Recent Trends on Additive Manufacturing Biomaterial Composites in Tissue Regeneration Future Perspectives,Challenges,and Road Maps to Clinics for Biomedical Applications—A Review

Iron composites

Iron (Fe) is regularly used for healthcare purposes owing to its ease of manufacture and mechanical reliability, as well as its high fracture strength, ⁵⁸ which are biodegradability and nontoxic. Biodegradability is a key feature of biomaterials with long-standing causes of inflammation in the body. ⁵⁹ Alloying Fe makes it degrade quicker and more evenly. These techniques might also produce mechanical characteristics similar to the insertion site. ⁶⁰ In recent years, FeMn and FeMnPd alloys exhibiting improved degradation rates and mechanical properties comparable with those of 316L SS have been developed for stent applications. These alloys are more suitable for use as an orthopedic implant because of their nonmagnetic properties, which provide compatibility with nuclear magnetic resonance and MRI analyses. ⁶¹ The addition of Fe-based alloys causes an increase in dramatic degradation, which is twice as fast than pure Fe. ⁶² Layers of calcium phosphate and oxide shield the implant from the body. These layers slow down implant degradation by limiting oxygen access. ⁶³

Technical specifications in AM of the Fe composites

Fe-based alloys can be processed using laser power bed fusion (LPBF) processes such as electron beam melting (EBM) and selective laser melting (SLM). To achieve the appropriate microstructure and mechanical characteristics, it is necessary to adjust the technical aspects, ⁶⁴ including laser power, scanning speed, and layer thickness. To achieve consistent characteristics and reduce residual stresses, it is crucial to maintain control over thermal behavior throughout the printing. ⁶⁵ Balancing energy input and powder layer thickness to make high-density Fe-based products without porosity is difficult with PBF. Lack of energy can lead to layer bonding issues, while excess energy can induce vaporization or oxidation. ⁶⁶ Fe’s high melting point (∼1538°C) requires more energy, leading to thermal strains and distortions. The study of tissue demonstration SLM-made Fe-based scaffold offers high mechanical strength and slow biodegradation, causing it suitable for bone tissue creation. Abudeen et al. found that SLM-created porous Fe scaffolds enhanced bone ingrowth and had acceptable degradation rates for bone regeneration. ⁶⁷ Binder jetting (BJ) involves selectively depositing the binder over the bed containing Fe powder, which can be used to manufacture green components. ⁶⁸ BJ is a technique that bonds Fe powder particles using a liquid binder. It is effective in tissue engineering, where the binder saturation level is fine-tuned to 80–95% for optimal strength. However, it faces challenges such as controlling shrinkage and oxidation during postprocessing. BJ has been used to create porous Fe scaffolds for bone regeneration, demonstrating promising biodegradability and osteoconductivity; this technique makes it possible for the creation of complicated shapes and is quicker than the other AM techniques. Directed energy deposition (DED) is the process of depositing molten metal onto a substrate to produce the desired structure, ⁶⁹ ideal for repairing or enhancing existing parts, and the ability to utilize a diverse range of feedstock materials, such as Fe-based powders or wires. DED methods fail to regulate microstructure owing to high cooling rates, resulting in residual stresses and inferior material characteristics. For tissue engineering, Fe biomaterial composites’ mechanical integrity and biodegradation rate might be affected. ⁷⁰ For biomedical implants, DED’s rough surface finish might require substantial postprocessing. Degradable implants are made from Fe–Mg composites using DED. Tunable porosity and degradation rates were achieved by Liu et al. using DED to make Fe scaffolds with biodegradable polymers, biocompatible and mechanically efficient scaffolds for tissue engineering. ⁷¹

Material extrusion (ME) entails layer-by-layer extrusion of an Fe-based filament to create the final product. It is widely used for prototyping and developing personalized implants. In comparison with other AM approaches, it is quite affordable and user-friendly. ⁷² This process is used to create Fe biomaterials by extruding Fe-polymer composite filaments. The process involves adjusting parameters such as extrusion temperature (220–250°C), layer height (0.1–0.3 mm), and print speed (30–60 mm/s) to ensure good bonding and prevent delamination. Postprocessing steps such as debinding and sintering are necessary to remove the polymer and densify the Fe structure. ⁷³ However, ME faces challenges in maintaining Fe powder homogeneity within the polymer matrix, resulting in uneven mechanical properties and unpredictable degradation behavior. ME has been explored for the fabrication of porous Fe scaffolds. Gorejova et al. utilized ME to produce biodegradable Fe-phosphate scaffolds with controlled porosity, demonstrating their potential for supporting cell growth and tissue regeneration in bone engineering. ⁷⁴

AM applications of Fe composites

3D-printed Fe with HA-coated scaffolds provides a boost to the development of stem cells in blood marrow. ⁷⁵ Intramedullary nails are used as implants. Fe intramedullary nails outperform steel in vivo. Also, provide a boost to leukocyte numbers and differentiation. ⁷⁶ Most Fe-based biomaterials enhance the focus of study on mechanical, electrochemical, and cell viability. There is no clarity about the effect of biomaterial in vivo. The impact of 3D printing on structure still needs more study. Current testing indicates that Fe 3D printed in vivo is feasible, but more research is needed. ⁷⁷ The extrusion-based 3D printing of porous Fe structures improved breakdown rates, prevented stress shielding, and encouraged cell growth. In vitro, porous Fe scaffolds deteriorate faster than pure Fe. ⁷⁸ 3D printing may have caused “sintered powder particle boundaries with a micropore network.” Electrochemical and immersion studies show that corrosion products reduce scaffold deterioration. ⁷⁹ There is a need for further research on the subject of a decrease in the corrosion of products. It can develop on the specimen’s surface and stop any further rusting. Polymeric composition, on the contrary, can moderate the corrosion rate and improve the scaffold’s biological performance. Polyethylenimine (PEI) has been investigated for decades and has found a niche in a variety of biological applications. Twelve weeks of corrosion are depicted in Figures 8 and 9 as cross sections of Fe and Fe-PEI samples.

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FIG. 8.

Before corrosion-metallographic cross sections of (A, E) Fe, (B, F) Fe-PEI1, (C, G) Fe-PEI2, and (D, H) Fe-PEI3. ⁸⁰ Reprint with permission.

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FIG. 9.

After 12 weeks of metallographic cross sections of (A, E) Fe, (B, F) Fe-PEI1, (C, G) Fe-PEI2, and (D, H) Fe-PEI3. ⁸⁰ Reprint with permission.

The Fe AM biomedical applications require the biocompatibility of Fe-based alloys manufactured via AM methods. Frequently, postprocessing operations are required to minimize irregularity and enhance surface quality, developing rigorous quality control procedures to ensure the printed component’s consistency and dependability, and comprehending and enhancing the mechanical characteristics of the Fe components that are printed to fulfill the particular demands of biomedical applications. Research gaps in Fe-based scaffolds include achieving the right balance of porosity and mechanical strength, as current SLM processes often lead to irregularities in pore size, affecting biological integration and vascularization. ⁸¹ Postprocessing effects on scaffold properties, such as annealing or surface modification, are insufficiently studied. Controlling the degradation rate of Fe-based scaffolds is a key challenge, as Fe degradation is slower than desired for tissue regeneration, requiring further research on alloy compositions or coatings. Existing literature shows that pore size and degradation behavior are crucial for scaffold performance, but there is a lack of research on how to achieve controlled porosity while maintaining mechanical integrity in Fe composites through SLM. ⁸² The BJ, it is difficult to get high-density parts without substantial postprocessing methods, technique used for tissue engineering has several research gaps. The postprocessing process, such as sintering, is crucial for achieving mechanical integrity, but its effects on the porosity, strength, and degradation rate of Fe-based composites are not well understood. In addition, the compatibility of Fe powders with binding agents is understudied, affecting the final properties of the scaffold. Despite limited research on Fe composites, studies on BJ for metallic scaffolds have mainly focused on Ti or SS. Only limited research has investigated Fe composites, leaving a significant gap in understanding how BJ can be optimized for Fe-based scaffolds. Addressing this gap could expand the application of this technique for biodegradable implants. ⁸³

DED: Preventing, cracking, and achieving the correct microstructure require precise control of temperature gradients during deposition. Comparing this AM approach with others, getting high-resolution characteristics might be challenging. ⁸⁴ Research gaps in the field of Fe composites include the biocompatibility of EBM-produced Fe composites, the impact of surface roughness on cell interaction, and the specific interaction between the Fe component and biological tissues. ⁸⁵ While EBM offers superior control over a microstructure, there is limited research on the specific interaction between EBM-produced Fe composites and biological tissues. In addition, the impact of surface roughness on cell attachment and proliferation in EBM scaffolds is underexplored. ⁸⁶ Addressing these gaps could improve the development of scaffolds that promote better tissue integration.

