Bone powder-laden hydrogel scaffolds in bone tissue engineering

Abstract

Bone powder-laden hydrogel scaffold is emerging as a promising bone graft material to offer solutions for bone defect repair in bone tissue engineering. Numerous hydrogel composite scaffolds loaded with xenobiotic bone powder or alloplast bone powder have been developed for preclinical experiments. In vitro experiments conducted on osteogenesis-related cells and in vivo studies using bone defect animal models have demonstrated that bone powder-laden hydrogel scaffolds exhibit favorable physicochemical properties, enhance the osteogenic behavior of osteogenesis-related cells, and improve the quality and efficiency of bone defect repair in animals. Bone powder-laden hydrogel scaffold can maximize the performance of individual components. However, this material has several limitations and has not yet been approved for clinical trials. Therefore, recent research has explored related products with superior properties, enhancing the mechanical, chemical, and biological characteristics of bone powder-laden hydrogel scaffolds, and proposed various strategies for improvement. This review summarizes the preparation procedures, therapeutic applications and possible improvements of various types of bone powder-laden hydrogel scaffolds.

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1 Introduction

Bone defects are a common challenge faced in the clinical setting. They are caused by various skeletal diseases, such as trauma, bone tumors, infections, congenital bone defects, and other diseases.¹ If bone defects are not promptly and effectively addressed due to factors such as systemic diseases, long-term use of anti-inflammatory drugs, related soft tissue injuries, or inappropriate treatment methods, they can lead to malunions and nonunions.² Large bone defects are unable to heal spontaneously and require the implantation of bone substitutes based on biomaterials to achieve bone tissue regeneration.³ Any biological or synthetic biomaterial used for implantation into the human body to restore bone health, maintain bone structure, or fill bone loss is considered a bone substitute. Over the past few decades, numerous experiments have been conducted to develop ideal bone substitute materials.

For the morphology of the bone substitute materials, Wang et al. pointed out that, in a rat model of a radial defect, particulate bone graft may accelerate bone defect healing compared with larger bone grafts.⁴ Thus, researchers often grind bone substitute materials into powder or nanoparticles and transport them to the transplant site. Most commercial bone graft materials also exist in powder form to enable complete filling of the defects in the bone. However, due to its loose morphology, bone powder is difficult to use in mechanically supportive areas.⁵ To overcome this challenge, biomimetic scaffolds have emerged as a promising strategy for promoting neovascularization and facilitating osteogenic ingrowth in bone regeneration.⁶ Among various scaffold designs, hydrogel-based systems have demonstrated particular potential. These scaffolds demonstrated significant osteoconductive potential, effectively promoting cellular adhesion, osteogenic differentiation, and subsequent matrix mineralization while establishing seamless osseointegration with host tissues. This multifaceted functionality culminated in the synergistic stimulation of coupled osteogenesis–angiogenesis processes essential for bridging critical-sized osseous defects.⁷ In recent studies, numerous researchers have combined bone powder with synthetic hydrogel scaffolds, achieving favorable outcomes in the fields of tissue engineering and bone regeneration (Table 1). Bone powder-laden hydrogel scaffold can mimic autogenous bone and repair the bone defect, allowing the bone powder and hydrogel components to each achieve their optimal effectiveness.

