Additive Manufacturing of Bioactive Glass and Its Polymer Composites as Bone Tissue Engineering Scaffolds: A Review

Bioactive glass (BG) and its polymer composites have demonstrated great potential as scaffolds for bone defect healing. Nonetheless, processing these materials into complex geometry to achieve either anatomy-fitting designs or the desired degradation behavior remains challenging. Additive manufacturing (AM) enables the fabrication of BG and BG/polymer objects with well-defined shapes and intricate porous structures. This work reviewed the recent advancements made in the AM of BG and BG/polymer composite scaffolds intended for bone tissue engineering. A literature search was performed using the Scopus database to include publications relevant to this topic. The properties of BG based on different inorganic glass formers, as well as BG/polymer composites, are first introduced. Melt extrusion, direct ink writing, powder bed fusion, and vat photopolymerization are AM technologies that are compatible with BG or BG/polymer processing and were reviewed in terms of their recent advances. The value of AM in the fabrication of BG or BG/polymer composites lies in its ability to produce scaffolds with patient-specific designs and the on-demand spatial distribution of biomaterials, both contributing to effective bone defect healing, as demonstrated by in vivo studies. Based on the relationships among structure, physiochemical properties, and biological function, AM-fabricated BG or BG/polymer composite scaffolds are valuable for achieving safer and more efficient bone defect healing in the future.


Introduction
The fracture, infection, and surgical removal of bones may all lead to the development of bone defects. With the number of new incidences of bone fracture having increased by 33.4% in the past three decades, the clinical and economic burden associated with bone defects has reasonably grown with time [1]. The incidence of bone defects results not only in chronic pain but, more importantly, in the loss of bone function (for example, the production of blood cells, the maintenance of mineral homeostasis, weight-bearing capacity, the assistance of limb movement, and the protection of organs/soft tissues by the cranium, ribs, and vertebrae) [2]. Nevertheless, the incidence of bone defects significantly impairs patients' quality of life [3].
Once damaged, the bones can initiate self-mediated regeneration. Nonetheless, when the size of the bone defect is beyond a critical size, for example, 1-2 cm or 50% of the circumvent of the bones, spontaneous bone regeneration cannot occur [3]. Autologous bone transplantation remains the gold-standard therapy for critically sized bone defects, which involves transplanting the patient's own bone to the defect site. Nonetheless, this technique involves numerous drawbacks, including limited donor bone availability, requiring additional surgeries to harvest the bone, and donor site morbidity [4]. Bone grafts 1.
Alkali metal cations within glass exchange with H + /H 3 O + in the surrounding medium (Si-O-M + H + → Si-OH + M + ).

2.
Hydrolytic attacks take place at the Si-O-Si bonds within the soluble SiO 2 , giving rise to hydrated silica (Si-OH) at the BG-liquid interface [38]. 3.
The hydrated silica undergoes polycondensation and repolymerization, which results in the formation of a silica-rich gel layer, as well as the depletion of metal cations from the BG.
The calcium phosphate layer crystallizes by incorporating OH − and CO 3 2− from the surrounding medium to form a hydroxycarbonate apatite (HCA) layer.
The HCA layer subsequently enables the adsorption of biological moieties, which modulate the attachment, proliferation, and osteogenic differentiation of stem cells. Eventually, the cells committed to osteogenesis produce a bone matrix and lead to new bone formation [47].
The five steps of HCA crystallization are regarded as the basis of the bioactivity of BG, and are dependent on the BG chemical formulation to a great extent. Typically, BG contains high CaO content and <60 mol.% of SiO 2 , as shown in Figure 1D [10]. Based on this diagram, more BG formulas (Table 1) are developed with variations in the type and ratio of metal oxides in the glass composition, with some formulas approved and commercialized for clinical application.
An evident drawback of silicate-based BG is the incomplete conversion of glass into bioactive apatite. As mentioned previously, a silica gel layer is formed at the surface of the BG in body fluid. Such a gel layer only allows metal cations from the glass to be leached into the surrounding medium, but protects the inner region of the BG from dissolution [48]. Consequently, the glass beneath the silica gel layer converts into a silicarich phase that cannot be replaced by newly formed bone tissues. It has been reported that the residual silica with a submicron size induces cytotoxicity and chronic inflammation, and the accumulation of silica in the spleen and liver may induce further complications [49,50].

Code Composition (Oxides of Each Element) Remark
Refs.
Unlike silicate-based BG, phosphate-based BG is fully degradable in an aqueous environment. After an ionic exchange of metal cations and H + /H 3 O + , the P-O bonds are gradually hydrolyzed, with no protective gel layer formed, resulting in the breakdown of the entire glass network [57]. Calcium and other doped elements, along with phosphorus, are then gradually released into the surrounding media, contributing to the formation of a calcium phosphate layer over the surface of the BG.
For phosphate-based BGs with P 2 O 5 content > 45 mol.%, the formation of abundant phosphoric acid after hydrolytic degradation often leads to increased acidity in aqueous media [58]. This acidity, in turn, accelerates the degradation of phosphate-based BG [59]. As such, the phosphate-based BG undergoes autocatalytic degradation. Moreover, these glasses are less likely to induce precipitation of the bioactive HCA layer after degradation, as increased acidity of an aqueous environment favors the precipitation of dicalcium dihydrate, while HCA is prone to dissolve in an acidic environment [60]. In contrast, several invert phosphate glasses (P 2 O 5 content < 40%) have been reported as bioactive, considering that the degradation media remained neutral or slightly basic to favor the nucleation of hydroxyapatite [61][62][63]. It has also been reported that titanium in an invert phosphate glass plays a vital role in the induction of bioactivity. Tetravalent titanium is found to ionically cross-link the phosphate units, thereby interrupting the glass network and causing a reduction in hydrolyzable P-O-P bonds, leading to reduced solubility of the phosphate glass. Acidification of the surrounding media is therefore impeded, giving rise to the precipitation of bioactive hydroxycarbonate apatite. Moreover, Kasuga et al. reported the formation of a Ti-enriched interfacial layer between 30P-60Ca-10Ti invert phosphate glass and bioactive apatite. The authors stated that the intermediate layer was likely to be a Ti-OH gel layer, and functioned similarly to the Si-OH gel layer on silica-based bioactive glass to facilitate apatite nucleation [64].