ME: Creating high-density parts with adequate mechanical characteristics might be difficult, due to its low resolution in comparison with other methods. It is only suitable for specific biological applications. ⁸⁷ Overall, Fe-based AM technology scaffolds face challenges in biodegradation control, mechanical versus biological properties, and surface modifications. Fe degrades too slowly in the body, hindering tissue regeneration. Addressing this issue is crucial for biodegradable implants to reach their full potential in clinical applications. Mechanical strength and biological performance trade-offs exist in AM-fabricated Fe composites, but optimizing biological aspects such as cell attachment, proliferation, and tissue ingrowth is an unresolved issue.^88,89 Bridging this gap could enhance Fe composites’ performance in tissue engineering. Surface modifications, although explored in some studies, are not comprehensively studied. Addressing this gap could lead to breakthroughs in scaffold design, creating structurally sound and biologically active scaffolds.

Stainless steel composites

The association of Fe and carbon with at least 11% chromium makes SS good for orthopedic implants with low cost and great mechanical, thermal, and biological compatibility. It is an issue that SS may outperform bone due to its capability to shift implants due to stress shielding. ⁹⁰ However, bone regeneration will take a long time once the metal corrodes. Corrosion may occur on several metals. Some patients are allergic to metal toxicity and leaching. Assuming metals in the teeth or oral cavity do not look very aesthetic, there is a need for polymeric biomaterials. It is possible to improve implant radiopacity through the use of radiopacity coatings or alteration to SS by alloy composition. Alloying other metals with SS may make it more biocompatible, and it has several varieties and grades for corrosion or oxidation resistance. SS alloys are very popular as there is no need for coating, treatment, or painting and with the ability to survive many years. ⁹¹ Combining chromium, carbon, and nickel into a material makes it more resistant to corrosion and increases its ability to withstand high temperatures. ⁹² It might be beneficial in sterilizing factors. SS may be enriched with anticorrosive elements such as molybdenum and other metals such as nitrogen and aluminum. Chromium oxide coating acts as a barrier to stop oxygen from getting into the steel and causing corrosion. Nitrogen, nickel, and molybdenum are other anticorrosive materials. Manganese may be added to steel to strengthen it and increase its durability. ⁹³ The most popular biomedical SS is 316L. It has nickel content but ample nitrogen by enriching nitrogen to form an austenitic structure, which enhances its mechanical characteristics. Hence, the low nickel content protects the body from the damaging effects of nickel ions. In addition to its high strength and ductility, 316L SS is highly biocompatible and resistant to corrosion and wear. Researchers indicate that a protective combination may be used in AM to maintain SS alloys against corrosion. ⁹⁴ 18-8sMo SS was the first anticorrosion implant material. There are various types of SS alloys, but SS 316L is a prominent alloy most widely used for medical purposes. ⁹⁵ Large quantities of this metal alloy are used in surgical tools. Its antibacterial and antistaining characteristics are important as a medical disinfectant. Several additional medical devices use SS, such as sensor probes, artificial heart valves, needles, bone fixations, syringes, orthopedic implants, and catheters for otolaryngology (ear surgery).

Technical specifications in AM SS composites

In the biomedical field, PBFs such as EBM and SLM are frequently used for SS. SLM is effective for processing metal-based composites such as SS due to its precise control over material properties and microstructure. ⁹⁶ The optimal conditions include a laser power between 200 and 400 W, a scanning speed of 0.5–2 m/s, a layer thickness of 20–50 µm, and an elevated build temperature (200–300°C) to reduce thermal stress and cracking, especially for SS composites. ⁹⁷ In SS composites, fast solidification causes high residual stresses during cooling, causing deformation or fracture. Heat treatment relieves tension after processing. SLM can make high-density components, however, SS microstructures are difficult to fabricate and need accurate laser settings. Tissue engineering scaffold research found that fine-tuning laser power and scan speed reduced porosity, increased mechanical strength, and improved biocompatibility in 316L SS composites. ⁹⁸ EBM is an advanced material metallurgy technique that is ideal for producing SS composites. It operates under specific conditions, such as a beam current of 5–20 mA, a build temperature of 600–800°C, and a vacuum environment to minimize oxidation. ⁹⁹ However, EBM faces challenges in surface finish, as it produces rougher surfaces compared with SLM, which may require postprocessing. In addition, its vacuum environment limits its compatibility with other heat-sensitive biomaterials, particularly those that are heat sensitive. For instance, EBM has been used to fabricate 316L SS scaffolds for load-bearing applications in bone tissue engineering. ¹⁰⁰

The biomedical applications of DED are broad due to its in situ alloying and multimaterial deposition capabilities. DED is a versatile tool for fabricating large parts and repairing existing components, making it ideal for tissue engineering implants using SS composites. Key parameters include laser power (500–1000 W), feed rate (10–30 g/min), and shielding gas flow (argon gas). However, DED faces challenges such as anisotropy, which can affect the reliability of SS tissue scaffolds, and surface roughness and precision, The deposited material layers are thicker compared with SLM and EBM, which require postprocessing for smoother finishes and higher precision. ¹⁰¹ For instance, DED was used to create porous 316L SS structures, demonstrating its potential as customizable scaffolds for tissue engineering, particularly in bone regeneration, due to its mechanical strength and corrosion resistance. ¹⁰² BJ is an AM technique that selectively deposits a binding agent onto a powder bed, creating complex geometries without high heat. ¹⁰³ The optimal conditions for SS composites include layer thickness of 50–100 µm, binder saturation of 70–80%, and postprocessing of sintering at 1200–1300°C for densification and removal of the binder. However, these parts may have lower mechanical strength compared with SLM or EBM, and the postprocessing complexity requires precise control over heating rates to prevent warping or defects. ¹⁰⁴

AM applications of SS composite

Researchers have used preosteoblast cells to ascertain the biocompatibility of SS alloys produced using 3D printing technology. Neither cell adhesion nor cell proliferation is affected by the 3D-printed SS alloy, high-strength, lightweight, and corrosion-resistant 316L SS alloy, made by selective laser sintering (SLS) 3D printing. Ethylene vinyl acetate copolymer coating was applied to the metallic powder. ¹⁰⁵ 3D printing yield and elastic strength is seen as similar to that of human bone. Laser energy density and sintering temperature play a vital role in porosity and mechanical characteristics. ¹⁰⁶

Gap consideration in AM SS composite

It is still difficult to get the best mechanical properties while maintaining biocompatibility in the SS SLM. ¹⁰⁷ Research gaps in surface roughness and porosity control during (SLM) are significant. Controlling surface roughness and porosity is crucial for cell adhesion, proliferation, and differentiation. However, achieving consistent surface properties remains challenging. In addition, the use of SS alloys containing chromium and nickel can induce cytotoxicity over time. ¹⁰⁸ To mitigate these risks, alternative alloy compositions or surface treatments are needed. Existing studies show that SLM can create scaffolds with tailored mechanical properties, but inconsistent porosity can lead to suboptimal biological responses. ⁹⁶ Addressing these gaps could lead to more reliable, biocompatible SS scaffolds with improved osteointegration. Improving surface roughness and porosity control could enhance cell behavior and mitigate long-term toxicity risks, resulting in safer, more effective implants and scaffolds for bone tissue engineering. ¹⁰⁹ Research gaps in the production of SS scaffolds for tissue engineering include microstructural control, postprocessing requirements, and the need for optimized EBM processes. Despite the advantages of EBM, controlling the microstructure of SS composites during processing remains an unresolved challenge as undesirable phases or grain structures could affect the scaffold’s mechanical properties and biological performance. ¹¹⁰ In addition, extensive postprocessing is required, increasing production time and cost. Studies by Smith et al. have shown that EBM can produce high-strength components, for tissue engineering applications. Addressing these issues could lead to optimized scaffolds with better mechanical properties and better biological outcomes, making AM of SS composites more commercially viable for tissue engineering. ¹¹¹ BJ, a cost-effective and scalable technique for creating sintered SS parts, has limitations in mechanical strength and porosity control. The process often requires sintering, resulting in lower strength compared with other methods such as SLM or EBM. ¹¹² Controlling porosity and density is also challenging, leading to poor mechanical performance and suboptimal biological outcomes. Studies by Bandyopadhyay et al. ¹¹³ highlight these limitations, preventing widespread adoption of BJ for load-bearing tissue engineering applications. Improving sintering processes and controlling porosity could lead to stronger, more reliable SS scaffolds, expanding BJ’s applicability to more demanding tissue engineering applications such as bone and cartilage regeneration, overall SS-based AM technology. Talha et al. ¹¹³ highlight concerns regarding the long-term biocompatibility of SS alloys, particularly those containing elements such as nickel and chromium. Addressing this gap through the development of safer alloy compositions or surface treatments would be a significant advancement in the field. In both SLM and EBM studies, Zhang et al. ¹¹⁴ stated the need for better control over microstructure and porosity, as these directly impact mechanical properties and cell behavior. Optimizing these parameters could improve the performance of SS scaffolds in tissue engineering applications.