Table 1 Representative bone powder-laden hydrogel scaffolds and applications

Composite scaffold	Bone powder		Hydrogel	Application	Reference
Abbreviations: DCB, decellularized bone powder; GelMA, gelatin methacryloyl; DBM, demineralized bone matrix; ADA, alginate dialdehyde; PEO, poly (ethylene oxide); PLGA, poly(lactic-co-glycolic acid); VEGF, vascular endothelial growth factor; GG, gellan gum; DBP, demineralized bone powder; sEVs, small extracellular vesicles; PEG, poly ethylene glycol; hDBM, DBM hydrogel; PP, pearl powder; PRF, platelet-rich fibrin; NP, nacre powder; SA, sodium alginate; NPP, nano-pearl powder; C–HA, chitosan–hyaluronic acid; DFO@PCL NPs, deferoxamine@poly(ε-caprolactone) nanoparticles; MnCO, manganese carbonyl; HA, hydroxyapatite; BC, bioactive ceramic; PPG, polyurethane, polyacrylamide, and gelatin bioink; ECM, extracellular matrix; BCP, biphasic calcium phosphate; Co, collagen; MSC, mesenchymal stem cell; PRP, platelet-rich plasma; Alg, alginate; PVA, poly(vinyl alcohol); TA, tannic acid.
DCB–GelMA hybrid scaffold	Animal-derived bone powder	DCB	GelMA	Rats cranial deficiency	Gao (2021)^8
GG–DBP hydrogel scaffold		DBP	GG	Rats cranial deficiency	Cho (2021)^9
Hydrogel/DBM powder/BMSC composite scaffold		DBM	ADA	Rabbits cranial deficiency	Li (2021)^10
hDBM/DBM/gTCP hydrogel		DBM/gTCP	hDBM	Rabbits femoral deficiency	Kang (2022)^11
PEO–GelMA/DBM/(PLGA/VEGF) microspheres hydrogel		DBM	PEO–GelMA	Rabbits cranial deficiency	Wu (2022)^12
PEG/hyaluronic acid–Bio-Oss® composite scaffold		Bio-Oss®	PEG-hyaluronic acid	Rats cranial deficiency	Zheng (2024)^13
NPP/C–HA porous composite scaffold	Animal-derived bonelike powder	NPP	C–HA	—	Li (2020)^14
VEGF-laden PP hybrid fish GelMA hydrogel		PP	Fish GelMA	Rats cranial deficiency	Yang (2023)^15
PRF-laden NP/SA scaffolds		NP	SA	Rabbits cranial deficiency	Liu (2024)^16
The injectable ECM hydrogels loaded with BCP powder	Alloplast bone powder	BCP powder	ECM	Rabbits femoral deficiency	Ventura (2020)^17
Agarose–gelatine–HA–minocycline nanocomposite		HA	Agarose–gelatine	Rabbits cranial deficiency	Zhang (2021)^18
HA/Co hydrogel seeded with MSC and PRP		HA	Co	Rabbits radial deficiency	Bakhtiarimoghadam (2021)^19
BC/PPG composite scaffolds		BC	PPG	—	Wang (2022)^20
DFO@PCL NPs/MnCO nanosheets/GelMA/polylactide/HA matrix biomimetically hierarchical scaffold		HA	GelMA	Rats femoral deficiency	Zhang (2022)^21
GelMA/HA porous composite scaffolds		HA	GelMA	Rabbits cranial deficiency	Song (2022)^22
Mineralized GelMA/sodium Alg–HA hybrid hydrogel		HA	GelMA/Alg hydrogel	—	Miao (2022)^23
PVA/HA/TA/Co hydrogel		HA	PVA	Rats femoral deficiency	Xiang (2024)^24

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Given the rapid advancement of the technology involving bone powder-laden hydrogel composite scaffolds, there is an urgent need to summarize this highly promising research endeavor. This review aims to provide an overview of the applications and latest developments of bone powder-laden hydrogel composite scaffolds, summarizing their preparation methods, characterization, and advantages. Additionally, it further discusses the limitations of this technology and presents future prospects for material improvements and technological integrations to promote bone regeneration.

2 The fabrication of bone powder-laden hydrogel

2.1 The fabrication of bone powder

Currently, available bone substitute materials can be categorized based on their sources: the patients themselves (autogenous grafts), a different donor from the same species (allogeneic grafts), donors from a different species (xenogeneic grafts), and synthetically created materials (alloplastics).²⁵ Among these, autogenous grafts are widely accepted as the gold standard due to their excellent osteogenic, osteoinductive, and osteoconductive properties. However, their clinical use is subject to numerous limitations, such as the limited quantity of autologous bone available and potential complications at the donor site, including bleeding, infection, chronic pain, functional impairment, and donor site fracture.²⁶ Allografts avoid donor site morbidity but entail risks of immune rejection and bacterial or viral contamination.¹ Both allografts and autografts have their respective advantages and disadvantages, whereas xenografts and alloplast bone powder offer an alternative solution. They have the ability to promote bone regeneration and often manifest osteoinductive and osteoconductive properties.²⁵ This review mainly discusses the preparation and application of xenobiotic bone powder (animal-derived bone powder and animal-derived bonelike powder) and alloplast bone powder.²⁷

