Tissue engineering, a discipline of biomedical engineering, uses a combination of cells, engineered materials and methods, and suitable biochemical and physiochemical factors to restore, maintain, improve, or replace various types of biological tissues. The primary criteria for a polymer to be qualified for tissue engineering applications are its high bioresorbability or biodegradability, high mechanical strength, and enhanced cell attachment ability [18]. Polylactic acid (PLA) is one such qualifier widely explored in bone tissue engineering. Gregor et al. [19] successfully fabricated PLA scaffolds with an average pore size of 350 microns and 30% porosity via FDM-based 3D printing. The authors reported high proliferation rates in osteosarcoma cells with 30% and 50% porous scaffolds while exhibiting mechanical properties to support load-bearing bone growth. Although alginate (Alg) hydrogels show porosity, biocompatibility, and solubility, they are restricted in 3D printing applications due to low mechanical strength, cell attachment, and easy degradation [20, 21]. However, a blend prepared by crosslinking Alg and gelatin (Gel) displayed high 3D printability and cytocompatibility with osteoblasts [22].

Recently, Kim et al. [23] prepared a bioink by combining Alg and silk fibroin (SF) protein to fabricate hydrogel scaffolds using DIW and visible light irradiation (Figure 2A). Due to the increase of cell compatibility through SF, these scaffolds supported the proliferation of fibroblasts with improved cytocompatibility than conventional Alg bioinks, making Alg/SF a promising material for tissue engineering.

In another effort, Lafuente-Merchan et al. [24] employed biopolymers such as Alg, nanofibrillated cellulose (NC), and hyaluronic acid to fabricate NC-Alg-hyaluronic acid scaffolds via extrusion-based AM. Mesenchymal stromal cells were used for cell viability analysis. Adding hyaluronic acid improved the scaffold properties, biocompatibility, and cell viability compared with NC-Alg scaffolds. Chameettachal et al. [25] demonstrated a DIW-3D printed enzymatic crosslinked silk-G bioink as a suitable material for 3D bioprinting of cartilage constructs.

SLM-based 3D printed hydrogel scaffolds containing platelet-rich plasma (PRP)-GelMA have also been studied [26]. Bioactive ceramics, another class of materials used in bone tissue engineering, have been successfully 3D printed to form complex bioceramic parts with dense and porous multifunctional structures [27]. While BJ and SLA have frequently been used to process bioceramics, various AM techniques, such as FDM, DIW, DED, and SLS, have been employed [28–30]. Ceramic-polymer composites have also been successfully 3D printed using polymers like polycaprolactone (PCL), PLA, or polylactide glycolic acid (PLGA) with calcium phosphate (CaP) [31, 32]. The incorporation of polymers improves the processability and flexibility of the fabricated composites during 3D printing. BGs, especially the highly abundant 45S5 composition (45 SiO₂, 24.5 CaO, 24.5 Na₂O, and 6 P₂O₅—in wt%), other melt-derived formulations, and sol-gel derived BGs have also been explored for 3D printing of scaffolds for bone tissue engineering. Recent research by Ma and co-workers [33] exhibited 3D printing of 45S5 BG-based scaffolds using SLA-based AM.

Using the composites prepared with 45S5 BG and tricalcium phosphate (TCP), Bose et al. [34] demonstrated successful 3D printing of scaffolds via BJ. 45S5 BG-based scaffolds have also been 3D printed via the SLS AM technique [35] and DIW [36]. Several studies have also been reported on the successful 3D printing of sol-gel-derived BGs. For instance, Wu and co-workers [37] described DIW 3D printed scaffolds of sol-gel-derived mesoporous BGs combined with polyvinyl alcohol (PVA). The study by Dai et al. [38] illustrated DIW 3D printed Gel/SF scaffolds incorporating sol-gel derived 1 wt% of Cu-doped BG particles for bone defect repair applications. Moreover, polymer-BG composites have also been successfully 3D printed. FDM technique has been successful with BG-based composites prepared from PCL [39], PLA [40], poly(hydroxybutyrate-co-hydroxyvalerate) (known as PHBV) [41], and polyolefin binders [42] while DIW has been mostly used with silk [43].