Magnesium alloy composites

Magnesium (Mg) alloys exhibit better biomechanical compatibility with human bone than SS and Ti. The unique properties of Mg need the development of alloys with higher degradation resistance. Another way to improve degradation rates is to coat the metal with a protective layer. Degradation resistance may be improved through the application of a silane coating on certain Mg alloys. However, coatings may not be effective due to their limited life span and the possibility of uneven cracking in the underlying materials. ¹¹⁵

Technical specifications in AM Mg alloy composites

PBF processes such as EBM and SLM: SLM is a process that uses a high-powered laser to melt and fuse metal powder particles layer-by-layer, creating Mg composites. The optimal parameters for Mg composites include laser power (100–200 W), scan speed (800–1500 mm/s), and layer thickness (20–50 µm). Mg’s low vaporization temperature presents challenges, as excessive power can lead to material evaporation, poor surface quality, or high porosity. ¹¹⁴ A study by Zhang et al. ¹¹⁶ showed that tuning the laser power to lower values reduced Mg oxidation during processing, enhancing the surface finish and mechanical properties of Mg-based composites for scaffolds. However, challenges include oxidation and porosity, which can compromise mechanical integrity. SLM-fabricated Mg composites have shown promise in biodegradable scaffolds, particularly in bone regeneration, as demonstrated by Zhou et al. ¹¹⁷ Ahmadi et al. ¹¹⁸ found that lower beam currents and slower scanning rates (1000 mm/s) reduced Mg vaporization in EBM, which led to higher density parts with better mechanical properties, and their biodegradable nature allows them to support bone healing without the need for surgical removal. Powder recycling faces a challenge in the fact that Mg powder degrades after being reused several times owing to oxidation, which increases the overall material cost. ¹¹⁹ Regarding the technical details, BJ is an economical method that works well for printing complicated geometries. BJ is a process that involves the use of Mg powders to create implants. The ideal particle size ranges from 10 to 45 µm, ensuring a balance between layer resolution and powder bed packing density. ¹²⁰ Binder selection and saturation are crucial for maintaining mechanical integrity. The optimal binder saturation level is between 50 and 60%, depending on powder characteristics and layer thickness. Postprocessing, such as sintering, is essential to prevent oxidation and maintain mechanical properties. However, the porous nature of the green parts can compromise the material’s strength. Ceramic reinforcements can enhance mechanical properties while maintaining bioactivity. Degradation behavior is another challenge in Mg-based implants. Polymer coatings such as polylactic acid can modulate the rate while maintaining the scaffold’s bioactivity. ¹²¹ A study by Balan et al. ¹²² on Mg-HA composites found favorable mechanical properties, with a compressive strength of 90 MPa and porosity levels suitable for bone integration. In vitro tests showed a controlled degradation rate over 12 weeks. The success of direct ink writing (DIW) with Mg composites relies heavily on the rheological properties of the ink. Mg particles are often mixed with polymers, ceramic additives, or bioactive materials to achieve the right flow properties. A shear-thinning behavior is ideal for DIW, allowing easy flow through the nozzle but retaining shape after deposition. A viscosity range of 1000–10,000 mPa·s is typically required for effective printing. Binders such as poly(lactic-co-glycolic acid) (PLGA) or alginate are often used to ensure adequate printability. ¹²³ Cross-linking agents such as calcium chloride are sometimes used to enhance the mechanical strength of the printed constructs postextrusion. Precise control over the nozzle diameter is crucial for achieving high-resolution structures. For Mg composites, a nozzle diameter of 200–500 μm is commonly used, balancing structural integrity and resolution. Maintaining uniform dispersion of Mg particles is a technical challenge, and techniques such as ultrasonication are often used to disperse particles homogeneously. Printed Mg scaffolds typically undergo postprocessing steps such as sintering or ultraviolet (UV) curing, depending on the ink composition. UV curing is common for Mg-based inks mixed with polymers, but sintering must be performed under carefully controlled temperatures (∼300–500°C) to avoid oxidation of Mg. However, one major issue is preventing Mg oxidation during printing and postprocessing, which can reduce the scaffold’s mechanical properties and biocompatibility. To address this, DIW processes often need to be conducted in controlled environments or with protective coatings. A study by Fellabaum et al. ¹²³ found that coating Mg particles with biodegradable polymers such as PLGA reduced oxidation and preserved the scaffold’s mechanical properties. A balance between printability and mechanical strength is also a challenge, with soft inks resulting in insufficient mechanical properties for load-bearing applications, and high-viscosity inks improving mechanical strength but increasing nozzle clogging risk. Controlling the biodegradation rate is crucial, as premature degradation could compromise scaffold stability before tissue regeneration.

AM application of Mg alloy composite

The use of several Mg alloys and 3D manufacturing techniques could be tried for enhancement of the implant characteristics. Mg is a good choice for usage in orthopedic implants and bone tissue engineering considering its mechanical qualities on human bone and its low weight. Furthermore, Mg implants are believed to improve osseointegration, a critical component of orthopedic implants. ¹²⁴ Surgical staples were tested in vivo using Mg staples. The incision was sealed without anastomotic leakage using Mg staples. An in vivo study found that Mg-based surgical staples help closure. ¹²⁵ Further research is needed on the study of the prevention of deterioration and hydrogen evolution in Mg materials. Improving Mg-based devices involves decreasing breakage. The paste extrusion method does not require high temperature or reduction in the material strength this technique might allow for the creation of magnetic implants that contain drug use of the paste extrusion process. Mg-ceramic composites can assist in 3D printing, which is essential for the development of Mg implants that mimic the bone structure of realistic bones. ¹²⁵

Gap consideration in AM of Mg alloy composites

SLM: Limited access to high-quality Mg powders for PBF. Research gaps in the SLM process include controlling porosity and pore interconnectivity, which are crucial for tissue ingrowth and nutrient flow. However, SLM has been found to produce irregular pore structures, which can negatively impact cell proliferation and vascularization. Excessive oxidation during fabrication can diminish mechanical properties and biocompatibility. Controlling the biodegradation rate of Mg composites remains a challenge in SLM-fabricated scaffolds, as rapid corrosion can lead to premature failure before sufficient tissue regeneration.^126,127 Studies by Ezhilmaran et al. ¹²⁸ highlighted the potential of SLM to fabricate bioactive Mg-Zn-Ca scaffolds and highlighted the issue of uncontrolled porosity and degradation. Addressing these gaps could lead to better control of scaffold architecture, enhance mechanical stability and biocompatibility of Mg composites, and extend the clinical applicability of Mg-based scaffolds in long-term tissue regeneration. BJ: Limited research on Mg BJ in biological applications ¹¹⁹ There is a lack of suitable binder materials that are biocompatible, degrade at appropriate rates, and do not negatively influence the mechanical properties of Mg-based scaffolds. Mg particles in BJ often suffer from poor bonding after sintering, resulting in weak microstructural integrity and reduced mechanical performance. Recycling of Mg powders during the BJ process is underexplored. This is crucial to improving the sustainability and cost-effectiveness of the technique. BJ has been relatively underexplored for Mg composites, with studies such as those by Salar et al. ¹²⁹ (2022) focusing more on polymers and ceramics. The lack of studies specifically targeting Mg-based materials points to a significant research gap in adapting this AM technique for biodegradable metals. Filling these gaps could make BJ a viable method for producing complex, patient-specific Mg scaffolds for tissue engineering applications, particularly in nonload-bearing environments. In addition, resolving issues related to binder chemistry and recycling could lower the environmental impact and enhance the scalability of Mg-based scaffolds. ¹³⁰ Research gaps in the development of bioinks for Mg composites are significant. There is a lack of bioinks that can achieve optimal rheological properties for DIW, which often does not balance printability, bioactivity, and degradation control. DIW-printed Mg scaffolds often lack mechanical strength due to binder content and particle agglomeration, which can compromise their ability to support load-bearing applications in bone tissue engineering. In addition, current DIW techniques require extensive postprocessing, such as sintering, which can compromise the material’s bioactivity or introduce structural defects. Existing research by Dutta et al. ¹³¹ demonstrated the feasibility of DIW for Mg-based composites, but few studies have addressed the development of bioinks specifically catering to Mg’s unique properties. Filling these gaps could lead to high-performance Mg composite scaffolds tailored for patient-specific tissue engineering needs, particularly in bone regeneration. Overall, in Mg-based AM technology, as mentioned by Han et al., ¹³² scaffolds must be controlled for porosity and degradation to promote tissue ingrowth and vascularization. The limitations of DIW investigations indicate that material compatibility and bioink development are necessary for mechanically stable and bioactive scaffolds. According to Li et al., ¹³³ process sustainability and scalability are needed to convert laboratory succeeds into clinical applications, although current research seldom addresses this problem. Addressing these deficiencies might lead to the creation of next-generation biodegradable alloys for AM and bioinks that increase cellular responsiveness and scaffold biointegration. The possibility to construct patient-specific, biodegradable scaffolds might revolutionize bone and tissue regeneration.

