Three-dimensionally printed mesoporous bioactive glass for craniomaxillofacial bone regeneration: Material evolution,functional mechanisms,and clinical translation

The clinical burden of craniomaxillofacial bone defects

The craniomaxillofacial (CMF) skeleton is a complex anatomical structure that not only dictates facial esthetics and identity but also serves critical physiological functions, including mastication, deglutition, phonation, and the protection of vital sensory organs ¹ . However, the integrity of this system is frequently compromised by a myriad of pathological conditions. Epidemiologically, the burden of bone defects is staggering ² . It is estimated that over 20 million patients worldwide suffer from devastating defects caused by trauma, tumor resection, infection, or congenital anomalies annually, with a global market for bone graft substitutes valued at approximately USD 3.1 billion in 2022 and projected to reach USD 4.6 billion by 2030^3,4.

Trauma remains a primary etiology. With the rising incidence of traffic accidents and sports-related injuries, facial fractures account for a significant proportion of emergency hospitalizations ³ . Furthermore, alveolar bone atrophy resulting from severe periodontitis—a condition affecting nearly 10–15% of the global adult population ⁵ —presents a ubiquitous challenge in restorative dentistry. Following tooth extraction, the alveolar ridge undergoes irreversible resorption, with vertical and horizontal bone loss reaching up to 40–60% within the first 2 years if left untreated ⁶ . This volumetric deficiency severely compromises the feasibility of dental implant placement, necessitating complex augmentation procedures ⁷ . In addition, congenital deformities such as cleft lip and palate, which occur in approximately 1 in 700 live births globally, require longitudinal, multi-stage bone grafting interventions throughout the patient’s growth and development ⁸ . These conditions not only impose a heavy economic burden on healthcare systems but also inflict profound psychological distress and quality-of-life deterioration on patients due to facial disfigurement and functional impairment.

The “gold standard” and its intrinsic limitations

To address these defects, the reconstruction of bone tissue requires a therapeutic strategy that provides mechanical stability while inducing osteogenesis. Historically and currently, autologous bone grafting (autografting) is widely regarded as the “gold standard” in clinical practice ⁹ . Harvested typically from the patient’s own iliac crest, fibula, rib, or calvarium, autografts possess the unique “triad” of essential properties: osteoconductivity (providing a scaffold), osteoinductivity (containing growth factors such as bone morphogenetic proteins [BMPs]), and osteogenicity (containing viable osteogenic cells) ¹⁰ .

Despite its superior biological performance, the clinical application of autografts is plagued by significant drawbacks. The most prominent limitation is donor-site morbidity, which is a direct consequence of the harvesting procedure. Clinical studies have reported complication rates at the iliac crest donor site ranging from 19.4% to 31% ¹¹ . These complications include chronic pain, deep infection, hematoma, nerve injury (e.g., paresthesia of the lateral femoral cutaneous nerve), and cosmetic defects. Furthermore, the quantity of harvestable bone is finite, creating a supply-demand mismatch for massive defects (e.g., subtotal maxillectomy). Even when successfully implanted, autografts face the risk of unpredictable resorption, with volume loss reported as high as 50% in the long term, potentially leading to graft failure and the need for revision surgeries ¹² .

Alternatives: Allografts and xenografts

To circumvent the morbidity associated with autografts, allogenic grafts (from human donors) and xenogenic grafts (derived from species such as bovine or porcine sources) have been extensively utilized^10,13. While these alternatives eliminate the need for a secondary surgical site and offer unlimited supply, they are not without shortcomings. Allografts carry a risk, albeit low, of disease transmission (e.g., HIV, hepatitis) and may elicit an immune response leading to graft rejection or encapsulation ¹⁰ . Moreover, the rigorous sterilization processes (e.g., gamma irradiation, freeze-drying) required to ensure safety inevitably devitalize the tissue, destroying osteogenic cells and denaturing osteoinductive proteins, thereby reducing the graft to a primarily osteoconductive scaffold with reduced mechanical strength^10,14.

Similarly, xenografts (e.g., deproteinized bovine bone mineral, DBBM) are widely used in dentistry due to their structural similarity to human bone mineral ¹⁵ . However, their high crystallinity often leads to an extremely slow degradation rate ¹⁶ . Clinical observations indicate that xenograft particles can remain encapsulated in the host tissue for years without being replaced by new bone, potentially interfering with the osseointegration of dental implants and altering the biomechanics of the regenerated region ¹⁷ . Consequently, there is an urgent and unmet clinical demand for synthetic, “off-the-shelf” biomaterials that are biocompatible, biodegradable, bioactive, and mechanically tailored to specific defects.