Borate-Based BG
It was at the beginning of the 2000s that borate-based BG caught the attention of researchers in the field of bone tissue engineering [65]. The glass network of a borate-based glass contains trigonal, planar BO 3, and tetrahedral BO 4 at the same time. As shown in Figure 1C, vitreous B 2 O 3 only contains BØ 3 units, where all three oxygens are bridging oxygens in between two boron atoms, i.e., B-Ø-B. With the addition of metal oxides, trigonal planar BØ 3 converts into a tetrahedral BØ 4 − unit composed of four bridging oxygens and one boron bearing negative charge [66]. With all oxygen atoms being bridging oxygens, the network connectivity of the glass network increases. The further addition of metal oxides, however, results in the reversible conversion of BØ 4 − into BØ 2 O − , where non-bridging oxygen emerges. The transition of borate units continues with an increasing number of metal oxides added into the network, leading to the formation of BØO 2− , BØ 2 O 2− , and eventually, BO 3− units. With a lower ratio of bridging oxygens, the glass network becomes depolymerized and the network connectivity is gradually reduced [38]. Along with the changes in the glass network connectivity, a specific physiochemical property of glass (e.g., solubility or strength) usually exhibits a non-monotonous change and maximizes/minimizes as the BØ 4 − content maximizes, corresponding to the maximum network connectivity. This feature is termed "borate anomaly" in the literature [67,68].
As boric acid has rather limited acidity, the degradation of borate-based glass containing oxides of alkali/alkaline earth metal typically creates a neutral or basic pH in the degradation media, which facilitates the formation of hydroxycarbonate apatite over the glass surface [69]. Unlike silicate-based BG, no protective gel layer is formed to protect the inner parts of borate-based glass from further dissolution. Therefore, a borate-based, silicate-free glass converts completely into a bioactive compound enriched with calcium phosphate after immersion in the body fluid [48].
As borate-based BG is fully dissolvable, care must be taken to control the degradation rate of the glass by tuning the chemical formulation. It has been reported that the viability of both murine bone marrow-derived stromal cells and murine pre-osteoblasts-like cells decreased with increasing boron concentration in culture media [70,71]. These findings indicate possible hazards due to the excessively rapid dissolution of borate-based BG. While the dissolution rate was determined by the chemical formulation of glass [65], it was also reported that the macroscale architecture of the glass scaffold, which could be accurately defined via AM, may be exploited to modulate the dissolution rate of borate-based BG [72].  Reprinted from references [73]. (E-J) Schematic diagram of glass network and the mechanism of hydrolytic degradation as well as bone-binding bioactivity of (E,F) silicate-based BG, (G,H) phosphate-based BG, and (I,J) borate-based BG. Reprinted and adapted from references [74,75].

BG-Based Polymer Composites
While bioactivity allows the BG to directly bond to bone tissues, the processing of BG into the desired geometry is difficult owing to the low fracture toughness and a high modulus of BG [76]. For this purpose, one strategy is to develop BG/polymer composites, where BG serves as a functional agent that induces bioactivity by releasing elements, and also mechanical reinforcement to improve the stiffness of the resultant material. The rest of the composite is primarily the polymer matrix, which binds the BG together and provides toughness to the construct. The polymer matrix also endows the composite with ductility against brittle fracture, which is an evident drawback of pure BG. In regard to Relationship between the composition of Si-P-Ca-Na glass-ceramic (P = 6 wt.%) and index of bioactivity (I b ), defined as the inverse of the time required for bonding formation at >50% of the bone-BG interface. Reprinted from references [73]. (E-J) Schematic diagram of glass network and the mechanism of hydrolytic degradation as well as bone-binding bioactivity of (E,F) silicate-based BG, (G,H) phosphatebased BG, and (I,J) borate-based BG. Reprinted and adapted from references [74,75].

BG-Based Polymer Composites
While bioactivity allows the BG to directly bond to bone tissues, the processing of BG into the desired geometry is difficult owing to the low fracture toughness and a high modulus of BG [76]. For this purpose, one strategy is to develop BG/polymer composites, where BG serves as a functional agent that induces bioactivity by releasing elements, and also mechanical reinforcement to improve the stiffness of the resultant material. The rest of the composite is primarily the polymer matrix, which binds the BG together and provides toughness to the construct. The polymer matrix also endows the composite with ductility against brittle fracture, which is an evident drawback of pure BG. In regard to bone tissue engineering, free space must be provided for subsequent bone growth, and thus, requires the polymers to be degraded upon implantation without inducing adverse effects on the biological systems [77].
Depending on the origin, biodegradable polymers are typically classified as natural polymers and synthetic polymers. Natural polymers, which include proteins (e.g., collagen, gelatin, and silk fibroin), polysaccharides (e.g., cellulose, alginate, hyaluronic acid, and chitosan), and polynucleotides, are naturally occurring materials harvested from animals, plants, or microorganisms [78]. Natural polymers show great similarity to the components of the extracellular matrix in human tissues [79]. For instance, type I collagen constitutes 90% of the weight of the organic phase of bones, whereas type II collagen is abundantly found in cartilage [80,81]. Chitosan, though not native to the human body, is structurally similar to glycosaminoglycans, which constitute the extracellular matrix [82]. This compositional similarity makes natural polymers more recognizable by the biological system, thereby lowering the risk of chronic inflammation and cytotoxicity after degradation [83,84].
Nonetheless, the qualities and properties of natural polymers may exhibit batch variations between different sources [79]. Moreover, pathogens and immunogenic/antigenic substances can be transferred from donors to recipients, causing infection or significant foreign body reactions [85,86]. Last, but not least, natural polymers, especially for proteins, are prone to denaturation due to heat, radiation, and chemicals applied during material processing, which causes impaired mechanical and biological performance. Therefore, care must be taken to control the parameters and chemicals used to avoid denaturation of the natural polymer matrix and altered properties of the composites thereof.
Synthetic polymers, in contrast, are artificially produced materials. The most investigated synthetic polymers are aliphatic polyesters, with representative materials including polyglycolic acid (PGA), polylactic acid (PLA), poly-ε-caprolactone (PCL), and their copolymers. These materials degrade mainly via the hydrolytic route. Following water absorption, the ester groups in the polymer backbone are cleaved via hydrolysis to yield oligomers/monomers. These smaller molecules are further metabolized into CO 2 and water and are eventually eliminated from the body [77]. Compared to naturally derived polymers, synthetic polymers are superior in the consistency of their properties among different batches. To satisfy the clinical requirement for the degradation rate and initial mechanical properties, the ratio among different monomers and the molecular weight of the resultant polymers can be finely configured on-demand [87,88]. Owing to good biocompatibility, excellent processibility, and mechanical properties, synthetic polymers have received approval from the FDA for clinical application [89][90][91].
However, as synthetic polymers do not exist in the biological system, implanted synthetic polymers and their wear debris may invoke a significant foreign body reaction. Most synthetic polymers are hydrophobic, making it difficult for water and protein to be absorbed on the polymer surface, consequently hindering the adhesion of cells [92]. Another safety issue is that CO 2 is generated after the metabolization of monomers. This is believed to cause increased acidity at the implantation sites, which impairs the survival of both osteoblasts and mesenchymal stem cells, meanwhile stimulating osteoclastic (boneresorbing) activities [93][94][95][96]. Interestingly, the addition of BG was reported to enhance the surface hydrophilicity, neutralize the post-degradation acidity, and retard the general degradation rate of composites with a synthetic polymer matrix, thereby ameliorating some of the concerns stated earlier [78].
A summary of the structural and mechanical properties of bone, cartilage, BG, and typical BG composites is listed in Table 2. While BG is too brittle to be machined into desired geometries, the presence of a polymer matrix endows the resultant composites with processibility. A wide range of techniques are available to yield composites of desired shapes and, more importantly, an intricate porous structure that determines the mechanical and degradational properties of scaffolds, as well as tissue ingrowth [97]. Gas foaming, freeze-drying, and thermally induced phase separation (TIPS) are frequently applied to induce porosity in the polymer phases. However, control of the shape, interconnectivity, size, and uniformity of pores has been proven difficult using these techniques [98]. For instance, during uncontrolled freeze-drying, the difference in heat transfer rate throughout the solution in a container leads to uneven pore size in the resultant scaffold [99]. Moreover, the use of organic solvents acts as an additional safety concern in TIPS [100]. The use of porogens of well-defined size, shape, and volume allows for better control over the porous structure, yet the on-demand spatial distribution of pores and complete leaching of porogens cannot be guaranteed [100,101]. There is a pressing need for a novel fabrication technique of BG and BG/polymer composites that can precisely configure the geometry and porous structure so as to achieve accurate fitting of the defect site and a well-defined porous structure that satisfies the need for effective tissue ingrowth.
The thermal treatment process and BG solid content were optimized, leading to lower linear shrinkage and improved compressive mechanical properties of glass scaffolds [122] 13-93 Glass 100 • Infiltration of biodegradable polymers into BG of interconnecting porosity led to enhanced toughness, stiffness, and strength (compared to pure BG) [76] Glass-Ceramic 100 648.14 ± 81.12 (C)~7 (C) • Avoided burst pH reduction due to autocatalytic PG degradation • Sustained release of calcium and phosphate [127] The incorporation of BG with optimized content at 33 wt.% upregulated the proliferation and osteogenic differentiation of rBMSC [128] 58S/PLDLA Glass/Polymer Composites 10  [130] Note: C/F/T in maximum stress or elastic modulus denote modes of mechanical testing. C = compression, F = flexure, T = tension.