2.1.1 Animal-derived bone powder. Animal-derived bone powder is made from animal bone, which is similar to human bone, directly by various methods; we call this biological bone powder. Animal-derived bone powder has been widely used as a biomaterial to promote osteogenic differentiation. It is composed of organic factors such as collagen, proteoglycans, residual calcium, and bone morphogenetic proteins.²⁸ Due to the similarity in bone structure, the sources of bone powder for experimenters predominantly include mammals such as pigs, cows, deer, rabbits, and horses (very few also come from poultry). The materials are histologically processed and ground to prepare osteogenic particles. Taking the commonly used pig bones as an example, the fresh bone tissue is first shattered and then washed to remove excess bone marrow. The cleansed bones are subsequently frozen using liquid nitrogen and ground into bone powder with particles smaller than 4 mm³.¹¹ This approach replicates the strength, morphology, porosity, bioactivity, and load-bearing capacity of living bone by retaining the bone composition and microstructure.²⁹

In contrast to the immunogenic protein components present in vascularized organs such as the kidneys, heart, and liver, the bone matrix contains scarce immunogenic protein components.³⁰ Freeze-drying, demineralized freeze-drying, or gamma irradiation can be employed to further reduce the antigenicity of bone grafts. The processing of freeze-dried bone (FDB) encompasses several steps, commencing with donor screening, followed by soft tissue removal, size reduction, decontamination, antimicrobial treatment, freeze-drying, dehydration, secondary size reduction, and terminal sterilization.³¹ These procedures not only mitigate the risk of antigenic immune reactions but also enable the long-term preservation of bone powder. Consequently, FDB demonstrates superior performance in terms of dimensional stability, new bone formation, and cost-effectiveness compared with untreated allogeneic materials.³² Additionally, the removal of water disrupts the lipid envelope, thereby inactivating enveloped viruses and further reducing the risk of disease transmission. Thus, compared with fresh-frozen bone, FDB is easier to maintain, exhibits lower allogeneic immunogenicity, and presents a reduced risk of infection.³³ Clinical evaluation by Morato et al. showed that patients receiving FDB grafts exhibited no foreign body-type reactions for up to 120 days post-surgery, confirming its low immunogenicity.³² Unlike FDB, demineralized freeze-dried bone (DFDB) is processed using varying concentrations of hydrochloric acid for different durations. The demineralization process in DFDB enhances the accessibility and release of various growth factors, including bone morphogenetic protein (BMP)-2, 4, and 7, thereby facilitating rapid revascularization and hard tissue growth at the bone defect site.³⁴ This promotes regeneration, rendering DFDB more osteogenic than FDB. Histological analysis by Wood et al. demonstrated significantly greater bone formation in human subjects 19 weeks after implantation of DFDB compared with FDB.³⁵ Gamma radiation has been used as a terminal sterilization technique to prevent disease transmission to recipients. Sterilization complements other steps in tissue banking, including stringent donor screening, aseptic procurement, aseptic handling during processing, and preservation methods, to produce safe allografts.³⁶ It is common practice to employ low-dose gamma irradiation to reduce the antigenicity of bone grafts while minimizing adverse effects on their bioactivity and mechanical strength.³⁷

Most researchers employ demineralization or decellularization techniques to process animal bone tissue into demineralized bone matrix (DBM) or decellularized bone powder (DCB). In the preparation of demineralized bone powder (DBP), the demineralization process results in the loss of some calcium and mineralized components, but retains collagen and non-collagen proteins including cytokines and growth factors.³⁸ Compared with the direct use of native animal bone tissue, demineralization not only effectively removes potential immunogenic components (such as α-Gal antigens) but also addresses the issues of slow degradation of natural minerals and their poor integration with newly formed bone. Subsequent controlled remineralization via the incorporation of synthetic calcium phosphate (e.g., β-Tricalcium Phosphatev (β-TCP)) enables precise regulation of degradation rates and ion release profiles, thereby better matching the requirements of bone regeneration processes.^11,38,39 The decellularization process, on the other hand, results in DCB with a high content of natural growth factors and preserves a fibrous microstructure similar to the collagen bundles in the extracellular matrix (ECM), thereby achieving higher bone regeneration efficiency.^8,40 DBM and DCB are simple, with low antigenicity and ethical requirements, and retain bone conductivity and induction, and are widely used in bone regeneration studies. They are often combined with natural polymers such as fibrin, collagen and gelatin as mouldable bone grafts for dental and orthopedic treatment.²⁸