Zinc composites

The biodegradable properties of Zinc (Zn) are considered to be superior to those of Fe and Mg ¹³⁴ due to their excellent corrosion resistance. Zn-based biomaterials are entirely bioresorbable and do not produce excessive hydrogen gas compared with Mg-based biomaterials. ¹³⁵ The primary drawback of pure Zn is its low fatigue strength and a high tendency to creep. Various Zn-based alloys are used in 3D printing. For improvement in mechanical qualities, including Zn-Al, Zn-Mg, and Zn-Ag, research is required on the subject of an increase in Zn biomaterials with higher strain capacity, yield, and tensile strength. ¹³⁶ Zn is a crucial trace element for human health. This enzyme affects growth and tissue regeneration. Zn alloys are options for biodegradable stents and implants. Zn-based biomaterials are used in wound closure, orthopedic, and cardiovascular devices. Zn has a rapid degradation rate, which makes it a promising biological substance. It is involved in bone development and mass maintenance. Zn-based biomaterials may help reduction in bone loss and simulation of bone. ¹³⁷

Technical specifications in AM of the Zn composites

Evaluate the impact of Zn powder’s particle size, morphology, and distribution on the ability to print and the characteristics of the material. Examine the most effective laser or electron beam parameters to get desirable microstructures and mechanical qualities in Zn alloys. ¹³⁸ The optimal laser power (around 200–400 W) and scanning speed (600–1000 mm/s) are crucial for uniform melting of Zn powder, while layer thickness ranges between 20 and 50 µm. Zn’s high thermal conductivity and low boiling points complicate thermal management during SLM. A controlled build chamber environment with temperatures close to its melting point (419°C) can mitigate issues. Maintaining consistent powder quality is challenging due to its oxidation sensitivity, and ensuring the correct inert gas environment is essential for recycling and preventing oxidation. ¹³⁹ EBM typically operates at higher temperatures than SLM, with the powder bed preheated to around 500–700°C. For Zn, this reduces thermal gradients but requires careful control to avoid vaporization due to Zn’s low boiling point ¹³⁸ as well as balancing mechanical strength with degradation rate. PBF techniques can help address these issues by controlling the build environment and refining laser parameters. For example, a study on SLM of Mg alloys found that laser power in the range of 150–300 W and scanning speeds of 400–800 mm/s yielded the most consistent results in terms of porosity and tensile strength for tissue engineering applications. Techniques such as laser remelting can also enhance the mechanical properties of Mg alloys, reducing porosity and increasing fatigue life. ⁵⁵ By incorporating precise control over thermal conditions and material handling, PBF techniques can be tailored to create high-quality Zn composites for tissue engineering.

3D bioprinting involves precise cell and biomaterial deposition to create complex tissue constructs, resulting in mechanical characteristics of Zn-based structures. Multimaterial printing: Investigate the practicality of precisely depositing various materials to fabricate intricate biomedical implants. ¹⁴⁰ The viscosity of the bioink, including Zn composite materials, must be optimized for extrusion. Shear-thinning bioinks with Zn nanoparticles are effective for maintaining scaffold integrity and promoting osteogenesis. Zn ion-incorporating hydrogels have been bioprinted to create tissue engineering scaffolds with ink viscosities ranging between 20 and 40 Pa.s and extrusion pressures of 50–100 kPa. Ensuring the bioink’s mechanical stability and bioactivity is complex due to Zn’s reactive nature. A study by Wang et al. (2022) showed that bioprinted Zn-hydrogel scaffolds significantly improved vascularization and bone regeneration in vitro. ¹⁴¹ In the binder jetting process, various techniques are employed to regulate sintering parameters in order to achieve the desired porosity, microstructure, and mechanical properties. BJ process with Zn composites is crucial for oxidation control, as it can alter the scaffold’s mechanical properties. To prevent this, sintering in a vacuum or using antioxidation coatings can be used. Achieving the right balance between porosity and mechanical strength is essential for Zn scaffolds in tissue engineering. Studies show that adjusting powder particle size and binder concentration can control porosity levels. A study on bone tissue engineering found that a layer thickness of 60 µm and a sintering temperature of 375°C under an argon atmosphere yielded optimal results. ¹⁴² A study by Snelling et al. found that BJ of Zn composites can produce scaffolds with high porosity and appropriate degradation rates. By optimizing binder concentration and sintering conditions, they achieved scaffolds with 60% porosity and a compressive strength of 20 MPa, comparable with cancellous bone. Wang et al. (2021) explored BJ of Zn–calcium phosphate composites, achieving scaffolds that promote osteoconductivity while maintaining mechanical strength for load-bearing applications in bone tissue engineering. ¹²⁸ The selection of the most effective AM technique for fabricating Zn-based composites in tissue engineering depends heavily on specific conditions such as temperature control, laser power, solution concentration, and layer thickness. Each AM method offers unique advantages, but challenges such as Zn oxidation, material homogeneity, and maintaining bioactivity remain prevalent. Incorporating precise technical data from studies on Zn composites in tissue engineering significantly strengthens the understanding of these challenges and offers avenues for future optimization.

AM applications of Zn composite

Research into Zn alloys has included analysis of their structure, production, and deterioration in laboratory and animal models. Through AM, Zn implants tailored to each patient can be created, despite their biodegradability and patient-specificity. Fused deposition modeling (FDM) can be used in the casting of Zn in polymer scaffolds, by making porous Zn scaffolds that can replace trabecular bone. A study of Zn scaffolds has been made for mechanical, topology characteristics, biodegradation, antibacterial capabilities, and cell-bio compatibility. Researchers have used FDM in the construction of scaffolds with low- and high-porosity values for investigations. Because of Zn’s low melting point, high oxidation tendency, and low boiling point, additively manufactured parts tend to be extremely porous. Using FDM, researchers have recently published articles on 3D-printed Zn metal. ¹⁴³ During cellular studies, the rigidity and rate of corrosion decreased with increasing pore size and porosity. Three days of incubation in Minimum Essential Medium (MEM) restored the ability of preosteoblast cells to adhere to and grow on porous scaffolds. All scaffolds used in the research were 100% antimicrobial and biocompatible. The low-porosity scaffold had a lower antibacterial rate than the high-porosity scaffold used in various investigations. These outcomes provide credence to the concept of using Zn in bone tissue engineering. ¹⁴⁴ FDM is replacing conventional processing methods, according to recent studies, for topographically organized Zn scaffolds and solid Zn components with interesting biological applications.