Additive Manufacturing of BG and BG/Polymer Composites and Their Application to Bone Tissue Engineering
Additive manufacturing, according to ISO/ASTM 52900:21, refers to a process of "joining materials to make parts from 3D model data, usually layer upon layer, as opposed to subtractive manufacturing and formative manufacturing methodologies" [131]. Depending on how the materials are added to form an integral object, numerous AM technologies have been developed, with melt extrusion, direct ink writing (DIW), vat photopolymerization (including stereolithography and digital light processing), and powder bed fusion being frequently applied to produce BG or BG/polymer composites ( Figure 2); these techniques are elaborated upon in the following section.

Additive Manufacturing of BG and BG/Polymer Composites and Their Application to Bone Tissue Engineering
Additive manufacturing, according to ISO/ASTM 52900:21, refers to a process of "joining materials to make parts from 3D model data, usually layer upon layer, as opposed to subtractive manufacturing and formative manufacturing methodologies" [131]. Depending on how the materials are added to form an integral object, numerous AM technologies have been developed, with melt extrusion, direct ink writing (DIW), vat photopolymerization (including stereolithography and digital light processing), and powder bed fusion being frequently applied to produce BG or BG/polymer composites ( Figure 2); these techniques are elaborated upon in the following section.
Following the introduction of AM technologies, this section summarized how additive-manufactured BG or BG/polymer scaffolds have been applied in bone defect healing owing to their ability to fabricate objects with highly complex geometry and intricate porous structures. Published works based on animal studies were reviewed in detail, highlighting the ability of AM to configure the structural properties of resultant objects and to satisfy the need for clinical application, for example, a patient-specific design and the ondemand distribution of multiple functional biomaterials.

Melt Extrusion
Melt extrusion refers to a process where molten feedstock is extruded from a nozzle to fabricate an object via on-demand material addition. The feedstock is typically a Following the introduction of AM technologies, this section summarized how additivemanufactured BG or BG/polymer scaffolds have been applied in bone defect healing owing to their ability to fabricate objects with highly complex geometry and intricate porous structures. Published works based on animal studies were reviewed in detail, highlighting the ability of AM to configure the structural properties of resultant objects and to satisfy the need for clinical application, for example, a patient-specific design and the on-demand distribution of multiple functional biomaterials.

Melt Extrusion
Melt extrusion refers to a process where molten feedstock is extruded from a nozzle to fabricate an object via on-demand material addition. The feedstock is typically a composite composed of a thermoplastic polymer. It is continuously delivered into a hot end, where heating is applied to melt and liquefy the feedstock. Under the control of a program, the molten feedstock is forced out of the heated nozzle, which travels across the deposition platform to deposit molten feedstock at the specified location. As the temperature drops, the molten extruded materials consolidate, yielding a "sliced" solid layer of the model [132]. Following this, melt extrusion is performed again over the previously deposited layer. With repetitive material deposition in a layer-by-layer manner, a solid object is eventually fabricated.
Most melt extrusion AM devices use feedstock in the form of a continuous filament (Figure 2A). This technology, typically termed "fused deposition modeling (FDM)", requires that the thermoplastic polymer and BG be homogenized and extruded into filaments beforehand. The incorporation of rigid BG, however, increases the brittleness of filaments, which are prone to breakage when being fed into the hot end [120]. To avoid this, other devices use pellets/powders of a BG/polymer mixture or BG/polymer composites as feedstock, which are melted, and then, propelled either via pneumatic pressure [37] or using a screw [133,134].
Owing to its advantages, including ease of use and the low cost of the device, melt extrusion is one of the most widely applied AM technologies [135], with several bone tissue engineering implants and drug delivery systems fabricated via melt extrusion already cleared for clinical application [136,137]. With the incorporation of BG, the BG/PCL composite scaffolds fabricated via melt extrusion have been reported to have enhanced surface hydrophilicity that improves cell adhesion, as well as the function to enhance in vivo bone regeneration via the elements relseased [138,139].
However, the spatial resolution of melt extrusion-based deposition methods is comparatively low. While nozzles with a small outlet (diameter down to 100 µm) are able to produce fine structures, the force required to extrude molten polymer from a fine nozzle increases dramatically and potentially leads to the buckling/breakage of feedstock filaments. An elevated temperature may reduce the viscosity of polymer melts to ease the extrusion; however, this would be at the expense of more pronounced thermal degradation of the polymer and compromised mechanical properties of the fabricated objects [140].