The evolution of bioactive glass

From 45S5 to mesoporous variants among synthetic bone substitutes, bioactive glasses (BGs) have garnered immense attention since the pioneering discovery of the 45S5 Bioglass^® by Larry Hench in 1969 ¹⁸ . Unlike inert bioceramics (e.g., alumina/zirconia) that trigger fibrous encapsulation, BGs are capable of forming a strong, interfacial bond with host bone ¹⁹ . This integration is mediated by a rapid sequence of surface reactions: ion exchange (Na⁺/Ca²⁺ with H⁺), formation of a silica-gel layer, and subsequent precipitation of an amorphous calcium phosphate layer that crystallizes into hydroxyapatite (HA), mimicking the mineral phase of natural bone ²⁰ . Furthermore, the dissolution products of BGs (e.g., Si⁴⁺, Ca²⁺ ions) have been proven to biologically stimulate osteoblast (OB) gene expression ²¹ .

However, traditional melt-derived BGs suffer from inherent limitations, including low porosity, low specific surface area (<10 m²/g), and a tendency to crystallize during thermal processing, which diminishes bioactivity ²² . To overcome these hurdles, mesoporous bioactive glasses (MBGs) were developed in the early 2000s using supramolecular surfactant templates (e.g., P123, F127) via sol-gel chemistry ²³ . MBGs are characterized by a highly ordered nanoporous channel structure (pore size 2–50 nm) and an exceptionally high specific surface area (300–1000 m²/g) ²⁴ . This unique texture significantly enhances in vitro apatite-forming ability (often within hours) and provides a versatile platform for the loading and controlled delivery of therapeutic drugs, growth factors, or antibacterial agents, representing a quantum leap in bone tissue engineering materials ²⁵ . Figure 1 illustrates the structural evolution from traditional BG to 3D-printed mesoporous scaffolds.

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Figure 1.

Schematic illustration of the structural evolution from traditional bioactive glass to 3D-printed mesoporous scaffolds. (a) Traditional 45S5 bioglass: characterized by a dense, non-porous surface with limited specific surface area, restricting its drug-loading capacity. (b) Mesoporous bioactive glass (MBG): synthesized via sol-gel templates, featuring highly ordered mesoporous channels (2–50 nm). This nanotexture significantly increases surface area, enabling the encapsulation and controlled release of therapeutic molecules. (c) Three-dimensionally printed MBG scaffold: by processing MBG into a 3D construct, a hierarchical pore architecture is achieved. The macroscopic pores (>300 μm) facilitate cell migration and vascular network formation (angiogenesis), while the intrinsic mesopores regulate drug delivery and ion release, realizing the synergy of structural support and biological functionalization.

The intersection of MBG and 3D printing

While MBG powders demonstrate excellent bioactivity, their translation into clinically usable 3D scaffolds has been hindered by the material’s inherent brittleness and the difficulty of shaping them into complex, patient-specific geometries required for CMF reconstruction ²⁶ . Traditional fabrication methods (e.g., polyurethane foam replication) often yield scaffolds with poor mechanical reliability and uncontrolled pore architecture^27,28.

The advent of additive manufacturing (AM), or 3D printing, has revolutionized this landscape ²⁹ . By integrating patient imaging data (computed tomography (CT)/magnetic resonance imaging (MRI)) with computer-aided design (CAD), 3D printing enables the fabrication of patient-specific implants (PSIs) that precisely fit irregular defects ²⁹ . More importantly, 3D printing allows for the construction of scaffolds with hierarchical porosity: macroscopic pores (>300 μm) to facilitate vascular ingrowth and cell migration, and mesoscopic pores (2–50 nm) to enhance surface bioactivity and drug delivery^{22,27,30–32}. When coupled with functional ion-doping (e.g., Sr²⁺, Cu²⁺, Ag⁺), 3D-printed MBG (3DPMBG) scaffolds emerge as a multifunctional platform capable of osteoinduction, angiogenesis, and immunomodulation^33–36.