Gap consideration in AM Zn composites

Discover techniques to improve the compatibility of Zn-based alloys created using PBF, taking into account cellular reactions and tissue integration. It is difficult to achieve precise surface finishes on Zn-based implants through PBF, as the smoothness of the surface can have an impact on the biocompatibility. ¹⁴⁵ Research gaps in Zn composite powders include poor flowability, which can lead to inconsistent layering and weak interlayer bonding, and thermal stress and cracking, which have been underexplored in Zn composites. The high energy involved in PBF can introduce thermal stresses, which have been well-studied for Ti and steel implants. Addressing these gaps could lead to Zn-based scaffolds with superior mechanical integrity and structural properties, reducing the likelihood of failure after implantation and broadening the use of PBF for creating large, complex implants for tissue engineering. ¹⁴⁶ DED: Optimizing mechanical properties and corrosion resistances are exploring techniques for controlling the microstructure of deposited Zn alloys to enhance their performance for biomedical applications. Develop techniques to monitor the quality of Zn-based structures during the DED process in real time. ¹¹⁵ BJ: Develop and research binders that are compatible with biomedical applications to prevent contamination and ensure the biocompatibility of Zn-based implants. Research gaps in Zn composites include insufficient research on binders that maintain mechanical strength and biocompatibility after sintering, with current focus on other metals such as Ti. Postprocessing requirements such as sintering can negatively impact the microstructure of Zn composites, and there is limited research on optimizing sintering conditions to preserve Zn’s bioactivity while achieving the desired mechanical properties. ¹⁴² Identifying appropriate binders and optimizing postprocessing could significantly improve the quality of Zn-based scaffolds, leading to larger, more structurally sound scaffolds with complex geometries, potentially supporting tissue regeneration and therapeutic agents. Investigate methods for accurately regulating porosity in Zn structures created using BJ, as variations in porosity can significantly affect both mechanical strength and tissue integration. ¹⁴⁷ FDM: To ensure the safety of implanted medical devices, it is necessary to develop filaments that are more biocompatible for use in Zn-based FDM prints. Research gaps exist in achieving optimal compatibility between Zn and biodegradable polymers or ceramics in extrusion-based methods. The challenge lies in ensuring homogeneous dispersion of Zn particles to prevent corrosion and premature degradation. ¹⁴⁸ Limited data exist on the printability of Zn composites, with most studies focusing on other biomaterials. Resolving these issues could lead to tailored scaffolds with tailored degradation rates, superior mechanical properties, and biodegradation profiles. In order to fabricate complex biomedical designs, it's crucial to improve print resolution and overcome obstacles related to Zn-based filaments. ¹⁴⁹ Overall, in Zn-based AM technology, the researchers are working on improving the oxidation resistance of Zn powders to create more durable and bioactive scaffolds for bone regeneration. They are also optimizing binder materials for BJ to create complex structures for soft-tissue applications. Tailoring the printability and material compatibility of Zn composites in extrusion-based methods could enable the creation of multifunctional scaffolds that combine structural support with controlled drug release. Addressing these research gaps could enhance the performance of Zn-based implants and open new avenues for innovation in tissue engineering, ultimately improving patient outcomes in regenerative medicine.

Titanium composites

Ti is widely utilized, with one million kilogram of Ti implanted yearly. Ti is highly biocompatible owing to its low electrical conductivity, strong corrosion resistance, thermodynamic state at physiological pH (potential of hydrogen), and low tendency for ion production in aquatic environments. ¹⁵⁰ Regarding biomedical uses, Ti alloys top the list of most desired metals in biomedical applications. Ti-6Al-4V (90% Ti, 6% aluminum, and 4% vanadium) is widely used in biological applications because the stress shielding can be lowered due to the material’s excellent biocompatibility and Young’s modulus, and mechanical property analogous to that of human bone. ¹⁵¹

Technical specifications in AM Ti composites

SLM makes use of a high laser power (150–500 W) to selectively melt and fuse Ti powder layer by layer and uses scanning speed (1000–2000 mm/s) to produce Ti and its composites. It is effective in producing parts with excellent mechanical properties due to its precise control of microstructure, resulting in the creation of intricate structures. Outstanding resolution and the capability to create intricate geometries successfully produce parts with high density and excellent mechanical properties. ¹⁵² However, challenges include high residual stresses due to rapid cooling rates and issues with porosity and surface roughness. SLM has been shown to produce Ti-6Al-4V, which exhibits high biocompatibility and osteogenic properties, mimicking those of bone. EBM is a method for fabricating massive Ti components by melting and fusing the powder in a high-vacuum setting ¹⁵³ that uses a beam power of 3–6 kW, a build temperature of 700–1100°C, and a scanning speed of 8000 mm/s to create Ti composites. This process is effective for Ti due to its high build temperature, reducing thermal stresses and warping, critical for complex implants. However, challenges include contamination of material properties and limited resolution compared with SLM. Ti-6Al-4V produced by EBM is widely studied for bone scaffolds due to its excellent strength-to-weight ratios and biocompatibility. LPBF uses a laser to melt and fuse Ti powder in a powder bed in a selective manner and has exceptional accuracy and the capacity to manufacture intricate structures with minute details. Hybrid approaches combine several AM methods with traditional production processes to improve attributes. ¹⁵⁴ Integration of processes such as machining or surface finishing to improve the end product of Ti material.

AM applications of Ti composite

The metal-AM technology when compared with other metals for biomedical applications such as SS and cobalt-based alloys, Ti is significantly more expensive but has some distinct advantages over the long run. ¹⁵⁵ The use of room temperature deposition and heating provides crack-free coatings. The Ti coatings do not flake off after bending. Ti-Nb, Ti-Cu, and Ti-Mo are common alloys. Ti-Nb has shape memory. Alloys containing Cu improve antibacterial characteristics, hardness, and corrosion resistance in Ti-Cu alloys. Superior mechanical qualities, including a high Young’s modulus, are optimal in Ti-Mo alloys. Among the evaluated commercial Ti-containing wires, the Ti-Mo-Zr-Sn and Ti-Nb wires are ideally suited for use in orthodontic and dental procedures. Micrographs were taken with a scanning electron microscope for both the Ti-containing materials before (in their as-received form) and after the corrosion of wires for testing; fluoride-free artificial saliva was used. Figure 10 shows the strengthening phase from γ grains in the aged Ti alloys having adequate mechanical characteristics, including a vast potential for tissue engineering, particularly implants, providing Cu and Ni in the alloys can be substituted by alternative biocompatible metals in research evaluation. Modification of Ti alloys and postprocessing treatment may be tried for improvement in mechanical properties. However, Ti-based alloys have drawbacks. As a result of material mechanical deficiency, Ti-6Al-4V Ti alloy (Ti-64) and CP-Ti (commercially pure Ti) are used for up to 90% of biomedical use. ¹⁵⁶ The most prevalent are dental, orthopedic, bone screws, prosthetic joints, plates, and hearts. The American Society for Testing and Materials (ASTM) 1 to 4 are unalloyed materials, whereas grade 5 contains 6% aluminum and 4% vanadium, which is more durable, as per ASTM F67 and F136. ¹⁵⁷

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FIG. 10.

Scanning Electron Microscopy (SEM) images of the microporous structures of the Ti alloys (Mo Cu Ni 462) ¹⁵⁸ . Copyright 2015, Elsevier.

The printed components’ mechanical qualities and biocompatibility might be impacted by residual stress and porosities introduced by SLM and problems with postprocessing. ¹⁵⁹ SLM has been successful in achieving the desired mechanical properties, but there is limited research on optimizing surface properties to promote cell attachment, proliferation, and differentiation. Most studies focus on bulk material properties, neglecting surface nano/microstructure for cellular interactions. SLM-fabricated Ti parts often lack bioactive coatings, which are crucial for osteointegration. Current studies have not explored how these coatings can be integrated with the SLM process. In addition, there is limited exploration of postprocessing techniques to enhance the surface topography and bioactivity of SLM-fabricated Ti composites. Future research should focus on these techniques. To fix surface roughness and get the surface quality, we have to do certain postprocessing processes. The resolution of EBM could be lower than that of other AM technologies, making it unable to produce very fine details or complex features. ¹⁶⁰ Research gaps in EBM-fabricated Ti composites include the challenge of controlling pore size and interconnectivity at the micron scale, which is crucial for nutrient transport and vascularization in tissue scaffolds. The study focuses on achieving desired porosity, but there is a gap in understanding how these porous structures affect long-term scaffold performance in vivo. In addition, the issue of material heterogeneity in EBM-fabricated Ti composites has not been thoroughly investigated, potentially affecting implant longevity and stability. LPBF: It can be difficult to control the powder bed and ensure uniform powder distribution, which might impair the final part quality. LPBF may have restrictions on the size of components that can be produced in a single build. Integration of processes may be difficult to create a smooth integration of several production processes to achieve the necessary qualities in hybrid processes ¹⁶¹ ; it is essential that the materials utilized be compatible with one another. There is lack of standardized testing protocols for evaluating the biocompatibility of AM-produced Ti components and limited information on the long-term biocompatibility and stability of AM-produced Ti implants.