Direct Ink Writing (DIW)/Robocasting
DIW, also known as robocasting, is another AM technology based on material extrusion. As shown in Figure 2B, the feedstock used in the DIW process is typically a viscous ink, which is extruded through a nozzle and deposited on the deposition platform/previously extruded ink for layer-by-layer material addition. Post-fabrication processing to convert the as-fabricated, semi-solid model into a consolidated object is usually required, with typical post-processing methods including solvent evaporation [121,141], solvent extraction [142], thermal crosslinking [143], photopolymerization [144], and chemical crosslinking [145][146][147].
The key to successful DIW is to obtain inks with ideal processibility. On one hand, the ink shall flow smoothly and uniformly from the nozzle, without pulsating extrusion or clogging of the nozzle. This requires that the BG particles be evenly distributed within the ink without forming large clusters. On the other hand, the extruded material needs to span across the gaps and retain its as-deposited shape without slumping before being converted into a solid. For this purpose, the rheological and viscoelastic properties of ink play a decisive role, as have been reviewed in detail elsewhere [148]. An essential requirement is that inks be shear-thinning such that the viscosity of flowing inks in the nozzle capillary, under a high shear rate, is low to facilitate extrusion. After deposition and the reduction in the shear rate, the viscosity of the ink should be adequately high to maintain its shape. To avoid slumping, M'Barki et al. concluded that the dynamic yield stress of deposited ink should overcome the synergistic effect of gravity and surface tension [149], while Chan et al. considered the product of storage modulus and yield stress as a simple, universal, and effective criteria to predict whether the deposited ink slumps [150]. However, the rheological properties of inks containing BG may not be universally predicted. In most cases, the viscosity of ink typically increases with the addition of BG. For alginate-based inks, the release of Ca 2+ from BG further induces the ionic crosslinking of alginate, and consequently, increases the viscosity of the ink [151][152][153][154]. Nonetheless, the viscosity of the composite ink may also decline with an increased ratio of BG in the ink, possibly due to the disentangled polymer network of inks and insufficient bonding strength between the BG and the polymer [155,156]. Therefore, the rheological behavior of BG-containing inks needs to be studied to determine the proper BG content within the ink, as well as the parameters of the DIW process.
When compared to melt extrusion, a major advantage of DIW is that a high temperature is not required for liquefying the feedstock. As such, the organic compounds are prevented from undergoing thermal decomposition or denaturation, making DIW a suitable technology to fabricate BG/polymer composite scaffolds composed of natural polymers, drugs, and protein. Zoledronic acid (ZA), an anti-osteoporosis drug with over 20 years of clinical application, was loaded into MBG/PCL inks for the DIW construction of scaffolds [157]. Owing to the low surface area of scaffolds (1.33 m 2 /g), only 28% of the ZA loaded was slowly released into the degradation media after 4 weeks, and it effectively suppressed the osteoclastic differentiation of murine macrophages at the early stage.
Besides polymer composites, DIW has been frequently applied to fabricate pure BG scaffolds based on an indirect process. Using an ink composed of micro-or nanoparticles of BG and a polymer matrix with proper rheological behavior, the DIW process was performed to fabricate green bodies of scaffolds, which are essentially BG/polymer composites. The green bodies were then subjected to heat treatment for thermal decomposition of the organic phase and sintering densification of the residual glass. Specifically, 6P53B [114], 13-93B [158], or 36B-18Si-2P-6Na-8Mg-8K-22Ca [159] were selected as the BGs in studies, owing to their lower crystallization tendency under high temperatures. The obtained BG objects were further spin-coated or dip-coated with MBG or drug-eluting polymers. The coatings endowed the resultant scaffold with increased surface areas for cell adhesion, lower degradation rates [160], as well as the function to locally deliver protein (BMP-2 [161]), a gene (ss-DNA [161]), and drugs (dexamethasone [161], HYSA [158]) to stimulate bone regeneration. Wang et al. spin-coated a borosilicate BG with a MoS 2 /PLGA solution. [159]. The coated scaffolds exhibited a photothermal response toward the near-infrared laser, demonstrating great potential for healing bone defects resulting from osteosarcoma-removal surgery.
A more recent study reported a novel process that seamlessly combined the sol-gel synthesis of MBG and DIW to prepare pure MBG scaffolds ( Figure 3) [119]. The key component in this process was acrylated F127, which served to direct the formation of a highly ordered mesoporous structure in sol-gel glass and enabled photopolymerization of the ink upon material deposition. The seamlessly additive-manufactured MBG scaffolds exhibited a highly interconnected macroporous structure with well-defined pore size and pore location while maintaining the mesoporous nature of the MBG. When compared to the control group obtained via the decomposition-sintering densification of MBG over PU foams, the seamlessly fabricated MBG scaffold induced a more pronounced osteogenesis rate of BMSC and more efficient new bone formation in a rat calvaria bone defect. The authors attributed the difference in the morphology of the scaffold to that fact that the concave region at the intersection of extrudates may contribute to the topical enrichment of calcium ions, which positively stimulated osteogenic activity, while the higher interconnectivity in the additive-manufactured scaffold favored tissue ingrowth. Owing to the mild conditions during fabrication, DIW with ink that encapsulates living cells is technically feasible. Known as extrusion-based bioprinting, this process aims to fabricate living constructs with a well-defined 3D structure to provide a tissue-mimicking structure for cell culture and cell delivery. DIW with bio-inks containing BG has been reported in several works. Owing to its high rigidity and its ability to release doped elements, BG within the bio-ink was reported to modulate the stiffness of bioprint ink, as well as the responses of cells, endowing the bioprinted construct with proper printability. Meanwhile, the BG may induce a specified biological function, depending on its chemical composition. For instance, Zhu et al. used copper-doped MBG to simultaneously enhance the angiogenetic and osteogenesis activities of stromal stem cells within a bioprinted construct [162]. In contrast, Li et al. reported that the silica-based BG nanoparticles within a gelatin-alginate bio-ink inhibited the angiogenic and osteogenic differentiation of BMSC. Instead, the stem cells maintained active proliferation activity as well as stemness, indicating the potential of the bioprinted constructs for stem cell therapy.
While the functions and benefits of BG incorporation in bio-ink were discussed earlier, it is noted that the addition of rigid BG particles may increase the viscosity of the bioink, and thereby hinder the spreading of cells ( Figure 4A) [163]. In addition, the collision, friction, and steric hindrance of rigid glass particles become more pronounced with increased size and content of the BG, leading to compromised viability and proliferation of cells within the bio-ink [164]. For instance, the addition of 1 wt.% BG microparticles (size = 13-50 µm) into a gelatin-alginate hydrogel ink led to six times greater viscosity at the bioprinting temperature. Consequently, the shear stress during extrusion increased, and the cell viability was markedly lower [154]. However, the addition of BG nanoparticles with an average diameter of 12 nm with doubled concentration (2 wt.%) did not induce evident cell death when compared to the blank gelatin-alginate ink. Nonetheless, with the BG nanoparticle content increased to 5 wt.%, significantly lower cell viability was detected [163]. Owing to the mild conditions during fabrication, DIW with ink that encapsulates living cells is technically feasible. Known as extrusion-based bioprinting, this process aims to fabricate living constructs with a well-defined 3D structure to provide a tissuemimicking structure for cell culture and cell delivery. DIW with bio-inks containing BG has been reported in several works. Owing to its high rigidity and its ability to release doped elements, BG within the bio-ink was reported to modulate the stiffness of bioprint ink, as well as the responses of cells, endowing the bioprinted construct with proper printability. Meanwhile, the BG may induce a specified biological function, depending on its chemical composition. For instance, Zhu et al. used copper-doped MBG to simultaneously enhance the angiogenetic and osteogenesis activities of stromal stem cells within a bioprinted construct [162]. In contrast, Li et al. reported that the silica-based BG nanoparticles within a gelatin-alginate bio-ink inhibited the angiogenic and osteogenic differentiation of BMSC. Instead, the stem cells maintained active proliferation activity as well as stemness, indicating the potential of the bioprinted constructs for stem cell therapy.
While the functions and benefits of BG incorporation in bio-ink were discussed earlier, it is noted that the addition of rigid BG particles may increase the viscosity of the bio-ink, and thereby hinder the spreading of cells ( Figure 4A) [163]. In addition, the collision, friction, and steric hindrance of rigid glass particles become more pronounced with increased size and content of the BG, leading to compromised viability and proliferation of cells within the bio-ink [164]. For instance, the addition of 1 wt.% BG microparticles (size = 13-50 µm) into a gelatin-alginate hydrogel ink led to six times greater viscosity at the bioprinting temperature. Consequently, the shear stress during extrusion increased, and the cell viability was markedly lower [154]. However, the addition of BG nanoparticles with an average diameter of 12 nm with doubled concentration (2 wt.%) did not induce evident cell death when compared to the blank gelatin-alginate ink. Nonetheless, with the BG nanoparticle content increased to 5 wt.%, significantly lower cell viability was detected [163].