Scope of this review

This review aims to provide a comprehensive and critical analysis of the state-of-the-art in 3DPMBGs for CMF regeneration. We will systematically cover: (1) the material science evolution from traditional BG to MBG; (2) the adaptability and optimization of various 3D printing technologies (direct ink writing (DIW), stereolithography (SLA), digital light processing (DLP), selective laser sintering (SLS), fused deposition modeling (FDM)) for MBG; (3) the biological mechanisms underlying osteogenesis, angiogenesis, and immune regulation; and (4) the current status of preclinical validation and the roadmap for clinical translation. Particular emphasis will be placed on the challenges of regulatory approval and standardization, identifying the gaps that must be bridged to move this promising technology from the bench to the bedside.

Search strategy

An extensive electronic literature search was performed across primary scientific databases, including PubMed, Web of Science, Scopus, and Google Scholar. The search timeframe encompassed articles published from inception up to February 2026, with a particular emphasis on publications from the last decade to capture the latest advancements in AM. The search strategy utilized Boolean operators (AND, OR) to combine the following key terms: (“Mesoporous bioactive glass” OR “MBG” OR “mesoporous glass”) AND (“3D printing” OR “additive manufacturing” OR “rapid prototyping” OR “bioprinting”) AND (“craniomaxillofacial” OR “maxillofacial” OR “oral” OR “dental” OR “alveolar” OR “bone regeneration” OR “osteogenesis” OR “angiogenesis” OR “macrophage polarization”).

Inclusion and exclusion criteria

To ensure the relevance and quality of the synthesized data, strict inclusion and exclusion criteria were applied.

Inclusion criteria: (1) Peer-reviewed original research articles (both in vitro and in vivo studies); (2) studies specifically utilizing 3DPMBG-based scaffolds (pure or composites); (3) studies reporting on bone tissue engineering, particularly in CMF or load-bearing applications; and (4) articles published in English.

Exclusion criteria: (1) non-English publications; (2) conference abstracts, editorials, or unpublished reports lacking peer review; (3) studies focusing solely on traditional melt-derived BGs (e.g., 45S5) without a mesoporous structure or 3D-printed architecture; and (4) studies exclusively investigating soft-tissue repair without a bone-regeneration component.

Quality assessment and data extraction

The quality of the included preclinical studies was critically evaluated based on experimental design, the presence of appropriate control groups (e.g., autografts, blank defects, or traditional bioceramics), and the reproducibility of the 3D printing parameters. Furthermore, the translational potential of the studies was assessed by prioritizing in vivo experiments that utilized critical-sized defect models (e.g., rat calvarial, rabbit femoral, or porcine models) and provided robust histomorphometric and radiological (micro-CT) quantitative data.

DIW: The rheological balancing act

DIW, also known as robocasting, is currently the most widely investigated technique for fabricating MBG scaffolds due to its cost-effectiveness, material versatility, and ambient processing conditions^39–41.

Vat photopolymerization (SLA and DLP): Precision meets chemistry

SLA and DLP represent the pinnacle of resolution in ceramic 3D printing. These techniques utilize a photosensitive resin loaded with MBG particles, which is cured layer-by-layer using an ultraviolet (UV) laser (SLA) or a projected light mask (DLP)^52,53.

SLS: The thermal dilemma

SLS employs a high-energy laser (CO₂ or fiber laser) to selectively fuse particles in a powder bed. Unlike DIW and SLA, SLS does not necessarily require a binder, offering a direct path to pure ceramic scaffolds ⁶⁴ .

FDM: Composite solutions

FDM (or FFF) is the most accessible 3D printing technology, typically using thermoplastic filaments, such as polycaprolactone (PCL) and poly(lactic acid) (PLA). For bone regeneration, MBG is incorporated as a filler to create polymer/ceramic composites^66,67.

Comparative analysis of AM technologies

The selection of a specific 3D printing technology for CMF regeneration is a decision matrix involving resolution, mechanical requirements, and biological functionality. Table 1 summarizes the key technical parameters, reported mechanical and biological outcomes from the literature, and clinical applicability of these methods to provide a comprehensive comparative analysis.

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Table 1.

Quantitative and comparative analysis of mainstream additive manufacturing technologies for MBG-based scaffolds.