Natural biopolymer composite

Natural polymers, including starch, alginate, chitin, polysaccharides, chitosan, and gluconate, are called natural polymers and contain many sugars and proteins, such as collagen, fibrin, and silk. Bacteria and fungi produce these natural polymers from plant- or animal-derived origins. Many natural polymers find use in biomaterial applications. For example, starch, which contains many cyclodextrins, is utilized in drug delivery systems. ¹⁶⁷ Chitin and chitosan are used in ligaments, and polysaccharides are also used in drug delivery systems. ¹⁶⁸ Proteins are used in coating biomaterials to make them biocompatible. ¹⁶⁹ The usage of collagen and silk scaffolds is sometimes seen in tissue engineering, but polymers, particularly natural polymers, are commonly used due to their biological recognition feature. Despite biocompatibility, the host systems are comfortable with natural polymer when implanted within the body. They have strong cell adhesion and differentiation, but poor mechanical characteristics and immunogenic qualities since they might be bacterial or animal, causing immunological reactions. ¹⁷⁰ As bovine serum albumin, chitosan, and collagen are animal-derived, animal contamination is a concern. Collagen is also limited; massive amounts of collagen cannot be made, there is also the need to sacrifice animals. With proper fermentation technology, a lot of bacterial polymers may be created. Bacteria in vast concentrations make linear glucan. ¹⁷¹

Synthetic biopolymer composite

When comes to synthetic polymers, (PLA), PCL, polymethyl methacrylate (PMMA), glycolic acid, and PEG, monomers are generated by the polymerization process chemically carried out to produce synthetic polymers. Polymers are synthesized. Some of them are hydrophobic, while others are hydrophilic. As a result, there is a broad variety of hydrophobicity and hydrophilicity; some have thermosetting qualities, others have thermoplastic capabilities, some inhibit bacterial addition, while others do not use biodegradable materials such as polylactic acid and PE. ²³³ Different properties may be obtained through the use of different polymer families. Polymers cannot match metals in certain ways, but they can compensate for them in other ways.

The synthetic polymer can be used in drug-eluting stents, which contain smart polymer coating that elutes the drug. Metal cannot be utilized in orthopedics for drug delivery, but biodegradable polymers can. Nothing beats polymers in drug delivery, considering their suitability for drug delivery, progressive release methods, and bioresorbable scaffolds. So polymers have a bright future and usage. ²³⁴ The problem with hydrophobic polymers is that they are water-repellent polymers, bacteria can attach to crystalline structure polyesters such as PE terephthalate. ²³⁵ Synthetic cells are not attached to polymers unless they recognize them as hosts. As a result, they have no additional signals. Based on the physical or chemical differences, polymers can be amorphous or semicrystalline, but not fully crystalline such as metals. PMMA and polystyrene are all amorphous polymers. ²³⁶ Polymers possess two molecular factors, namely, weight average molecular weight (Mw) and number average molecular weight (Mn), and are distinct concepts. M w = Σ w i m i w i M n = Σ n i m i n i(1)

Mechanical functionality and stability are needed in polymer scaffolds for the prevention of structural damage during an early postoperative function for patient activity and maintenance of the structural integrity of the scaffold holes during cell regeneration. Figure 12(A) illustrates variations that are seen in mechanical strength based on the polymer structure, such as when the network structures have much higher mechanical properties. The second one is the high strength seen in cross-linked structures lower than the network. The next is a branch, which is lower than cross-linked, and the final one is a liner much lower than the above three structures. ²³⁶ Other important factors affecting strength include cross-linking, crystallinity, and Mw. The tensile strength of a material improves as its Mw rises. Polymers will be very strong and durable when they are extremely cross-linked. A crystalline material has greater mechanical characteristics than an amorphous one. Figure 12(B) illustrates that a crystalline structure proved to be significantly stiffer than an amorphous one. ²³⁷

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FIG. 12.

SEM of pristine and 16-week degraded samples (A) PLCL6040 and (B) PLCL7030. ²³⁷ Copyright 2021, Elsevier.

In other words, elongation to break indicates the ductility of any material and the percentage change in its length before it fractures. So, length increases by tugging. Usually, it breaks when it reaches a certain length, especially for athletes and sportspersons, with considerable physical activity experience. ²³⁶ It is a measure of ductility, so thermoplastics are >100%, and thermosets are <5% elongation to break. ²³⁸ Polymers have Young’s modulus, but substantially lower than metals. ²³⁹ Polymers sometimes are brittle when they have a greater Young’s modulus, such as ceramics, but elastomers have a lesser Young’s modulus with a flexible nature. ²⁴⁰ The two primary forms of deformations are elastic and viscous. Stress has a direct relationship with deformation in elastic material. Stress returns to its original shape when removed. It permanently deforms when subjected to a viscous tension. Viscoelastic material gradually rises and then falls but never reaches its ultimate state. So, polymer possesses both elastic and plastic behaviors. Depending on the strain rate and the temperature of the specimen described, such greater strain rate elastic behavior is exhibited at low temperatures, and lesser strain rate viscous behavior is determined when the temperature increases. These stress relaxation and creep recovery features are very important, especially in biological settings. Polyester and hysteresis are referred to as viscoelastic material. ²⁴¹ Many human body systems exhibit this type of viscoelastic property. So, biological tissues are viscoelastic in nature, for example, the aorta. ²⁴²

Thermal characteristics of polymers are also highly important because polymers, particularly the amorphous structure, exhibit both glassy and rubbery states at lower temperatures. Polymer molecules vibrate but with no significant movement, causing the resemblance of polymers to glass and hence the term “glassy state.” This causes them to become brittle, stiff, and rigid. The amorphous portion of the polymer becomes rubbery with an increase in temperature, which means it becomes soft and flexible, much similar to rubber. The crystalline component of the polymer begins to form at certain temperatures due to changes in the melting temperature. ²⁴³ The melting temperature of the polymer might range between 200°C and 300°C. Although biomaterials within the body are not subjected to such intense temperatures, the focus of researchers has been on additives to minimize the high temperatures in biopolymers. So, the significance of these temperature factors influences the polymer’s volume and dynamic mechanical characteristics. ²⁴⁴ Table 3 represents the various biopolymer types and their biomechanical properties.

Technical details: FDM is the layer-by-layer deposition of molten polymer filaments to construct 3D structures. Precise control of printing parameters such as temperature and layer height is crucial for attaining desired material qualities, ²⁶³ a technique that involves extruding thermoplastic polymers through a heated nozzle to create structures, such as PLA, PCL, and polyether ether ketone (PEEK). The optimal printing conditions range from 190–220°C for PLA to 240–260°C for PCL, depending on the polymer’s glass transition temperature. Maintaining the print bed at 40–60°C for PLA and 55–65°C for PCL improves adhesion and reduces warping. A slower printing speed (30–50 mm/s) can result in higher precision and better surface finish. ²¹³ However, FDM faces challenges such as biocompatibility concerns, resolution and surface finish, and achieving uniform HA dispersion due to its extrusion-based nature. Biocompatibility concerns arise from residual stresses and shrinkage during cooling, which can cause delamination or distortion in scaffold structures, impacting mechanical strength and pore uniformity. Microscale resolution for intricate tissue scaffolds is challenging, especially for vascularized tissues. PCL-HA composites have shown enhanced osteoconductivity in bone tissue scaffolds, but uniform HA dispersion remains a challenge due to FDM’s extrusion-based nature. ²⁶⁵ SLS uses a laser to selectively fuse powdered polymer ingredients layer by layer. Polyamide (PA) materials are extensively used in biomedical applications because of their biocompatibility and mechanical robustness and are suitable for producing complicated geometries without the need for support structures. ²⁶⁶ SLS is a process that uses a laser to sinter powdered materials, such as PA (nylon), PEEK, or custom blends of biocompatible polymers, to form a solid structure. SLS requires precise control of laser power (20–50 W) and a scanning speed (500–800 mm/s) to ensure uniform sintering without overheating, which can cause material degradation. ²⁶⁷ Maintaining the powder bed temperature close to the polymer’s melting point is crucial to reduce thermal gradients and warping. However, SLS faces challenges such as material constraints, the use of a limited number of synthetic biopolymers, and controlling porosity, which is essential for tissue scaffolds, especially for vascular networks. Research into nylon-based composites with HA nanoparticles has shown promising results in bone regeneration, but controlling the nanoparticle distribution remains a key challenge. ²⁶¹