Vat Photopolymerization
Vat photopolymerization was first developed in the 1980s by Hull as an AM technology [166]. The feedstock used in the vat photopolymerization process is a liquid mixture of photopolymerizable oligomer/monomers, photopolymerization initiator, and other functional additives (e.g., dispersant, dye, and fillers) [166]. Vat photopolymerization begins with the coating of the deposition plane with a thin layer of liquid feedstock. Next, UV or visible light is introduced onto the deposition plate to initiate photopolymerization, converting the fluid feedstock into a layer of polymerized solid pattern. The deposition plate then travels the distance of one layer, allowing the consolidated layer to be coated by the feedstock for the next layer to be photopolymerized. These steps are repeated until the desired model is fully fabricated through layer-by-layer polymerization.
Two types of vat photopolymerization technology have been widely used to fabricate BG or BG/polymer composites. The first one is stereolithography (SLA), during which a light beam(s) moves across the deposition plane and instantly initiates the photopolymerization of feedstock within the light spot ( Figure 2C). In the digital light processing (DLP) process, the light is introduced onto the digital micromirror device (DMD), which is an array of micromirrors that can be individually controlled by a program to alternate between the "on" and "off" states. This allows a patterned light reflected from the DMD to irradiate the liquid feedstock, thereby initiating the photopolymerization of a whole layer at once ( Figure 2D). Reprinted and from references [154]. (B) Schematic view of light diffraction due to the presence of BG particles, as well as the consequent variation in cured depth and width of the fabricated products. Reprinted and adapted from references [165]. (C-F) Scheme and corresponding SEM images of the SLS-fabricated composites using a simple mixture (C,D) or composite microsphere (E,F) as feedstock. Orange polygonal particles represent BG, and blue spheres represent polymer. Scale bar = 500 µm. Reprinted and adapted from references [113,129].

Vat Photopolymerization
Vat photopolymerization was first developed in the 1980s by Hull as an AM technology [166]. The feedstock used in the vat photopolymerization process is a liquid mixture of photopolymerizable oligomer/monomers, photopolymerization initiator, and other functional additives (e.g., dispersant, dye, and fillers) [166]. Vat photopolymerization begins with the coating of the deposition plane with a thin layer of liquid feedstock. Next, UV or visible light is introduced onto the deposition plate to initiate photopolymerization, converting the fluid feedstock into a layer of polymerized solid pattern. The deposition plate then travels the distance of one layer, allowing the consolidated layer to be coated by the feedstock for the next layer to be photopolymerized. These steps are repeated until the desired model is fully fabricated through layer-by-layer polymerization.
Two types of vat photopolymerization technology have been widely used to fabricate BG or BG/polymer composites. The first one is stereolithography (SLA), during which a light beam(s) moves across the deposition plane and instantly initiates the photopolymerization of feedstock within the light spot ( Figure 2C). In the digital light processing (DLP) process, the light is introduced onto the digital micromirror device (DMD), which is an array of micromirrors that can be individually controlled by a program to alternate between the "on" and "off" states. This allows a patterned light reflected from the DMD to irradiate the liquid feedstock, thereby initiating the photopolymerization of a whole layer at once ( Figure 2D).
The addition of BG particles is known to affect the rheological and optical properties of feedstock, which may affect the microstructure and properties of the produced parts. For instance, an increased content of BG particles led to higher viscosity of the feedstock [167,168]. As such, a longer time was required for the liquid feedstock to cover and level on the deposition plate, otherwise contributing to the uneven thickness of the polymerized layer. Meanwhile, the presence of BG also affects the photopolymerization behavior of feedstock. A higher concentration of BG may intensify the scattering of light within the feedstock, and thus, alter the depth and width of the polymerized structure, leading to a compromised precision of the manufacturing process [165,169] (Figure 4B). Last, but not least, the influence of BG on photopolymerization kinetics shall not be ignored. Par et al. reported a decreased degree of polymerization with increased BG content in photopolymerizable resin, and attributed this effect to the electron transfer to the oxides in BG [170,171]. A lower degree of polymerization and more residual oligomers could lead to lowered mechanical properties of the produced objects and cytotoxic monomer residuals, thereby jeopardizing the safety of the obtained composites for biomedical application.
Two cases reported the use of DLP to prepare BG scaffolds in an indirect manner (green body fabrication-sintering). Su et al. fabricated gyroid scaffolds composed of 45S5 Bioglass ® and biphasic calcium phosphate (BCP) [167]. During sintering, the BG reacted with BCP and hindered the densification of scaffolds, resulting in greater microporosity of the scaffolds. Meanwhile, the reaction products, which include CaSiO 3 and various sodium calcium phosphates, were more reactive than pure BCP in inducing in vitro hydroxyapatite precipitation. Xu et al. used photopolymerizable slurry-like feedstock containing AP40mod glass-ceramic as a feedstock to fabricate BG-ceramic scaffolds via DLP [172]. With endothelial progenitor cells (EPC) and BMSC seeded onto the scaffolds at an optimized ratio (EPC:BMSC = 2:1), more efficient formation of blood vessels and bone was observed at the implant site, suggesting the potential of the DLP-fabricated AP40mod scaffold as a platform for cell-aided bone defect repair.