Demineralization of bone tissue is commonly achieved through acid treatment. At room temperature, the ground bone matrix is continuously stirred in HCl for 24 hours to remove minerals. The extent of demineralization is confirmed by scanning electron microscopy (SEM) to observe structural changes, by deoxyribonucleic acid (DNA) quantification (showing a reduction to 3.43 ng mg⁻¹ in DBM, p < 0.005), and by gel electrophoresis, ensuring effective removal of mineral and cellular constituents.^11,41 After rinsing, a 1 : 1 mixture of methanol and chloroform is used to eliminate lipids from the DBM. The key step in decellularization of DBM involves trypsin treatment: after rinsing, trypsin and ethylene diamine tetraacetic acid (EDTA) solution are added, and the mixture is incubated at 37 °C. After 24 hours, excess enzymatic proteins are removed through methods such as rinsing with sodium dodecyl sulfate (SDS), followed by freeze-drying for storage.^8,11 Although prolonged chemical processing may partially degrade environmentally sensitive growth factors such as bone morphogenetic protein-2 (BMP-2) and TGF-β, enzyme-linked immunosorbent assay (ELISA) results show that DBM retains bioactive BMP-2 (346.7 pg g⁻¹), TGF-β1 (152.2 pg g⁻¹), and vascular endothelial growth factor (VEGF) (57.5 pg g⁻¹), which continue to promote the expression of osteogenesis-related genes, indicating that its biological activity is preserved.^8,11 To further reduce the antigenicity of DCB, Gao et al. conducted experiments using pig bones with the key antigen α-1,3-galactose (α-1,3-gal) knocked out, which is responsible for xenogeneic rejection.⁸ Numerous studies have demonstrated that α-1,3-gal is crucial in triggering hyperacute or chronic immune rejection reactions, leading to delayed tissue repair and regeneration. Donor transplants deficient in α-1,3-gal can survive longer in the recipient's body, opening up new avenues and improvements for successful xenotransplantation.³⁹

Notably, the preparation method significantly influences the biological performance of DBM or DCB, with physico-chemical combined approaches demonstrating superior bioactivity preservation. For instance, in the study by Leng et al., porcine DBM prepared using ultrasonication combined with chemical reagents (methanol/chloroform buffer supplemented with 0.5% dispase and 1% Triton-X100) not only achieved complete cell removal (as confirmed by DAPI and HE staining) but also exhibited mechanical properties comparable to native bone, with a maximum compression force of 295.11 N and an elastic modulus of 0.003 N mm⁻². Furthermore, when combined with platelet-rich plasma (PRP), this DBM significantly enhanced new bone formation in a rabbit critical-size radial defect model. Micro-CT analysis revealed nearly complete defect healing at 12 weeks, outperforming scaffolds prepared using solely mechanical or chemical methods.⁴²

In contrast, conventional mechanical methods, while partially reducing antigenicity, exhibit notable limitations in completely eliminating immunogenicity, which may trigger immune rejection in recipients. Experimental evidence shows that co-culture of human peripheral blood mononuclear cells (PBMCs) with mechanically prepared wild-type porcine bone extracts induced robust immune activation: PBMC proliferation was significantly enhanced, along with elevated levels of pro-inflammatory cytokines such as tumor necrosis factor-α (TNF-α) and interleukin-1β (IL-1β). Concurrently, CD4⁺ T cells were activated (the proportion of CD69⁺ cells increased from 8.7% in controls to 21%), secreting key rejection-related cytokines interferon-γ (IFN-γ) and interleukin-17 (IL-17). In comparison, bone tissue from α-1,3-galactosyltransferase knockout pigs markedly attenuated these immune responses.³⁰ This indicates that purely mechanical methods (e.g., supercritical freeze-drying), although better at preserving structural integrity, may retain immunogenic components such as the α-Gal epitope, while purely chemical approaches often lead to growth factor loss.