Feature	DIW	Vat photopolymerization (SLA/DLP)	Selective laser sintering (SLS)	FDM
Principle	Extrusion of viscoelastic paste	Light-curing of photosensitive resin	Laser fusion of powder bed	Extrusion of molten thermoplastic
MBG solid loading (%wt)26,45	40–75% (high loading in hydrogel/paste)	30–55% (limited by resin viscosity & light scattering)	80–100% (powder bed, binderless possible)	10–40% (limited by filament brittleness)
Resolution27,52	200–500 μm	25–100 μm	100–300 μm	200–400 μm
Print speed	Extrusion speed: 10–50 mm/s	Vertical build: 10–50 mm/h	Laser scan: 500–2000 mm/s	Extrusion speed: 10–60 mm/s
Mechanical strength (compressive) 26	2–20 MPa (highly dependent on porosity)	10–50 MPa (dense struts post-sintering)	5–30 MPa (residual internal porosity)	5–15 MPa (modulus: 0.5–2 GPa)
Bioactivity (cell viability)	>95% (room temp, highly biocompatible)	>85% (requires rigorous post-wash to remove toxic monomer)	>85% (crystallization slightly reduces ion release)	>90% (polymer masking delays initial interaction)
In vivo performance (estimated BIC)	~45–60% at 8–12 weeks	~40–55% at 8–12 weeks	~35–50% at 8–12 weeks	~30–45% (delayed onset)
Drug-loading capacity 31	High (up to 10–20 %wt, mixed directly into paste)	Very limited (<2 %wt, interferes with UV curing)	None (0%, extreme heat destroys biologics)	Limited (<5 %wt, thermal degradation risk)
Key limitation	Nozzle clogging; die swell effect	Resin toxicity; refractive index mismatch	Thermal devitrification (crystallization)	Polymer masking effect; low ceramic loading

DIW, direct ink writing; SLA, stereolithography; DLP, digital light processing; SLS, selective laser sintering; FDM, fused deposition modeling; BIC, bone-implant contact. The reported quantitative parameters (e.g., %wt, mechanical strength, speed, and biological outcomes) represent typical generalized ranges synthesized from the cited literature and may vary based on specific MBG formulations, hierarchical pore designs, and specific 3D printer hardware.

Summary and selection strategy

For alveolar ridge preservation where geometry is small but intricate, DLP is increasingly preferred due to its ability to print highly permeable gyroid structures that facilitate rapid vascularization.

Conversely, for large-volume mandibular defects requiring significant drug delivery (e.g., antibiotics for osteomyelitis), DIW remains the gold standard for drug and cell delivery. This is fundamentally dictated by its material-processing conditions. Unlike SLS or FDM, which involves extreme thermal processing that denatures delicate biologics, or SLA, which relies on cytotoxic photo-initiators and high-energy UV irradiation, DIW provides a mild, bio-friendly environment^58,69. Because it operates via the room-temperature extrusion of aqueous or hydrogel-based pastes, high doses of heat-sensitive therapeutic agents, or even living cells, can be homogeneously mixed directly into the ink prior to printing without compromising their biological activity ⁵¹ . Future hybrid printers combining DIW (for drug delivery) and FDM (for mechanical frame) represent the next frontier in functional scaffold fabrication.

Optimal MBG compositions and topological lattice configurations

The biological efficacy of 3DPMBG is fundamentally dictated by its multiscale design: the nanoscale elemental composition and the macroscale topological lattice.

Stoichiometric optimization of dopants

Regarding chemical composition, a high silica network (typically 60–80 mol% SiO₂) is mandatory to maintain the ordered mesoporous channels and provide sufficient silanol (Si-OH) groups for rapid apatite nucleation ³² . For enhanced osteo-angiogenesis, trace element doping requires precise stoichiometric optimization. Efficacy must be perfectly balanced against potential heavy metal cytotoxicity or glass network disruption. The optimal doping windows and specific biological roles of common therapeutic ions (e.g., Sr, Cu, Co, Mg) are quantitatively summarized in Table 2^34,36.

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Table 2.

Quantitative optimization of elemental compositions in MBG scaffolds.

