspired by trabecular bone architecture to support the adhesion, growth, and differentiation of osteoblast-like cells and microporous structures. They demonstrated that these scaffolds mimic natural bone hierarchies and promote increased bioactivity. Moreover, Zhang et al. [44] developed a scaffold that imitates Haversian bone, designed as a multicellular delivery system with osteogenic and angiogenic/neurogenic cells distributed in specific locations. This system actively supports bone tissue engineering and serves as a proof-of-concept for a 3D structure-based co-culture platform. Due to its biocompatibility, biodegradability, and adaptability, PCL, a thermoplastic polymer, is regarded as a polymer material appropriate for the fabrication of porous scaffolds utilized in bone tissue engineering [56]. Rosales-Ibáñez et al. [45] found that grafting gelatin molecules onto PCL, which is inherently hydrophobic and inhibits cell attachment, enhances cell adhesion, and increases cell survival. Ebrahimi et al. [46] reported that coating PCL scaffolds with HA/collagen (COL) creates a bone-like surface that accelerates healing and that surface roughness enhances cell adhesion and proliferation. In our previous study, we used 3D printing and PCL to fabricate a scaffold loaded with rifampicin and cefazolin for osteomyelitis treatment. By employing computer-aided design and manufacturing (CAD/CAM) and printing rifampicin-loaded PCL scaffolds at 80°C, we successfully preserved the antibacterial activity of both drugs [57,58]. Bioactive glass (BG) is a well-known osteoinductive and osteoconductive material that has been combined with biopolymers to create composite scaffolds for bone engineering [59,60]. Distler et al. [49] developed PLA-BG composite filaments for 3D scaffold manufacturing and demonstrated the printability of these filaments, as well as the bioactivity and cytocompatibility of the resulting PLA-BG scaffolds, using pre-osteoblast MC3T3E1 cells. Luo et al. [47] described a novel biodegradable organic-inorganic composite scaffold with tailored porosity, mechanical strength, and excellent bioactivity as a potential candidate for bone regeneration. Various bioink materials have been explored for producing bone-like scaffold additives to demonstrate bone conductivity. However, clinical applications have been only partially successful due to issues such as inadequate bone conduction signals, biomaterial-related infections, suboptimal cell survival, lack of reproducibility, and insufficient patient specificity [48].

An osteoregenerative bioink capable of being printed into complex geometries at high production rates is known as hyperelastic bone (HB) [61]. Due to its distinct biomechanical and biochemical properties, HB bioink can facilitate new bone growth and vascularization without requiring additional biological components [48]. Researchers have demonstrated the ability to 3D-bioprint highly porous, reproducible, and defect-specific bone grafts using HB bioink.