Therefore, combined physico-chemical treatments offer a distinct advantage in balancing mechanical integrity, bioactivity retention, and immunogenicity reduction.

Despite its numerous advantages, DBM may fail to function as an adequate scaffold due to the lack of calcium and phosphate, leading to tissue collapse after implantation and rapid resorption prior to new bone formation.¹¹ Consequently, mixing DBM with materials rich in calcium phosphate is crucial for the preparation of DBM composites. One of the most commonly utilized forms of calcium phosphate in synthetic bone substitutes is β-TCP, which exhibits excellent osteoinductivity, osteoconductivity, biocompatibility, and bioresorbability. Additionally, TCP possesses an appropriate degradation rate, facilitating the release of calcium and phosphate ions to induce bone formation.⁴³ Therefore, bone powder derivatives mixed with calcium phosphate are frequently employed as bone grafts in research. Compared with preserving natural minerals in bone, controlled remineralization using synthetic calcium phosphate (such as β-TCP) offers significant advantages. β-TCP can continuously release calcium and phosphate ions during degradation, promoting osteoblast expression of key genes such as runt-related transcription factor 2 (RUNX2), alkaline phosphatase (ALP), and osteopontin (OPN) by activating the ERK1/2 and PI3K/Akt signaling pathways. By precisely adjusting the crystal phase, porosity (20–500 μm), and particle size of β-TCP, controllable degradation and ion release ranging from several days to months can be achieved, wherein pore sizes >100 μm facilitate vascular ingrowth, and surface roughness <100 nm influences cell adhesion. After combining 5 w/v% porous β-TCP particles with DBM, the bone volume/tissue volume (BV/TV) in animal models reached 45.2 ± 6.3%, nearly double that of DBM alone (23.0 ± 8.1%), effectively preventing early scaffold absorption and collapse. This strategy of precisely tuning material properties overcomes the limitations of natural minerals (fixed Ca/P ratio of 1.71) with their stable structure, providing a more predictable and highly osteogenic solution for bone defect repair.^11,39

In addition to using animal-derived bone processing and bone grafting, commercial animal bone meal is commonly used in research. At present, deproteinized natural bovine mineral bone powder (Bio-Oss®) extracted from cow bone is the most widely used bone substitute, and its osteogenic role has been recognized.⁴⁴ Bio-Oss® is a carbonate apatite crystal extracted from bovine bone. After special treatment, proteins and other organic components are removed, rendering its structure almost identical to human bone. Consequently, Bio-Oss® shares a similar structure with human cancellous bone and exhibits high porosity, allowing for cell adhesion and growth. It demonstrates excellent biocompatibility, good osteoconductivity, low biodegradability, and minimal tissue reaction. As such, it has been widely used as a primary bone graft material in dental implants and stands as one of the earliest and leading commercially available xenogeneic grafts, long regarded as the gold standard.^44–46 In recent years, other commercially available animal bone powder composites have gradually been developed as alternative materials for bone substitutes. For instance, Lim et al. compared A-Oss® and Bio-Oss®, both of which are low-temperature sintered inorganic bovine bone materials, and found no significant differences in bone regeneration efficiency, quality, or graft volume loss.⁴⁷ Another xenogeneic graft, InterOss®, based on bovine hydroxyapatite particles, has recently been applied in small preparations for implantation in the condyle of rats and in a study of critical-size defects in the alveolar bone of beagle dogs, demonstrating its bone regeneration efficacy.⁴⁸ Despite the continuous design and development of commercially available animal bone powder composites, often combined with synthetic scaffolds for research purposes, Bio-Oss® remains the most mainstream commercially available bone substitute at present.

2.1.2 Animal-derived bonelike powder. Animal-derived bonelike powder is a bone substitute material extracted from calcified corals and algae or mollusc shells. Given that the exoskeletons of such animals are not homologous to human bones, the calcium-containing compounds in their compositions also differ. The coral exoskeleton exhibits an intrinsic porous structure with pore diameters ranging from 150 to 500 μm, structurally analogous to human cancellous bone.⁴⁹ However, as a bioactive bone-filling ceramic, its degradation rate is excessively rapid for achieving an ideal repair effect.⁵⁰ To address this issue, a highly pure crystalline β-TCP bone graft substitute was developed using propagated hard coral and combined with hydroxyapatite. Evaluations have demonstrated that this composite represents a promising bone graft substitute.⁵¹ Among algae, only a limited number of species, such as L. corallioides, are capable of producing calcareous skeletons, containing key molecules similar to those in bone and possessing high mineral content.⁵² More commonly utilized is a commercial product derived from this alga, termed Aquamin®, which is applied as an oral calcium powder to promote bone formation.⁵³