Powder Bed Fusion
As the name suggests, powder bed fusion is an AM process where feedstock in the form of loose powders is thermally fused into an integral part. As shown in Figure 2E, the process begins with the spreading and rolling of the feedstock powders over a horizontal building platform to obtain a thin powder layer of uniform thickness. A beam of laser or electron is then introduced onto the powder layer. As the beam scans over the powder layer, the irradiated and heated powders fuse to form a sliced cross-section of the model. The building platform then descends by the thickness of the powder layer, and the abovementioned powder spreading-laser scanning is repeated until the desired object is completely fabricated. The entire process is typically performed in a closed chamber with an inert atmosphere to avoid oxygen-aided thermal decomposition of the materials [173].
Among the powder bed fusion technologies developed, selective laser sintering (SLS) has been widely recruited to produce BG and its composites. This technology utilizes a CO 2 laser beam as a heat source to partly melt the feedstock powders for fusion. Regarding the fabrication of BG/polymer composites, a low-power (<5 w) laser of high scanning speed (100-1000 mm/s) is typically used, and the SLS process occurs in an inert atmosphere. This configuration is considered to prevent thermal degradation and oxidation of the polymer phase. The feedstock powders are most commonly prepared by blending the fine particles of the BG and the matrix polymer [174]. This process typically yields a highly rough and porous surface over the fabricated composites ( Figure 4C,D). In comparison, when composite microspheres or polymer-coated BG particles serve as feedstock, the BG and polymer are more closely integrated at the interface, and the surface of the obtained composites is much smoother ( Figure 4E,F), contributing to reduced stress concentration and enhanced mechanical properties of the SLS-fabricated objects [175].
As the sintering of the polymer phase has a low demand for energy input, the transient heating generated by the low-powered, fast-scanning laser may offer energy that is sufficient for polymer sintering without thermally decomposing other organic compounds. Using composite microspheres containing 10 wt.% MBG nanoparticles, dexamethasone, and PLLA, Sun et al. prepared dexamethasone-eluting MBG/PLLA scaffold via SLS for bone regeneration [113]. Dexamethasone was continuously released in 4 weeks, leading to evident osteogenesis of BMSC cultured in a dexamethasone-deficient culture medium, and more pronounced formation of blood vessels and bones in animal models. Shuai and colleagues used polydopamine (PDA)-coated MBG as reinforcement in a PLLA-based polymer composite. The coating applied led to strengthened bonding at the MBG/polymer interface, more homogeneous distribution within the polymer matrix, and improved surface hydrophily of the scaffolds, thereby enhancing the compressive mechanical properties of the scaffolds as well as the adhesion profile of MG63 osteosarcoma-like cells [176]. The PDA coating was further exploited as a reductive agent, allowing graphene oxide and Ag + to be (partially) reduced and immobilized over the surface of PDA-MBG [125,174,177]. These works demonstrate the versatility of SLS-fabricated scaffolds containing PDA-MBG for bone regeneration, with enhanced mechanical properties and additional antibacterial functions.
Alternatively, BG particles without polymer could be directly produced into BGceramic constructs via SLS. Under laser beam radiation, the heated glass particles develop into a viscous flow to fuse with the surrounding glass particles or the substrate below. As no polymer binder is used, the obtained part does not require a heat treatment (typically pyrolysis) to remove the binder, thereby being free from potential contamination of the residual binder and shrinkage of size. For instance, Cao et al. successfully produced fully amorphous 13-93 BG scaffolds via SLS [178]. To sinter the BG particles (~100 nm), which require a significantly high temperature to be softened, the laser applied had elevated power (5-9 W) and a lower scanning speed (100 mm·min −1 ) compared to the parameters applied in the SLS of polymer composites (e.g., 5 W/40 mm·s −1 [129], 0.09-0.2 W/1 mm· −1 [113], and 2.3 W/100 mm·s −1 [174]). Using the same device, researchers also succeeded in the direct SLS fabrication of 45S5 [179] and 58S BG [180].
Inevitably, the high-powered, slow scanning laser applied during SLS increases the tendency of thermal crystallization of BG, which leads to impaired bioactivity compared to amorphous BG [181,182]. Whether the BG crystallizes after direct SLS depends on the temperature profile, as a high temperature is required to make the BGs sufficiently soft to fuse with each other, and the transient temperature is likely to be higher than the onset temperature of crystallization (T c ). For instance, it is reported that 45S5 glass scaffolds produced via direct SLS were partially crystallized. This is believed to be related to a narrow sintering window (T w , which is approximately 87 • C for 45S5), namely, the difference in temperature between the glass transition temperature and T c [179,183]. In contrast, 13-93 glass has a wide sintering window (~100 • C) and a high T c (~825 • C), and the glass is believed to be softened enough before reaching T c so that glass fusion without crystallization can occur [184,185]. Regardless of the wide T w , crystallization/devitrification of an amorphous glass may still occur considering that the transient temperature during laser processing is above the T c . Rodrigo-Vázquez et al. reported that crystalline peaks corresponding to pseudowollastonite were detected in the additive-manufactured 62W glass (T w ≈ 150 • C) scaffold [186]. As the 62W glass underwent devitrification after at 1 h of heat treatment under 900 • C, it is believed that the transient temperature during laser processing was much higher than 900 • C to induce the partial crystallization of 62 W glass in a short period of laser irradiation [187]. Therefore, a systematic analysis of the temperature-viscosity profile of BG and meticulous configuration of the SLS processing parameters are required for the direct SLS of a fully amorphous BG.
In summary, numerous AM technologies have been successfully applied to fabricate BG or BG/polymer composite scaffolds for bone tissue engineering. A comparison of these technologies is presented in Table 3. The significant variance in both the geometry of bones and the demographic characteristics of patients emphasizes the need for patient-specific bone defect treatments that, to a great extent, rely on a customized geometrical design of the material filling the defect. On one hand, anatomical fitting along with a maximized contact area between the defect and the implanted material provides optimal post-implantation stability, which prevents the undesired dislodgement and loosening of the implant. On the other hand, the geometry of implanted material determines the aesthetic aspects of bone defect healing. This point is especially valuable in the repair of craniomaxillofacial bone defects, as the highly unique yet complex geometry of bones conventionally calls for meticulous intraoperative shaping of the implanted material, making the operation more time-consuming and technically challenging. Owing to its outstanding capability to fabricate complex geometries, AM has become a powerful tool for the fabrication of patient-specific implants. Based on medical images, computer-aided design, and finite element analysis, scaffolds may not only end up with a structural geometry that recapitulates the original bone, but also a customized topological design that restores the biomechanical loading behavior of the original bones [23].
Evidence that BG-containing tissue engineering scaffolds with customized geometry led to efficient bone defect healing in animal models was recently reported by Han et al., who prepared a borate-based BG/PCL composite scaffold with a case-specific design via SLS ( Figure 5) [188]. To recapitulate the geometry of actual bones, the radii of rabbits were subjected to CT scanning, and the resultant images were converted into 3D models.
A Boolean crossover operation between the 3D radii models and porous body-centered cubic units was then performed, producing 3D models with customized geometries and a microporous structure capable of inducing bone ingrowth. Using a mixture of borate-based BG microparticles and PCL powders as feedstock, customized tissue engineering scaffolds were fabricated via SLS and finally implanted into bone defects in rabbit radii. When compared to the control group, where the osteotomy sites were left blank, the presence of customized scaffolds, regardless of the material composition, induced bone regeneration into the interconnecting pores while following the geometry of the customized scaffolds. Moreover, scaffolds containing 20 wt.% BG were most effective at inducing osteogenesis and angiogenesis at the defect site, while the growth of fibrotic tissues (a sign of a foreign body reaction) was minimal around the scaffold, possibly owing to the optimal dosage of ions released by the scaffolds. The authors also emphasized the role of BG content within the composites, reporting that a BG content of 40 wt.% resulted in reduced viability and ALP activities in human BMSC, which may be attributed to the excessively high pH of the extracellular environment after the degradation of scaffolds.

Scaffold with the On-Demand Spatial Distribution of Biomaterials
The flexibility of AM can also be utilized to control the spatial distribution of materials within a single object, yielding bone tissue engineering scaffolds with heterogeneous porosity, and thus, tailored mechanical properties and degradational behavior in different regions. More recently, the development of multi-material and multi-disciplinary/hybrid AM has brought further possibilities to combine different biomaterials into a single object, unchaining the potential of a tissue engineering scaffold that is conventionally limited to a single material composition. Here, we introduced two cases where multi-material AM of BG-containing materials was performed to fabricate scaffolds with the heterogeneous spatial distribution of different biomaterials.
Degenerative disease of the joints results in damage to the articular cartilage and, if not treated promptly, defects in the subchondral bone. The distinctive characteristics of the two neighboring tissues call for a dual-or multi-component tissue engineering construct that facilitates the regeneration of cartilage, subchondral bone, and, ideally, the interfacial tissue (calcified cartilage) in between [189]. For this purpose, Gao et al. reported the use of DIW to fabricate dual-module scaffolds ( Figure 6) [117]. The polymer phase was composed of poly(N-acryloyl 2-glycine) (PACG) and methacrylated gelatin (GelMA). By tuning the concentration of the two polymers, inks with tailored mechanical properties and degradation rates were obtained to satisfy the need for cartilage regeneration (soft matrix, rapid degradation) and bone regeneration (stiff matrix, slow degradation). Furthermore, MnCl 2 and MBG nanoparticles were selectively loaded into the optimized inks for cartilage repair and bone regeneration, respectively. Single-material scaffolds prepared using Mn 2+ -releasing inks effectively upregulated the chondrogenesis-related genes of hBMSC, whereas those prepared with BG-containing inks induced more pronounced osteogenesis. Next, the two functional inks were used to fabricate the dual-module scaffolds via DIW, with the chondrogenic module deposited on top of an osteogenic module. Following implantation into rat knees, histological sections showed that both the cartilage and subchondral bone regenerated robustly. This work not only demonstrates the potential of the dual-layered scaffold in the repair of cartilage/bone defects at the articular joints, but more importantly, emphasizes the utility of AM to fabricate multi-material, multi-structural constructs intended for multi-tissue regeneration. Bioengineering 2023, 10, x FOR PEER REVIEW 25 of 42