Elemental dopant	Optimal range (mol%)	Primary biological role & mechanism	Potential drawback at excessive doses
Silicon (Si)	60–80%	Acts as the primary network former; essential for maintaining mesoporous integrity and inducing apatite nucleation.	High doses (>80%) limit the incorporation of other bioactive network modifiers (Ca, P).
Calcium (Ca)	10–20%	Network modifier; accelerates degradation and provides raw material for carbonated hydroxyapatite (CHA) deposition.	Excessive Ca disrupts the ordered mesoporous channels and causes rapid bursts of local alkalinity.
Strontium (Sr)	1–5%	Promotes osteoblast Wnt/β-catenin signaling and inhibits osteoclastogenesis (dual-action).	>10% induces severe cytotoxicity and disrupts the silicate glass network due to its large ionic radius.
Copper (Cu)/Cobalt (Co)	1–2%	Simulates hypoxia by stabilizing HIF-1α; powerfully stimulates VEGF secretion and rapid angiogenesis.	Highly toxic heavy metals; >2–5% leads to severe oxidative stress, DNA damage, and cell apoptosis.
Magnesium (Mg)	5–10%	Enhances early mechanical strength; regulates the initial degradation rate of the glass network.	Competes with Ca2+ during apatite formation, potentially delaying the early onset of mineralization.

The optimal molar percentages (mol%) are synthesized from prevailing literature evaluating the balance between therapeutic efficacy and cytotoxicity. HIF-1α: hypoxia-inducible factor-1α.

Mathematical design of TPMS architectures

Beyond chemistry, the 3D topological configuration profoundly influences cellular mechanotransduction and fluid dynamics. While traditional orthogonal “woodpile” (grid) patterns are easily fabricated, advanced TPMS architectures are currently recognized as the optimal biomimetic geometries ³⁸ . TPMS structures possess zero mean curvature, perfectly mimicking the continuous, anisotropic architecture of native trabecular bone. Mathematically, these complex topologies are generated using nodal approximations based on trigonometric parametric equations ³⁸ . The three most prominently utilized TPMS structures in bone tissue engineering are the gyroid (G), diamond (D), and primitive (P) surfaces, defined by the following fundamental level-set equations (where x, y, and z are spatial coordinates and c defines the solid volume fraction):

Gyroid (G-surface): sin(x)cos(y)+sin(y)cos(z)+sin(z)cos(x)=c

Diamond (D-surface): sin(x)sin(y)sin(z)+sin(x)cos(y)cos(z)+cos(x)sin(y)cos(z)+cos(x)cos(y)sin(z)=c

Primitive (P-surface): cos(x)+cos(y)+cos(z)=c

Combined with an optimal macropore size of 300–500 μm, these mathematically derived TPMS lattices uniquely distribute localized stress, maximize fluid shear stress for OB mechanosensing, and provide ideal non-linear spatial geometry for rapid vascular network penetration ³⁰ . A comparative analysis of these topological architectures is detailed in Table 3.

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Table 3.

Comparative topological analysis of 3D printed lattice configurations.

Topological configuration	Curvature feature	Mechanical behavior	Permeability & fluid dynamics	Biological efficacy & in vivo outcomes
Orthogonal (Grid/Woodpile)	Sharp angles (90° joints)	High compressive strength but highly anisotropic; prone to stress-shielding and early fatigue failure at joints.	Straight channels lead to excessively high fluid velocity in the center, causing “wash-out” of cells.	Baseline: Poor cell attachment at sharp corners; limited surface area restricts rapid vascularization.
Gyroid (TPMS-G)	Zero mean curvature (continuous)	Isotropic mechanical properties; excellent energy absorption mimicking natural spongy bone.	Highly tortuous pathways provide optimal wall shear stress for mechanotransduction.	Optimal: High specific surface area maximizes cell adhesion and spreading; tortuosity accelerates complex capillary network formation.
Diamond (TPMS-D)	Zero mean curvature (tetrahedral)	Exhibits the highest elastic modulus among TPMS structures at equal porosities.	Moderate permeability; localized fluid vortices enhance nutrient mixing.	Excellent: superior load-bearing capability makes it ideal for load-bearing mandibular defect reconstruction.
Primitive (TPMS-P)	Zero mean curvature (cubic)	Lowest mechanical strength among the three TPMS structures.	Highest permeability; large open spherical pores with small interconnecting necks.	Moderate: Favors mass transport but limited specific surface area reduces the available sites for osteogenic cell proliferation.

TPMS, triply periodic minimal surface. Mechanical and fluid dynamic characteristics are highly dependent on the mathematically defined parameter c (solid volume fraction) and unit cell size.