In animal-derived bonelike powders, the most frequently employed is nacre powder (NP). NP is a composite material derived from the innermost layer of the animal shell, and it has become a promising substitute for bone. NP has a structure similar to native bone and consists of inorganic mineral phases and an organic matrix. The inorganic part of NP is mainly aragonite calcium carbonate, accounting for more than 95% of its composition.¹⁶ Besides, the organic matrix is mainly composed of proteins and polysaccharides. Although the content of organic matrix in the nacre layer is small, it plays a crucial role in controlling the nucleation and shape of CaCO₃ polymorphs, so the pearl powder also has certain osteogenic activity.⁵⁴ For this reason, nacre powder has been reported to have better bioactivity and bone induction compared with hydroxyapatite (HA). Pearl and nacre are similar in morphology and features, which may be due to that they are both formed in shell, so researchers can grind pearls or their products directly to make bonelike powder.¹⁴

2.1.3 Alloplast bone powder. Alloplast bone powder is an artificially synthesized, mostly inorganic substitute for bone transplantation, encompassing two important types: calcium phosphate and bioactive glasses.²⁷ These synthetic alternatives offer remarkable advantages in terms of availability and production, posing no risks of disease transmission or immune-related effects. Furthermore, they possess satisfactory physical properties and plastic resorption rates, allowing for large-scale production and use with minimal adverse reactions.⁵⁵ Bioactive glass is renowned for its capacity to form a direct chemical bond with bone tissue. A representative example is the standard bioactive glass system 45S5 which has a composition of 45-SiO₂, 24.5-CaO, 24.5-Na₂O, and 6-P₂O₅ (wt%).⁵⁶ These materials are biocompatible and can facilitate the nucleation of a bonelike apatite layer on their surfaces upon immersion in simulated body fluid. Through this process, bioactive glasses can build strong chemical bonds with bone tissue by forming a surface layer of hydroxyapatite, thereby enhancing osteointegration between implant material and living tissue.⁵⁷ Bioactive ceramics (BCs), classified as calcium phosphate materials, are the most common types of alloplast bone powder. These materials exhibit an intrinsic porosity and can be combined with other synthetic scaffolds, leaving room for the recruitment and proliferation of osteoblasts. Moreover, degradation products and released ions can enhance cellular activity and accelerate bone repair.⁵⁸

BC is composed of calcium (Ca), magnesium (Mg), silicon (Si) and other bioactive elements, and is an inorganic material with adjustable biological response and high biological activity. These compounds include hydroxyapatite, α-TCP and β-TCP or complex formation.⁵⁹ BC can release bioactive ions, participate in the proliferation, differentiation and growth of osteoblasts, and induce the formation of the osteogenic phenotype of bone marrow mesenchymal stem cells (BMSCs). BC has shown advantages in orthopaedic, craniofacial or dental, and bone tissue engineering applications. At present, the polyphase BC containing Ca, Mg and Si is generally prepared by sol–gel method. Specifically, Si[OC₂H₅]₄ (TEOS) is firstly mixed with deionized water and nitric acid, and the mixed aqueous solution of Ca(NO₃)₂·4H₂O and Mg(NO₃)₂·6H₂O would be added to the TEOS solution after complete hydrolysis of TEOS. Then NH₄HF₂ nucleating agent is added, and the mixture is stirred continuously until gel formation. The gel is aged overnight at room temperature, dried, ground, and sifted and calcined several times to obtain the BC powder.²⁰

Among calcium phosphate-based materials, the most commonly used is HA. As the major inorganic component of bone, it can provide a strong affinity for the host bone due to its similarity to the inorganic bone segments. HA, with excellent bone conductivity, biocompatibility and biointegration, and non-inflammatory and nontoxicity properties, as well as its bone conduction properties and bioactivity, has been widely used to make synthetic materials for bone grafting.^19,60 In addition, there are duplex calcium phosphate powder and other materials used for research and synthesis of bone replacement materials.