Scaffold with the On-Demand Spatial Distribution of Biomaterials
The flexibility of AM can also be utilized to control the spatial distribution of materials within a single object, yielding bone tissue engineering scaffolds with heterogeneous porosity, and thus, tailored mechanical properties and degradational behavior in different Reprinted and adapted from reference [117].
The incorporation of BG into bioprinting ink enables the responses of living cells (e.g., the maintenance of stemness [163], the upregulation of proliferation, and osteogenic differentiation [162]) within a 3D construct to be modulated. Nonetheless, the increased stiffness of the matrix, collision/friction between cells and rigid BG particles, and the increased shear stress within the flowing ink, which are caused by the presence of a rigid BG within the bio-ink, have been shown to impair cell viability [154]. This conflict was smartly resolved in a recent study [190]. As shown in Figure 7, the authors employed MBG/PCL as the feedstock for the FDM fabrication of a porous 3D framework. Once a layer of the framework was fabricated, bioprinting with a BG-free bio-ink was seamlessly performed. Thus, the MBG modulated the cellular responses "remotely" through its degradation products, while the cells were protected from the adverse effect due to the presence of rigid BG particles within the bio-ink. As a proof-of-concept, the authors designed The incorporation of BG into bioprinting ink enables the responses of living cells (e.g., the maintenance of stemness [163], the upregulation of proliferation, and osteogenic differentiation [162]) within a 3D construct to be modulated. Nonetheless, the increased stiffness of the matrix, collision/friction between cells and rigid BG particles, and the increased shear stress within the flowing ink, which are caused by the presence of a rigid BG within the bio-ink, have been shown to impair cell viability [154]. This conflict was smartly resolved in a recent study [190]. As shown in Figure 7, the authors employed MBG/PCL as the feedstock for the FDM fabrication of a porous 3D framework. Once a layer of the framework was fabricated, bioprinting with a BG-free bio-ink was seamlessly performed. Thus, the MBG modulated the cellular responses "remotely" through its degradation products, while the cells were protected from the adverse effect due to the presence of rigid BG particles within the bio-ink. As a proof-of-concept, the authors designed 3D scaffolds where the degradation of MBG triggered functional expression of the encapsulated cells. The stem cells within the bio-ink were transfected with a lentiviral vector harboring Tet-on-BMP2, thereby enabling controlled transcription of the downstream BMP2 gene in the presence of doxycycline, which was loaded into MBG embedded in the MBG/PCL framework. The results showed that doxycycline, a wide-spectrum antibacterial agent, was continuously released from the MBG/PCL framework and inhibited the growth of pathogens typically related to orthopedic surgeries, thereby rescuing murine stem cells from suppressed survival and proliferation in the presence of MRSA. Meanwhile, BMP-2 transcription was dramatically enhanced in stem cells cultured with doxycycline-eluting MBG/PCL, thereby enhancing the osteogenic activities of stem cells. The hybrid constructs were found to induce ectopic bone formation after subcutaneous implantation, with a greater amount of bone and fewer bacteria present in the surrounding tissues relative to the control group (MBG without doxycycline uptake). With the ability to simultaneously stimulate bone regeneration and prevent bacterial infection, the hybrid construct offers a promising solution to the clinical treatment of large-sized bone defects where bacterial infection typically impairs the efficacy of therapy.  Red boxes indicate tissues adjacent to remaining scaffolds that are examined under higher magnification. Reprinted and adapted from reference [190].

Perspectives on Future Research
The case studies elaborated upon in Section 4 demonstrate the value of AM in scaffold fabrication, which includes the ability to prepare BG or BG/polymer composite scaffolds of the desired shape, the on-demand spatial distribution of biomaterials, and a well- Red boxes indicate tissues adjacent to remaining scaffolds that are examined under higher magnification. Reprinted and adapted from reference [190].

Perspectives on Future Research
The case studies elaborated upon in Section 4 demonstrate the value of AM in scaffold fabrication, which includes the ability to prepare BG or BG/polymer composite scaffolds of the desired shape, the on-demand spatial distribution of biomaterials, and a welldefined porous structure. These structural features play important roles in determining the physiochemical properties of scaffolds that affect their performance in clinical bone defect healing. The relationships between structure, physiochemical properties, and biological function will be, in our view, a key focus of future research on the AM of BG and its composites. This requires the development of AM technologies that are compatible with BG or BG/polymer feedstocks. These advancements are anticipated to remove the technical barrier of fabricating tissue engineering scaffolds with complex structures, which essentially determines the biological function of the scaffolds to realize safer, more effective, and more patient-specific therapy. Here, we summarize our perspectives on how future studies may be directed in these areas.

Toward a Higher Spatial Resolution
A higher spatial resolution of the manufacturing process enables precise tuning of the structural properties of a BG-containing tissue engineering scaffold. Spatial resolution down to the submicron and even the nanometer scale is useful in the precise control of porous structures, which enables the degradation profile of the BG to be finely adjusted. Moreover, the high-precision manufacturing process imparts cell-sensible topological cues to the surface of scaffolds; these topological cues can be exploited for the rapid (within several hours of contact) induction of cell morphology, and afterward, cellular responses, presumably before a critical concentration of ions are released through BG degradation [191].
Recently, the two-photon polymerization (2PP) technique has been successfully applied in the fabrication of glass with fine structures at the submicron scale, catching the interest of researchers. During the 2PP process, the photopolymerization initiator is attacked by two photons, generating free radicals in a highly localized region near the laser focal spot to initiate the polymerization process, thereby achieving a high spatial resolution of the manufacturing process ( Figure 8A). Based on a 2PP AM-thermal debinding and sintering process, Kotz et al. prepared micro-structured models, such as a micro-lens and filtering elements with approximately 55-µm pores ( Figure 8B) [192]. Further optimization of the feedstock pushed the resolution of the 2PP process to the sub-200 nm scale ( Figure 8C) [193]. With the silica nanoparticles replaced by BG nanoparticles, it is possible to fabricate bioactive BG or BG/polymer scaffolds with submicron structures, which potentially enables more complex and precise control of the physiochemical properties and biological functions of the resultant scaffolds.