Ionic orchestration of osteogenesis

The osteogenic capacity of 3DPMBG is driven by the “ionic dissolution products” released into the extracellular matrix^21,31. These ions are transported into cells via ion channels or endocytosis, where they trigger specific intracellular signaling pathways that govern the fate of mesenchymal stem cells (MSCs) and OBs^21,72,73.

Angiogenic coupling: The “Type H” vessel hypothesis

Bone regeneration is metabolically demanding; no bone forms without blood. In large CMF defects, the core of the scaffold often fails due to hypoxia-induced necrosis. 3DPMBG addresses this by promoting rapid vascularization, specifically targeting CD31ʰⁱEmcnʰⁱ (Type H) capillaries, which are known to couple angiogenesis with osteogenesis^92,93.

Osteoimmunomodulation: From “bio-inert” to “immuno-active”

Traditionally, biomaterials were designed to be “bio-inert,” aiming to avoid the host immune response. However, modern “Osteoimmunology” dictates that the immune system is the primary orchestrator of healing ¹⁰² . The implantation of 3DPMBG inevitably triggers a foreign body response (FBR), where macrophages play the central role ¹⁰³ .

In vivo degradation and ion-release kinetics

While in vitro simulated body fluid (SBF) assays provide baseline dissolution data, the in vivo degradation of 3DPMBG is a far more complex, dynamically driven process governed by physiological fluid perfusion, cellular activity, and local pH fluctuations. Unlike the static conditions in vitro, in vivo release kinetics exhibit a distinct biphasic profile.

Initially, the exceptionally high specific surface area and hierarchical porosity facilitate rapid fluid permeation, resulting in an initial burst release of therapeutic ions (e.g., Si, Ca, Sr) within the first 3–7 days ²¹ . This early burst is clinically advantageous, as it rapidly establishes the localized concentration gradient necessary for M2 macrophage polarization and initial angiogenesis.

Subsequently, the kinetics transition into a sustained, diffusion-controlled release phase lasting for weeks to months. This deceleration occurs because a dense layer of carbonated hydroxyapatite (CHA) rapidly precipitates on the scaffold surface in vivo, acting as a natural diffusion barrier that slows further dissolution of the underlying glass network^74,75. Furthermore, in vivo degradation is actively mediated by osteoclasts; the mildly acidic microenvironment (pH ~4.5) within the osteoclastic resorption lacunae dynamically accelerates localized glass dissolution^82,103. Crucially, the macroscopic degradation timeline of optimized 3DPMBG scaffolds (typically 3–6 months) aligns perfectly with the physiological rate of autologous bone creeping substitution ¹¹¹ . This synchronized degradation prevents premature mechanical collapse—a common pitfall of highly soluble biomaterials (e.g., borate glasses)—while avoiding the permanent stress-shielding and FBRs associated with non-resorbable crystalline bioceramics^10,112.

Synergistic effects and temporal control

In clinical reality, these mechanisms do not occur in isolation. The ideal 3DPMBG scaffold acts as a “temporal conductor” ¹¹³ .

Crucially, this temporal orchestration by 3DPMBG presents a significantly accelerated healing timeline relative to traditional bone substitutes. While highly crystalline bioceramics (e.g., conventional HA or inert titanium) often exhibit a delayed onset of osseointegration due to their slow degradation and lack of active ionic stimulation, the exceptionally high specific surface area of MBG ensures an immediate and proactive interaction with the host^9,10. This rapid initial ionic burst effectively shortens the prolonged inflammatory phase typically seen with inert materials and fast-tracks the onset of angiogenesis (shifting it to the critical first week)^34,113,114. Consequently, 3DPMBG shifts the entire regenerative timeline forward, transitioning the material from a passive space-maintainer to an active catalyst for early-stage healing, as comparatively summarized in Table 4.

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Table 4.

Comparative timeline of in vivo healing events: 3DPMBG scaffolds vs. traditional bone substitutes.