Contribution from 4D printing towards tissue engineering has also been evidenced in recent literature [44, 45]. Live cells containing bioinks have been 4D printed, enabling the fabrication of functional tissues for successful organ transplantation and efficiently repairing damaged tissues [16]. Gugulothu and Chatterjee [46] demonstrated a successful 4D printed bioink consisting of a blend of GelMA and PEGDM with a photoinitiator and a photoabsorber via DLP technique. This 4D bioprinted material acts as a shape-morphing and cell-laden hydrogel for tissue engineering applications by supporting cell viability and proliferation while altering its shape upon hydration, a cell-friendly stimulus (Figure 2B and C) [46]. Díaz-Payno and co-workers [47] recently prepared an extrusion-based 4D printed smart multi-material system using two hydrogel-based materials, hyaluronan and Alg. This scaffold self-bends upon differential swelling between the two zones, mimicking the natural cartilage structure. Ding et al. [48] showed a DIW-based 4D bioprinting to produce a shape-morphing cell condensate-laden bilayer system. The technology facilitates the creation of tissue constructs with precise control over cellular organization and distribution and would pave the way for future research on scaffold-free tissue regeneration.

Using ADA and Gel, Kitana et al. [49] exhibited DIW-based 3D printed tubular structures that self-transformed into a T-junction after immersing in water, hence showing 4D printing behaviour. The transformation of the 4D printed crosslinked ADA/Gel, into a tubular T-junction after swelling in pure water is depicted in Figure 2D [49]. Human endothelial cells seeded on the T-junction showed outstanding growth properties and excellent cell viability. This finding could pave the way for future vascularized tissues for the survival and function of larger engineered organs. Furthermore, 4D printing has become promising for fabricating patient-specific, functional organs that can be adapted and integrated within the recipient’s body [50].

Drug delivery refers to a broader scientific field involving various approaches, formulations, manufacturing techniques, storage systems, and technologies that transport a pharmaceutical compound to a specific target site to obtain a desired therapeutic effect. Polymers such as PLGA, PCL, and other materials like CaP ceramics, BGs, bioactive ceramics, and ceramic-polymer pastes have been explored in this field. BJ AM technology has been widely utilized to fabricate BMMs in this field. VP, PBF, and material extrusion AM techniques have also been used to 3D print relevant scaffold structures. During drug delivery, porous scaffold structures are first 3D printed from the relevant material, followed by the loading of the drug. This strategy avoids the degradation of drugs upon high-temperature processing. There has been extensive research on processing, mechanical property measurements, and biocompatibility evaluations in vitro and in vivo of many CaP ceramic scaffolds.

Ceramic-polymer composites have been 3D printed successfully. Adding polymers such as PCL, PLA, and PLGA into CaP ceramics improves the processability and flexibility of the printed part [31, 32]. The BJ AM technique enables the delivery of heat-labile molecules such as growth factors and antibiotics by fabricating low-temperature CaP-based scaffolds [28]. Controlling pharmacokinetics is essential in drug delivery applications. Inzana et al. [51] described that a PLGA-based post-printing coating of CaP scaffolds achieved first-order drug release kinetics over 14 days. Drug/growth factor-loaded composite CaP scaffolds could be successful via extrusion-based AM under low temperatures and mild post-processing conditions. Mineralized slurry or paste compositions, extrudable under physiological temperature, become ideal for this approach [28]. Martínez-Vázquez et al. [52] described successful DIW-based 3D printing of porous silicon-doped hydroxyapatite (HASi) and Gel composite scaffolds for delivering vancomycin antibiotics. Mild scaffold fabrication conditions maintained the antibiotic’s antimicrobial activity in standard in vitro assays. In a similar DIW-based approach, Akkineni et al. [53] demonstrated the fabrication of vascular endothelial growth factor (VEGF) or bovine serum albumin (BSA) loaded CaP-based scaffolds. The mechanical properties of these scaffolds were comparable to those of trabecular bone and showed biocompatibility with mesenchymal stem cells for up to 21 days. Poldervaart et al. [54] exhibited a rare in vivo effort to 3D print composite macroporous Alg scaffolds via extrusion, laden with Gel microparticles (GMPs) and mesenchymal stem cells. However, concentrations greater than 3% w/v Alg could not be extruded due to the high viscosity factor of the printing ink.