Binder-Free AM of Pure BG Objects
Owing to their direct bone-bonding ability, pure BG scaffolds continue to hold great research interest for bone defect treatment. Regarding the AM of pure BG parts, the "indirect AM", during which the additive-manufactured BG/polymer green bodies are subjected to binder removal and glass sintering at high temperatures, remains the most commonly applied routine. Nonetheless, this process is not only tedious, but is also linked to numerous issues. During sintering, the as-fabricated composite scaffolds undergo significant shrinkage, leading to distortion of the porous structure and deviation in its size [147]. Moreover, the high temperature applied results in the generation of cracks [194] and carbonaceous residuals that are difficult to thermally decompose [195]. In the context of bone tissue engineering, these problems may jeopardize the geometrical fidelity, loadbearing capacity, and biocompatibility of the resultant BG scaffolds. As a result, the effectiveness of bone defect healing is significantly compromised.
Recently, the direct deposition of molten glass has been reported in several studies. Zaki et al. investigated the melt extrusion of phosphate glass with low glass transition temperatures ( Figure 8D,E). The melt-quenched phosphate glasses were thermally drawn into a rod-like preform with a diameter of 1.90 mm. A desktop FDM 3D printer equipped with a high-temperature (~500 • C) extruder and build plate was used to additive manufacture pure phosphate glass models in FDM mode, with the layer resolution reduced to 100 µm [196]. In another work, Liu et al. investigated the melt deposition of silica glass, a material typically requiring high temperatures (>1000 • C) to process [197,198]. As shown in Figure 8F, the feedstock was a fused silica glass filament with a diameter of 196 µm, while four CO 2 laser beams served as an energy source, focusing on the tip of the filament to locally melt the glass. With fine-tuning of the laser power and the speed of feeder movement, the width of the deposited line was controlled at 240-330 µm, and a 10.5 × 3.5 × 21-mm 3 (L × D × H) prism with no built-in stress was successfully produced. While the filaments with a diameter of 200-1900 µm required a long time to melt and led to a slow material deposition rate (100 mm/min) applied in the process, Spirrett et al. developed a system that continuously jets fine glass powders (D 50 = 45 µm) onto a building platform, followed by irradiation with a continuous-wave, ytterbium-doped fiber laser ( Figure 8G) [199]. After optimization of the laser power and the glass feeding rate, the laser scan speed was increased to~700 mm/min, which significantly improved productivity. When compared to earlier works on the melt deposition of glass, the technologies developed in these recent studies seem superior in terms of spatial resolution [198]. Although none of these studies used BG as a raw material, it is possible that by meticulously tailoring the glass formula, BGs with desired biological functions, along with proper thermal and rheological properties, could be acquired to gain compatibility with these novel methods.  [193]. (B) Photographs of silicate-based glass object produced via 2PP. Scale bar = 500 µm [192]. (C) Scanning electron microscopy image of silicate-based glass objects produced via 2PP. Scale bar = 20 nm (image) and 1 nm (inset) [193]. (D,E) Photographs of phosphate glass preform, and a scheme of direct FDM of phosphate glass [196]. (F,G) Scheme of laser-assisted direct deposition of molten glass using (F) glass rod [197] and (G) jets of glass particles [199] as feedstock. Reproduced from references listed above.

Scaffold for the Regeneration of Multiple Tissues at the Bone Defect
The healing of bone defects, especially those caused by trauma, is far more complex than a simple regeneration of bone. Typically, a bone defect involves damage to the bones as well as the surrounding bone-attaching tissues, including the periosteum [200], articular cartilage [201], and ligaments [202]. Therefore, complete healing of the bone defect requires the regeneration of multiple tissues affected and, more importantly, the restoration of a biological bonding of different tissues at their interfaces.  [193]. (B) Photographs of silicate-based glass object produced via 2PP. Scale bar = 500 µm [192]. (C) Scanning electron microscopy image of silicate-based glass objects produced via 2PP. Scale bar = 20 nm (image) and 1 nm (inset) [193]. (D,E) Photographs of phosphate glass preform, and a scheme of direct FDM of phosphate glass [196]. (F,G) Scheme of laser-assisted direct deposition of molten glass using (F) glass rod [197] and (G) jets of glass particles [199] as feedstock. Reproduced from references listed above.

Scaffold for the Regeneration of Multiple Tissues at the Bone Defect
The healing of bone defects, especially those caused by trauma, is far more complex than a simple regeneration of bone. Typically, a bone defect involves damage to the bones as well as the surrounding bone-attaching tissues, including the periosteum [200], articular cartilage [201], and ligaments [202]. Therefore, complete healing of the bone defect requires the regeneration of multiple tissues affected and, more importantly, the restoration of a biological bonding of different tissues at their interfaces.
For this purpose, multiple "modules" are expected to be integrated into a single scaffold, with each module mimicking the material composition, microstructure, and cell phenotypes, which are selectively determined to upregulate the regeneration of the targeted tissues. A more challenging issue is constructing interfaces between different modules. It is at the limited region close to the interface that the abovementioned properties display a gradient transition, which prevents an abrupt change in the mechanical properties, thereby ensuring an effective load transfer and mechanical stability without stress concentration [202]. Preferably, different modules should be fused or chemically linked to each other at the interface, which offers stronger bonding strength compared to a stratified structure where different modules are simply stacked together [203].
AM with BG or BG/polymer composites offers an opportunity to satisfy the abovementioned requirement. Because of its high flexibility, AM has long been exploited to produce functionally graded materials. With the spatial distribution of various feedstocks, a monolithic scaffold with gradient porosity and material composition is yielded, with the degradation rate, permeability, and load-bearing capacity locally tuned [204][205][206]. Thus, the possibility further multiplies with BG added into the feedstock. As the type and ratio of functional elements in BG can be flexibly tuned, BGs with variable degradation rates and element release profiles can be obtained through this approach, which enables selective stimulation of the regeneration of different targeted tissues. Finally, the advent of multimaterial AM, as well as hybrid AM, further pushes the boundaries of complexity in the composition and structure of BG-containing scaffolds [117,190]. With recent studies reporting the function of BG in stimulating the regeneration of tendons [207] and cartilage [208], additive-manufactured scaffolds doped with different BGs seem to hold much potential for the regeneration of multiple tissues in bone defects.

Conclusions
The current study offers an overview of additive-manufactured BG or BG/polymer composites as bone tissue engineering scaffolds. AM technologies based on melt extrusion, DIW, vat photopolymerization, and powder bed fusion have been successfully applied to process feedstock containing BG and polymeric binders. With the selection of AM for fabrication, the resultant BG or BG/polymer composites present well-defined geometries and intricate porous structures. The BGs within the additive-manufactured parts maintained their biological function (e.g., upregulating osteogenesis inducing bone-binding bioactivity), while the well-defined shape, size, and porous structure satisfy the need for anatomic fitting of the bone defect and effective bone tissue ingrowth. Specifically, the advent of bioprinting technology allows, for the first time, a viable construct containing BG to be prepared, showing significant potential for stem cell-based therapy for bone defects.
A unique advantage of AM is that it allows a scaffold with a complex shape, intricate porous structure, and even multiple material compositions to be fabricated with highprecision. In this regard, we have proposed three directions that call for future research: improving the resolution of the AM process to a submicron scale, applying binder-free AM technologies to fabricate pure BG objects, and fabricating multi-module scaffolds that stimulate the regeneration of multiple tissues in the bone defects. With more effort applied to these research areas, additive-manufactured BG-containing scaffolds will undoubtedly become more valuable to induce the safe and efficient healing of bone defects.