Healing phase (estimated timeline)	3DPMBG scaffolds (active regeneration)	Conventional bioceramics (e.g., highly crystalline HA)	Inert implants (e.g., titanium, polyether ether ketone [PEEK])
Phase 1: Inflammation & immunomodulation (days 1–7)	Accelerated M2 switch: Rapid Si/Ca burst effectively suppresses prolonged M1 phase; early anti-inflammatory niche.	Prolonged M1 phase: Slow degradation leads to delayed debris clearance and extended inflammation.	Foreign body response: Chronic FBR risk; initiation of fibrous encapsulation.
Phase 2: Angiogenesis & vascularization (weeks 1–4)	Early onset: Hypoxia mimicry (e.g., Cu, Co) triggers rapid Type H vessel ingrowth into macropores.	Delayed onset: Relies purely on passive host blood vessel invasion; poor core vascularization.	Avascular interface: Lack of internal porosity; vessels can only bypass the implant surface.
Phase 3: Osteogenesis & remodeling (months 1–6)	Rapid creeping substitution: High BIC (>50%); active Wnt/β-catenin stimulation matches scaffold degradation rate.	Slow appositional growth: Material remains unresorbed for years; acts only as a passive osteoconductive scaffold.	Mechanical interlock only: No material resorption; permanent stress-shielding risk.

HA, hydroxyapatite; FBR, foreign body response; BIC, bone-implant contact. The comparative biological responses and estimated timelines are conceptually synthesized from the in vivo and in vitro literature discussed in the Synergistic effects and temporal control section.

Hierarchical preclinical evaluation: From rodents to large animals

The validation of CMF bone substitutes typically follows a “pyramid” strategy, moving from proof-of-concept in small animals to translational verification in large animals that closely mimic human physiology^112,118.

The regulatory landscape: The paradox of personalization

Despite promising preclinical data, the commercialization of 3DPMBG faces a unique regulatory hurdle: the conflict between “mass standardization” (the basis of traditional regulations) and “mass customization” (the essence of 3D printing) ¹³¹ .

Technical and economic barriers to commercialization

Beyond regulations, practical hurdles impede the routine clinical use of 3DPMBG.

Current clinical status and the drive for commercial translation

While 3D-printed titanium and polyether ether ketone (PEEK) implants are routine in CMF surgery, 3D-printed bioactive ceramics are in their infancy. Most human applications of 3DPMBG are currently restricted to pilot clinical trials or investigator-initiated studies (IITs) ³⁷ . Preclinical evidence indicates that 3D-printed borate and silicate BG scaffolds facilitate bone regeneration in critical-sized craniofacial defects. Their tailored structural designs notably enhance osteogenesis and vascularization. Specifically, scaffolds fabricated by SLS show improved regenerative capacity, especially when combined with osteoinductive factors ¹²⁷ . In addition, 3D-printed borosilicate scaffolds incorporating MBG coatings demonstrate significantly boosted OB activity and bone formation in femoral defects, highlighting the critical role of mesoporous architectures ¹⁴⁴ .

However, transitioning these encouraging early outcomes into widespread clinical adoption requires navigating a complex commercial landscape. Currently, the global CMF and periodontal bone graft market is heavily dominated by particulate xenografts (e.g., Geistlich Bio-Oss®) and traditional BG putties (e.g., NovaBone®). While clinically ubiquitous, these commercial mainstays possess inherent functional ceilings. Particulate xenografts often exhibit extremely slow resorption rates, remaining encapsulated in vivo without integrating into the natural bone remodeling cycle^15,17. Conversely, while commercial 45S5 BG putties offer excellent osteostimulation, they completely lack the macroscopic mechanical integrity required for the structural reconstruction of complex, load-bearing maxillofacial defects (e.g., severe alveolar ridge atrophy or segmental mandibular resections)^18,22.

This technological gap highlights the immense translational value of 3DPMBG, positioning it not merely as another bone filler, but as a next-generation “patient-specific regenerative device.” Realizing this potential requires a robust “industry-university-research-clinical” translational ecosystem. There is a particularly urgent strategic demand to industrialize high-end oral medical devices to break the monopoly of traditional imported products and address the massive epidemiological burden of CMF defects. Transitioning MBG from an academic research project to a standardized, commercially viable medical device hinges on overcoming scalable manufacturing bottlenecks, establishing non-destructive QC, and conducting large-scale randomized controlled trials (RCTs) benchmarking 3DPMBG against autogenous bone grafts^131,134,145. Successfully navigating this commercialization pathway will not only disrupt the current dental biomaterials market but also provide a superior, cost-effective regenerative solution tailored to the specific anatomical and biological needs of the patient.