On the other hand, high-temperature scaffold fabrication or processing enables CaP-based ceramics to achieve improved mechanical properties. However, high-temperature processing lacks uniform printability of cells and bioactive molecules. As a precaution, additional post-processing routes can be taken, such as the incorporation of bio-factors and cells onto the printed structures, including surface adsorption or surface modifications, regardless of the 3D printing technology [28]. Moreover, doping CaP with silicon improved both the bioactivity and mechanical properties of these scaffolds. Combined with bone morphogenetic protein (BMP) in the form of recombinant human BMP-2 (rhBMP-2), these scaffolds enabled bone ingrowth, osseointegration, and vascularization. Ishack et al. [55] fabricated biphasic CaP [15% hydroxyapatite (HA) and 85% β-TCP] scaffolds via extrusion-based AM. After loading with either BMP-2 or dipyridamole and implanting into a mouse calvarial defect, they promoted bone regeneration eight weeks post-operatively. Koski et al. [56] demonstrated the use of naturally sourced gelatinized starch as a natural binder system with HA ceramic to obtain extrusion-based solid-freeform fabricator (SFF) scaffolds. These scaffolds showed improved compressive strength and in vitro biocompatibility with osteoblast cells without crosslinking or post-processing.

4D printing is applied in numerous advanced drug delivery systems to improve efficiency in treatment outcomes. Controlled release of drugs, patient-specific dosing, and targeted delivery are the main advantages of these 4D printed structures compared to conventional systems [57]. Researchers can fabricate devices that release drugs at a precise rate over a predetermined period by utilizing the ability of the 4D printed scaffolds/structures to respond to specific stimuli like temperature and pH changes [58].

Tran et al. [59] devised 4D printed smart hydrogel systems that respond to thermal, magnetic, electrical, photo, pH, and water stimuli. For example, a pH-responsive hydrogel capsule releases its content gradually in an acidic environment of the stomach. The authors reported that this controlled drug-releasing strategy is highly beneficial for minimal side effects and improved therapeutic efficacy. Moreover, AM techniques such as SLA, DLP, two-photon photopolymerization (2PP), and extrusion have successfully fabricated these hydrogels. Cancer treatment is another area where 4D printed personalized drug-eluting implants have recently become a highly versatile technique. Upon responding to external stimuli like pH changes or biomarkers, these implants could release chemotherapeutic agents at a controlled rate, enabling precise and timely distribution of drugs to specific sites, minimizing side effects and enhancing the overall efficiency of the treatment [60]. Makvandi et al. [61] prepared a 4D printed microneedle patch for personalized pain management. This BMM could release analgesic drugs in response to inflammation/pain signals. Importantly, 4D printing approaches ensure effective and efficient drug delivery with minimal side effects.

Some of the recent 3D and 4D printed BMMs targeted for drug delivery applications are depicted in Figure 3. These include PLA/PVA-based FDM-3D printed mouthguard (Figure 3A–D) loaded with food-grade flavor vanillic acid (VA) and clobetasol propionate (CBS) model drug [62], polymer/carbon-based magnetoelectric responsive porous nanocookie conduit 4D printed via DLP (Figure 3E) [63], and ethyl cellulose/hydroxypropyl methylcellulose (HPMC)/polyvinyl pyrrolidone (PVP)/cellulose acetate-based controlled drug release shell (Figure 3F and G) 3D printed via pressure-assisted microsyringes (PAM) technology [64].

Medical implants are devices placed in or on the body surface. The most common implants are prosthetics intended to replace a damaged or missing body part. Apart from prosthetics, other implants deliver medicines to internal organs and tissues, support internal structures and monitor body functions. Implants can be made from biological materials such as bones, tissues, skin, metal, plastic, ceramic, and other composite materials.