SLA uses a laser to selectively polymerize liquid resin layer by layer, resulting in complex and high-resolution structures. Biocompatible resins, such as photopolymerizable acrylates, are extensively used in biomedical applications. It has a super surface polish and resolution, making it suitable for applications demanding high accuracy. ²⁶⁹ Photoinitiators used in synthetic biopolymers are highly sensitive to UV light (∼365 nm). Laser powers around 100–250 mW are typical, depending on the resin’s composition. Reducing layer thickness (20–100 µm) and carefully controlling exposure times (1–5 s) can yield high-precision structures with smoother surface finishes. Many synthetic biopolymer resins used in SLA may release toxic by-products upon degradation, limiting their application in tissue engineering unless postprocessing steps (e.g., washing) are done to remove residual monomers. While SLA allows for high-resolution features, the printed scaffolds often have low mechanical strength, making them unsuitable for load-bearing applications without additional reinforcement; SLA-printed scaffolds using PEGDA composites have shown promising applications for cartilage regeneration. ²⁷⁰ However, the low mechanical strength and long-term stability of these scaffolds remain problematic. Inkjet printing creates droplets of liquid polymer that harden layer by layer allowing for the introduction of numerous materials and even cells into the printing process. Bioprinting, a subcategory of inkjet 3D printing, focuses on generating tissue-like structures. ²⁷¹ Maintaining an ink viscosity is crucial for DIW, as it must allow flow during extrusion but retain shape postdeposition. Controlled extrusion pressure (100–200 kPa) and slow speeds (1–10 mm/s) are necessary for precision in printing intricate scaffold designs, such as those mimicking extracellular matrix (ECM) structures. Cross-linking is required postprinting to stabilize the structures. ²⁷² Achieving uniform cross-linking can be challenging, especially in the case of composite inks containing synthetic polymers. DIW is prone to structural collapse during printing due to gravity, especially for large scaffolds. ²⁷³ This often limits the technique’s application to smaller, softer tissue scaffolds. DIW of PCL/collagen composites has been studied for cartilage repair, where the combination of natural and synthetic polymers aids in both mechanical support and biological integration. However, ensuring homogeneity of the composite material during the extrusion process remains a challenge. ²⁷⁴

AM applications of synthetic biopolymer composites

Initially, scaffolds were made using traditional manufacturing processes. Recently, many porous and complex 3D structures have been made mostly via additive manufacture, which addresses the limits seen in traditional production methods, controlling pore geometry, connectivity, pore size, and more effective customization of produced scaffolds. Scaffolds are 3D constructs that are commonly used in regeneration and tissue engineering. Scaffold pore size and porosity are crucial in biological applications. ²⁷⁵ The open porosity of scaffolds is an important concern in their design and fabrication. It ensures the flow of the blood or culture media, and hence, the continual delivery of nutrients and metabolites, as well as the transfer of oxygen. Furthermore, porous scaffolds promote tissue development. While offering enough mechanical strength for implantation and transplanting in the human body, increased porosity and enhanced mechanical characteristics aid in scaffold proliferation, in vitro cell development, and mineralization. ²⁷⁶ As a result of the growth of 3D printing technology, a substantial study on printing materials, allowing the production of a wide range of high-performance polymers with good mechanical properties similar to metals and the evaluation of structures built with AM technologies, is made based on their features, including mechanical characteristics such as impact resistance, flexural strength, tensile strength, economic factors such manufacturing time and material usage, form correctness, and dimension. To achieve their intended mechanical characteristics, a thorough comprehension of the connection between technical process parameters and mechanical efficiency is required. ²⁷⁷ Mechanical qualities cannot be controlled through alteration in a single parameter; instead, management of numerous factors is required for the achievement of good results. Finding appropriate material is still a problem that researchers must overcome. When compared with other thermoplastic polymers studied, PLA’s mechanical reaction is better, and AM-produced PLA has the highest tensile strength. ²⁷⁸

Gap consideration in AM synthetic biopolymer composite

Gaps in FDM compared with other AM processes: It has lower resolution and surface polish, and high-temperature polymer printing challenges that might benefit specific biological applications. While FDM has been successful with a limited range of synthetic biopolymers, many tissue engineering applications require materials that are more biomimetic or include bioactive molecules. Current FDM techniques are limited in their ability to incorporate sensitive biomolecules or living cells without compromising their functionality due to the high temperatures involved. ²⁷⁹ This gap has been noted in studies focused on bioprinting scaffolds for bone regeneration, where osteoinductive factors could not be included. Although FDM is effective for fabricating structures with specific mechanical properties, it struggles to recreate the heterogeneity of native tissues. ²⁶⁵ For instance, layered tissues such as skin or vascular structures require complex gradients of mechanical properties and bioactive cues, which are not easily achievable with standard FDM processes. Postprocessing could be necessary to increase biocompatibility and surface quality. ²⁸⁰ In the SLA process, the main limitation is the availability of biocompatible and biodegradable photopolymerizable resins. Many of the commercially available resins are either toxic or nondegradable in the body. Advances are needed to develop novel resins that combine mechanical strength, bioactivity, and biodegradability. ²⁸¹ SLA is excellent at the microscale, but tissue regeneration often requires architectures at the nanoscale to better mimic the ECM. Recent studies have emphasized the role of nanoscale topographies in guiding cell behavior; however, SLA’s resolution limitations prevent the creation of such features The necessity for support structures might complicate the design and need additional processes for removal. ²⁸² SLS has limited availability of US Food and Drug Administration (FDA)-approved biological materials. Postprocessing measures could be needed to obtain optimum biocompatibility. The key limitation is the high energy required for sintering, which can denature proteins or degrade bioactive molecules incorporated into the biopolymer powder. This limits the ability of SLS to integrate bioactivity into scaffolds, as noted in studies that aim to develop SLS-based bioactive scaffolds for bone and cartilage repair. The ability to precisely control the printing process for complex structures in bioprinting remains a challenge ²⁸³ Hydrogels used in DIW often lack the mechanical strength required for load-bearing tissues such as bone. Studies have shown that despite good bioactivity, the mechanical properties of DIW-printed scaffolds are inadequate for certain applications, such as tendon or cartilage repair. ²⁷² Many DIW-fabricated constructs require extensive postprocessing to achieve the desired properties, such as cross-linking or further material reinforcement. This complicates the fabrication process and introduces additional challenges, such as ensuring uniformity in the cross-linking throughout the structure. In comparison with other AM processes, in the DLP there are less biocompatible materials available. Scaling up the printing size without affecting resolution is difficult. ²⁸⁴

Overall synthesis of biopolymer AM technology: The techniques discussed offer significant opportunities for tissue engineering advancement, and also reveal clear research gaps. Identified research gaps and their impact on the field: The availability of biocompatible, biodegradable, and bioactive materials in additive manufacturing (AM) techniques remains a major challenge, impacting the ability to create scaffolds that fully mimic the native tissue environment. The resolution of many AM techniques, particularly at the nanoscale, is insufficient to fully replicate the complexity of tissues such as bone, cartilage, or skin. This gap hinders the full functionalization of scaffolds. ²⁶⁷ Integration of bioactivity is also a challenge in many AM techniques, such as SLS and FDM, due to the high energy or temperatures involved in the process. Integrating growth factors, peptides, or proteins without compromising their activity is vital for advancing tissue regeneration strategies. Solving this problem could enable more effective scaffolds that actively promote tissue regeneration. ²⁶⁷ Mechanical strength for load-bearing applications is also a gap, particularly for tissues such as bone and cartilage. While bioinks used in DIW or SLA offer great flexibility and cell compatibility, their mechanical performance needs to be enhanced to expand their use to a wider range of tissue types. ²⁸⁵ Addressing these gaps could lead to major advancements in tissue engineering, such as material innovation, functional scaffolds with bioactivity, and customization for complex tissues. By addressing these gaps, the field could achieve a more complete synthesis of materials science, biology, and AM technologies, leading to new insights into tissue repair and regeneration, ultimately improving clinical outcomes.

Dental implant composites

Dental implants are quite ancient, the use seen from many centuries when people had their teeth replaced with artificial materials. Ti is a biocompatible metal with high strength and harness. As a result, it is widely used in orthopedic applications and also in dental equipment such as screws. ⁴¹³ Alumina is a fairly biocompatible material despite being a ceramic or an inorganic material. Therefore, it finds extensive use in various dental implements. HA is also used in dental equipment to fill gaps. ²⁵⁹ Then, there are tooth fillings in several forms such as mercury, PMMA, and acrylic acid. When it comes to dental implants, the thermal expansion coefficient is the most important factor to consider. ⁴¹⁴ As a result of temperature changes, the shape of an item shifts as well.

The following equation may be used for the calculation of the thermal expansion coefficient: Linear coefficient of expansion : α = ( Δ L L ) Δ T(2) Volumetric thermal expansion = 3 α(3)

In dental implants such as PMMA are very important when dealing with materials that need sterilization. The use of methylacrylate is recommended considering all of the polymerization. Polymerization also raises the temperature significantly, which can cause variations in thermal expansion and volumetric expansion. ⁴¹⁵ Changes in the temperature inside the implants and the dentin teeth placed may also occur. Changes in their linear and volumetric growth rates lead to many consequences. So, materials should be designed to have volumetric or linear expansion coefficients comparable with dentin and other parts of the mouth cavity. Another part is the directed tissue regeneration membrane and the void filling in the dental region, filling up the tooth space. Development of material is seen following extraction. Therefore, it should degrade gradually. Especially in acid environments, ceramics may get damaged due to changes in pH. Dental implants such as calcium phosphate and HA are susceptible to changes in pH. ⁴¹⁶ Decision on the use of ceramics depends on the specific situation.