Intrinsic limitations and translational pitfalls of MBG

While 3DPMBG presents immense regenerative potential, explicitly acknowledging its intrinsic limitations and potential clinical pitfalls is crucial for a realistic translational roadmap. First, despite the structural optimization provided by AM, the inherent brittleness and low fracture toughness of pure silicate glasses remain a significant pitfall. This fundamentally limits their standalone application in major load-bearing CMF sites (e.g., segmental mandibular resections). To mitigate this mechanical fragility, current translational research is increasingly pivoting toward composite formulations, incorporating secondary biomaterial phases such as HA, tricalcium phosphate (β-TCP), or medical-grade flexible polymers like PCL. These composite strategies effectively deflect crack propagation and significantly enhance the macroscopic fracture toughness and load-bearing capacity of the scaffolds^26,66,67,141.

Second, the primary advantage of MBG—its extraordinarily high surface area and rapid ion release—can paradoxically become a toxicological pitfall. If the formulation is not perfectly optimized, an uncontrolled “burst release” of therapeutic heavy metal dopants (e.g., Cu²⁺, Co²⁺, Ag⁺) can easily exceed the local physiological tolerance, inducing severe cytotoxicity, oxidative stress (ROS generation), and adverse FBRs^21,35,114. Finally, there is a critical risk of a “degradation-regeneration mismatch.” In highly vascularized defect sites, if the mesoporous scaffold degrades and loses its structural integrity faster than the host’s osteogenic rate, premature mechanical collapse of the regenerated site may occur^10,112. Future translational efforts must specifically address these pitfalls through precise multiscale composite engineering before routine clinical adoption can be achieved.

Concluding remarks: An evolution in regeneration

The emergence of 3DPMBG represents a highly promising trend in CMF surgery. Although largely still in the preclinical and early translational stages, we are witnessing an encouraging evolution from the era of “passive substitution”—where inert blocks merely filled voids—to the era of “active regeneration,” where smart, hierarchical scaffolds possess the potential to actively instruct the host body to heal.

By synergizing the intrinsic bioactivity of sol-gel glasses with the geometric freedom of AM, 3DPMBG overcomes the historical dichotomy between biological function and mechanical competence. The evidence synthesized in this review confirms that through precise ionic doping (Sr²⁺, Cu²⁺, Ag⁺) and topological optimization (TPMS structures), these scaffolds can successfully orchestrate the “bone-vascular-immune” triad. However, the journey from a promising prototype to a standard-of-care medical device is fraught with challenges in standardization, stability, and regulatory compliance. As we stand on the precipice of clinical translation, the field must evolve beyond “trial-and-error” fabrication toward intelligent, dynamic, and predictive regenerative medicine.

Future horizons: The next generation of CMF scaffolds

To transcend current limitations, future research must converge materials science with digital intelligence and developmental biology. We propose three transformative directions that will define the next decade of research, offering a conceptual illustration of next-generation paradigms (Figure 5).

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Figure 5.

Conceptual illustration of next-generation paradigms for CMF regeneration. (a) AI-driven inverse design: unlike traditional forward engineering, this approach utilizes machine learning algorithms to reverse engineer scaffold topology based on specific clinical requirements (e.g., target elastic modulus, porosity, and degradation rate), enabling the creation of functionally graded structures that precisely match the host bone mechanics. (b) Four-dimensional printing (shape morphing): by introducing “time” as the fourth dimension, stimuli-responsive smart scaffolds can alter their geometry (e.g., expansion or unfolding) in response to physiological triggers (e.g., body temperature or pH). This capability facilitates minimally invasive implantation and allows the scaffold to dynamically adapt to skeletal growth or defect geometry after the surgery.

Final thought: Moving toward active regeneration

The development of 3DPMBG reflects an ongoing shift in biomaterial design for CMF reconstruction, moving away from passive space-maintaining grafts toward more biologically instructive microenvironments. By integrating the intrinsic bioactivity of mesoporous glass with the precise geometric control of AM, and anticipating future integrations with AI and 4D printing, the field is moving closer to bridging the gap between materials science and clinical application.

Ultimately, the successful translation of 3DPMBG relies on utilizing these scaffolds as active modulators of the local bone-vascular-immune axis. Overcoming the remaining translational barriers—specifically, establishing standardized regulatory frameworks for personalized devices and validating non-destructive QC methods—will require sustained interdisciplinary efforts among material scientists, bioengineers, clinicians, and regulatory agencies. Through such collaboration, 3DPMBG holds the potential to advance the clinical objective of CMF surgery from mere structural repair to restitutio in integrum—achieving comprehensive anatomical and functional regeneration.