AM is one of the most used techniques for manufacturing medical implants. 3D printing is preferred over conventional implant manufacturing methods mainly due to the ease of manufacturing a customized implant with enhanced compatibility and clinical results. 3D printed implants are popularly used in reconstructions of the spine, shoulder, and hip and for facial surgery and dental implants [90]. Patient-specific implants and prostheses are fabricated using a wide range of medical-grade metallic, ceramic, polymer, and composite materials [91]. Metallic biomaterials are widely used in the medical field due to their superior mechanical properties and long lifetime. 3D printed metallic implantable medical devices are commonly made with alloys of Ti, Co-Cr-Mo, and stainless steel (SS) [92] using DED and PBF AM techniques [93]. Commercially pure Ti and its alloys (Ti-6Al-4V, Ti-6Al-7Nb, Ti-5Al-6Nb, and Ti-13Nb-13Zr) are the commonly applied bone implants due to low density, lightweight, and suitable tribological and mechanical properties [93]. Ti-based alloys possess the highest biocompatibility than any other metallic content, but they are still considered bioinert materials compared to bioceramics [94, 95]. Recently, many research attempts have focused on improving the quality of Ti implants, aiming for enhanced biocompatibility, osseointegration, and antimicrobial properties. Many studies have attempted to improve biocompatibility by surface modifying the 3D printed Ti implants. Some of the surface modifications on Ti scaffolds include the application of a homogeneous layer of microporous TiO₂ and calcium-phosphate [96], genetically modified elastin-like recombinamers (ELRs) containing specific cell adhesive (RGD) and osteoinductive (SNA15) moieties [97], coating of aspirin (ASP)/PLGA [98], titania nanotubes via electrochemical anodization and bioactivation through HA coating [99], and chimeric peptides [100]. Studies on adding antimicrobial properties were carried out by surface coating of the Ti implants with gallium nitrate [101], vancomycin hydrochloride [102], flavonoid quercitrin [103], chitosan (CS)-modified MoS₂ coating loaded with AgNPs [104], and calcium titanate [105], to prevent bacterial adhesion and proliferation on the surface. Today, biodegradable metallic implants such as magnesium alloys are gaining popularity as promising alternatives for metallic permanent prostheses. These biodegradable magnesium implants degrade gradually over time, matching the healing rate of surrounding bones and transferring the load back to the healing bone. Due to transparency towards X-rays, Mg-based implants do not interfere with radiographic techniques, allowing efficient monitoring of the implant. Further, Mg-based implants do not need to be surgically removed, preventing risks associated with additional surgical procedures [106, 107]. Despite these beneficial properties, Mg-based implants are also associated with unfavorable characteristics, such as granular tissue formation around implants and rapid degradation of the implant before the bone heals, preventing their wide applicability. Recently, numerous attempts have been made to avoid implant degradation by surface modifications [107].

Ceramics are also used for making medical implants and can be categorized into bioactive ceramics (bioglass, HA wollastonite, phosphates) or bioinert ceramics (alumina, zirconia and titania) and composite materials [108]. 3D printed bioceramic scaffolds such as akermanite (Ca₂MgSi₂O₇, AKT), HA, β-TCP, and BGs have been employed in creating multifunctional implants for osteosarcoma treatments. These implants may function as bone substitutes, filling the space and facilitating attachment, proliferation, and differentiation of bone cells, promoting bone regeneration. These scaffolds can be further functionalized by adding anti-tumor functional agents, nanoparticles, and even engineered microbes to display additional functions [109–111]. Ceramic biomaterials such as CaP, halloysite, alumina and zirconia contain many applications in dentistry. HA is one of the optimum ceramic materials for dental implants due to its excellent biocompatibility [112]. However, due to high elasticity modulus, HA is brittle and often used as a coating associated with other materials [113]. Zirconia is another bioactive ceramic material commonly used for dental applications due to higher biocompatibility, suitable mechanical and tribological behavior, less dental plaque production, and resistance to staining [114]. Zirconia has been used extensively for the 3D printing of dental implants and prostheses, and a recent review by Branco et al. [115] summarized the recent advances of 3D printed zirconia-based dental materials.