Another important metal is also used in dental alloys for filling the teeth. Dental amalgams are used for filling in teeth. Mercury-based dentistry has been around for hundreds of years, for example, silver and mercury, copper and mercury, and strontium and mercury. Mercury is malleable and so can be molded to fit any form or gap, filling it and solidifying rapidly. ⁴¹⁷ As a result, amalgam is still widely utilized in the form of mercury, it poses a concern due to the risk of leaks, which can be harmful to users. However, one of its key benefits is its ability to solidify into almost any shape, making it ideal for filling voids, holes, and other irregularities. Since then, dental amalgam has become highly popular. Crowns and bridges are made of noble metals and base metal alloys. ⁴¹⁸

The PMMA-based front end can be used in partial dentures since it matches the color of natural teeth. ⁴¹⁹ Face reconstruction or jaw reconstruction has to be done when an accident occurs, due to a major injury. Components such as wires and brackets are used as orthodontic materials. These materials include nickel–Ti alloys, Ti, and SS, in which Ti and Ti bridges are used as dental implants. A high level of inherent strength and fracture resistance is essential, since teeth undergo much mechanical activity with jaws moving teeth up and down, touching one another, food, and then bodily fluids, enzymes, and germs. So, teeth constitute a hard environment, where there will be much wear and tear since they move constantly. In addition, any dental material needs to be biocompatible as it comes in contact with body tissues and the mouth’s inner portion as well, such as the tongue, saliva, and other fluids, and in contact with numerous supporting tissues. Dentistry is a very complex field. So, the development of materials that are durable, strong, and wear-resistant is a challenging task.

As for inlays, onlays, and endodontic posts, they must be strong and tough. Wear resistance is due to a lot of mechanical activities, between teeth and food, and particle corrosion resistance as they are subjected to body fluids, as well as touch with both hard and soft tissues. Therefore, compact thermal expansion coefficients of metals and porcelain cladding are issues that have been engaging the attention of researchers for a long time. So, metals are covered with porcelain, due to strong bonding and good corrosion resistance.

Dental wires can be made from a broad range of composite materials, including Ti, SS, nitinol, cobalt, and chromium-nickel. They have high yield strength, but the elastic modulus, which is important to match bone modulus that is critical, is low. So novel material combinations such as nitinol are used. ⁴²⁰ Dentists use alloys such as gold silver alloys, silver copper alloys, various forms of silver and copper, and silver tin alloys, for construction in bridges, inner and outer cladding, and other dental components. ⁴²¹ Other alloys include cobalt, chromium, nickel, and nitinol, for example, “hyperelastic wires” made of nickel–Ti alloys. ⁴²² Apart from orthodontic wires, which hold the jaw in place, and cranial wires, which hold different fractured portions of the skull together, most of their application is in wires. As a result, metals abound in this specific application. Scanning or imaging technologies such as CT and MRI, and intraoral scanning combined with PMMA applications, are acquiring greater importance in dentistry. ⁴²³

Orthobone composites

Bone deficiency either congenital or caused by an accident is referred to as a bone defect. Biomaterials, such as cellulose and collagen, are used to fill the defect. ⁴²⁴ In joint prostheses, bone plates are used in the replacement of joints. They are made of various materials or even a combination of two or more materials. Researchers have attempted at cultivation of cells on surfaces with scaffolds for tissue engineering, a discipline that has seen significant growth over the past 5–10 years. Consequently, once they are perfected, they can be utilized to replace human tissues that have been injured or compromised by disease. ⁴²⁵ The human body has 270 bones at birth and 206 bones as one gets older due to some bones getting fused. Human bones, the humerus at the radius, outside, the ulna, which is within the arms, hands, and fingers, the femur, lower fibula, tibia, and feet, as a result, many bones are subjected to tremendous tension and compression. ⁴²⁶ During orthopedic surgery, such as bone replacement or defect filling, the biomaterial is a metal, a combination of metals and ceramics, or a polymer. Different mechanical properties of the biomaterial require consideration when it comes to orthopedic materials.^427,428 Metals such as Ti, SS 316L, and alloys are commonly utilized in the orthopedic field. Pure Ti cages are used in the treatment of long bone abnormalities. Special action is required to deal with the increase in deformations. In this case, Ti cages cure long bone segmental abnormalities. ⁴²⁹ The mechanical characteristics of these biomaterials must meet the requirement. They must either be in a tension or compression state. The stress-to-strain ratio is the elastic modulus. The most significant problem is in finding a material with the same elastic modulus. Bone has a low elastic modulus, and hence, any material with a high elastic modulus such as SS can be used. So then comes Ti. Now, researchers are looking into additional alloys for the decrease in the elastic modulus. ⁴³⁰ Materials having a higher elastic modulus are generally used for bone replacement. However, researchers realize such materials near the bone could cause problems due to their different elastic modulus values, beginning with deterioration in bone strength primarily supported by materials with a great elastic modulus, called stress shielding. ⁴³¹

Much research is currently being conducted on materials with extremely low modulus of elasticity with the inclusion of ceramics and some of the oxides. Ceramic segments, including aluminum oxides, ⁴³² zirconia, ⁴³³ calcium phosphate, ⁴³⁴ calcium sulfate, ²⁷⁹ carbon, ⁴³⁵ and glass ⁴³⁶ ceramics, are grouped, they are highly biocompatible and fit flawlessly into the human body without causing any negative effects. HA, calcium phosphate, and calcium sulfate are components of bone and other soft tissues; besides, they are biocompatible, with the same compressive strength and elastic modulus as the bone. Hardness is the next quality to be considered in biomaterials. A certain level of hardness is required considering the nature of the bone. To be safe, they should be hardened but not harder than the bone itself in surgery. The substance adjacent to the bone may penetrate the bone. Thus, matching the hardness is quite crucial. ⁴³⁷ Fracture strength is the maximum allowable stress before a material fracture. That is why it is important to have high fracture strength. Another important one is fracture toughness, which refers to the progression of the material break. This is particularly true for ceramics, which keep the biomaterial’s long-term clinical viability and serviceability. Thus, the durability of the biomaterials is important. Fracture toughness refers to the rate at which a material’s tensile strength is sustained. In other words, fatigue results from a lack of these properties. The diaphragm heart valve closes and opens repeatedly. Repeated loading or cyclic loading is another name for this mechanism. Therefore, it might have tensile or compressive strength properties. ⁴³⁸ Another important point is Poisson’s ratio. A metal implant in the hole: tension enlarges the implant length but produces shrinkage, which may lead to the formation of a gap in the diameter. So, a high Poisson ratio spells trouble. Strain elongates the length but reduces the diameter. Therefore, these biomaterials are separated from the bone, which might be harmful (or) dangerous. As observed in oral surgery, commonly used screws include Ti and SS for prosthetics in orthopedic cases. ⁴³⁹ When these screws are subjected to tensile stresses, they may elongate a bit. Poisson’s ratio for metals is approximately 0.4; hence, a moderate diameter shrinkage could also occur. ⁴³⁹ The tensile strength of bone is measured between 50 and 150 MPa. ⁴⁴⁰ Ti is 345 MPa ⁴⁴¹ and SS ranges from 465 to 900 MPa. ⁴⁴² Even though Ti has a lower tensile strength than SS, durability has kept it in widespread usage.

Materials such as alumina, bioglass, and HA (inorganic oxides) have very low tensile strength, used for bone application. ³⁹³ Bone has a compressive strength of between 100 and 230 MPa. The organic oxides have strong compressive strength and are found being used in bone replacement. It is vital to note that SS has a similar range of compressive strength to alumina. In comparison, the elastic modulus of bone is between 7 and 30 GPa. SS (200 GPa) and Ti (102 GPa) have an extremely high elastic modulus, and the elastic modulus of alumina (400 GPa) and hydropathies (120 GPa) is also relatively high. The crack and how it propagates is the term that refers to fracture toughness. It is too low for the bones (2 to12 MPa·m^1/2), compared with the fracture toughness of other materials, the alumina (6 MPa·m^1/2) and HA (1 MPa·m^1/2). Their fracture toughness is acceptable as an alternative to bone.