Polymers are a diverse group of natural or synthetic materials with favorable mechanical and physicochemical properties for applications in the medical field. Easy processing, low cost of production, compatibility with multiple 3D printing techniques, and the possibility of modifications are some of the advantages associated with polymers. 3D printing techniques such as FDM, SLA, SLS, and DIW are commonly used 3D printing methods for polymers [111, 116]. Synthetic polymers used in medical applications can be categorized into biodegradable and non-biodegradable polymers. Among the non-biodegradable polymers such as polymethyl methacrylate (PMMA), polyether ether ketone (PEEK), and polyether ketone ketone (PEKK) have all been applied in the preparation of medical implants via AM [116–118]. PMMA is a commonly used polymer for orthopedic and bone grafting implants, with specific applications for fixing orthopedic prosthetics in the shoulders, knees, and hips. However, PMMA-based bone cement has many disadvantages. Its limited interactions with the bone and non-biodegradability have prevented it from extensive usage as an implant material [119]. In a recent study by Chen et al. [120], an embedded 3D printing methodology combined with a special post-curing technique showed the potential to enhance the future fabrication of patient-specific, complex, and functional PMMA-based implants. PEEK is another leading high-performance thermoplastic organic polymer commonly used to produce medical tools, implants and prostheses via AM [121]. PEEK possesses many favorable characteristics, such as good mechanical properties, temperature stability, high wear resistance, low coefficient of friction, high processing capability and excellent biocompatibility making them suitable for a plethora of medical applications [121, 122]. One of the main advantages lies in its modulus of elasticity being similar to that of human bone, making it a suitable candidate for cranial, orthopedic, trauma and spinal implants via AM-based techniques [123]. However, the bioinertness of PEEK hinders the bone attachment to the implant surface, resulting in poor osseointegration. Recently, there have been many attempts to improve the bioactivity of PEEK by incorporating HA and using other binder agents, such as PLGA, to load the PEEK surface with other beneficial compounds to provide favorable features [123]. Among biodegradable synthetic polymers, polyglycolide acid (PGA), PLA, PLGA, and PCL have been used in the manufacturing of medical implants [119]. Both PGA and PLA are commonly used for biodegradable screws, nails, and plates to fix orthopedics. However, the wide application of these biodegradable polymers is limited due to the rapid degradation properties of PGA and intrinsic brittleness, poor toughness, and a slow degradation rate of PLA [119, 124]. de Oliveira and co-workers [125] successfully 3D printed a PLA-based interference screw via the FDM technique. This device has shown an excellent tendon-to-bone fixation comparable to its Ti counterpart, with promising results.

4D printing technology has contributed remarkably to fabricating smart implants and prosthetics. Its ability to respond to specific stimuli, such as temperature, moisture, light, and magnetic fields, has yielded customizable and adaptable implants and prosthetics. Khorsandi et al. [126] covered recent contributions of 4D printing towards dentistry and maxillofacial surgery in fabricating relevant implants. With the ability to perfectly fit into the oral structure of the patient and also to adapt changes in the jawbone over time, these implants offer several advantages, such as optimal functionality, reduced discomfort, and improved patient satisfaction. 4D printing contributions towards orthopedics surgery should also be acknowledged. Customizable patient-specific implants fabricated via 4D printing ensure a precise fit and diminish complications. Zamborsky et al. [127] illustrated the applicability of 4D printing in manufacturing blood vessels, tissues, intelligent bandages, and efficient wound healing via 4D printed latticework. With the ability of these 4D implants to be adjusted to the body changes of patients with time and advancements in artificial intelligence (AI) technologies such as robotics, satisfied recovery and repair, a key goal of precise orthopedics could be achieved [128]. Lin et al. [129] demonstrated successful FDM-based 4D printing of biomimetic intestinal stents using shape memory biocomposites. The design was based on wavy biomimetic networks mimicking the nonlinear stress-strain nature of biological tissues. High flexibility, facilitation of reduced irritation of the intestinal wall, biodegradability, and near-body-temperature (NBT) triggered nature of these 4D printed stents are considered next-generation intelligent implants.

Zhou et al. [130] introduced 4D printed shape-memory vascular stents of βCD-g-PCL, altering their shapes in response to fluctuations in blood flow or vessel diameter. These properties not only supported affected blood vessels but also minimized complications and additional surgical interventions for the stent, enhancing the efficiency of the treatment. Previous literature also showed the successful fabrication of 4D printed spinal implants that could gradually alter the shape supporting the spine as it heals. This could potentially minimize complications and support efficient recoveries [131, 132].

Some of the recent 3D and 4D printed BMMs tested for implants and prosthetics, reported in the literature are showed in Figure 5. The ones illustrated represent a variety of BMMs including Ti-6Al-4V-based porous channel dental implants 3D printed via DMLS (Figure 5A) [133], cross-linked PLA-based thermomagnetic responsive vascular stent 4D printed via DIW (Figure 5B) [134], PCL/acrylates-based thermo responsive vascular conduit 4D printed via DIW (Figure 5C) [135], FDM-3D printed acrylonitrile butadiene styrene (ABS)-based human skull and PEEK based porous implant applied on the skull (Figure 5D) [136], and FDM-3D printed PLA/antibiotic based interference fixation screws (Figure 5E) [137].

The above discussed most recent 3D and 4D printing techniques, materials, and printed BMMs for applications in tissue engineering, drug delivery, surgical and diagnostic tools, and implants and prosthetics are summarized in Tables 1 and 2.