2.2 The fabrication of hydrogel

Natural bone tissue includes both inorganic and organic parts, with the organic part being the ECM crucial for bone remodeling. Bone tissue engineering aims to replace traditional bone transplants by creating scaffolds that mimic the ECM and using bone replacement materials like bone powder. An ideal bone scaffold needs open pores of the right size for nutrient flow and cell movement, similar mechanical properties to natural bone, and biocompatible, biodegradable materials to minimize inflammation during healing.⁹

To date, various artificial scaffolds have been created for bone defects. They are classified as naturally derived or synthetic materials. Inorganic materials like calcium phosphate have good biocompatibility and osteoconductivity but are fragile and flexibility-limited. Synthetic organic materials, like polylactic acid, can be easily synthesized with desired mechanical properties and biocompatibility but lack bioactivity. Naturally derived materials are promising for bone regeneration scaffolds due to their biocompatibility and low cytotoxicity.⁶¹ Hydrogels, a type of polymeric scaffold consisting of a three-dimensional polymer network formed by crosslinked hydrophilic chains, have demonstrated significant utility in fabricating space-filling scaffolds for new bone formation due to their similar characteristics to the ECM.⁶² Hydrogels incorporating hydrophilic moieties or domains with three-dimensional (3D) architectures can be fabricated from either synthetic or natural polymers.⁶³ Synthetic hydrogels, such as those fabricated from polyethylene glycol, polyvinyl alcohol, and poly(2-hydroxyethyl methacrylate), offer advantages including tunable degradation rates and adjustable mechanical strength. In contrast, naturally derived hydrogels, such as those produced from collagen, gelatin, and hyaluronic acid, possess inherent biological properties that support cell signaling, cell–matrix interactions, and biodegradability. However, they are often limited by poor mechanical strength and uncontrolled degradation behavior.⁶⁴ To address these limitations, combinations of synthetic and natural hydrogels have been explored to complement the properties of each type.⁶⁵

Gelatin methacryloyl (GelMA), a hydrogel based on modified gelatin, is currently the most widely employed material among naturally derived hydrogels. GelMA is synthesized by modifying gelatin with methacrylic anhydride or methacrylate groups to introduce unsaturated bonds. It can photocrosslink in the presence of a photoinitiator, forming 3D structures under ultraviolet (UV) light with controllable mechanical properties.⁶¹ These hydrogels can be fabricated into customizable geometries suitable for implantation and exhibit tunable mechanical properties that meet the requirements of scaffolds under diverse conditions. Furthermore, they display a range of beneficial characteristics, such as low immunogenicity, low cost, and the presence of naturally derived Arg-Gly-Asp (RGD) motifs that promote biointeractions between cells and the scaffold.⁴⁵ These attributes have rendered such hydrogels highly attractive for bone tissue regeneration and have spurred their widespread development for a variety of applications spanning from drug delivery to tissue engineering, establishing them as a promising candidate among diverse scaffold materials.⁶⁵ GelMA is usually prepared as follows. The gelatin is dissolved in phosphate-buffered saline (PBS), followed by the addition of methacrylic anhydride. The methacrylation reaction is terminated by adding excess PBS, and it is then dialyzed before being frozen.^45,66 To date, GelMA-based biomaterials have been widely studied for their physical and biochemical properties with applications in drug delivery and tissue engineering.

To better model the structure and function of the ECM, natural biological materials are often considered for incorporation into scaffolds. Some researchers suggested that GelMA scaffolds have high formability for filling bone defects but their mechanical strength may be still insufficient. Beck et al. extracted ECM from fresh cartilage and added it in GelMA scaffold construction. The incorporation of the ECM significantly improved the mechanical properties of the GelMA hydrogel, and ECM supported cell attachment and proliferation and promoted the formation of new bone.⁶⁷

In addition to GelMA, various other natural hydrogels have been widely utilized, such as alginate. Alginate is a naturally occurring polysaccharide composed of α-L-guluronate (G unit) and β-D-mannuronate (M unit) arranged as linear homopolymeric and heteropolymeric blocks, and it has been widely applied in drug, gene, and cell delivery systems due to its biocompatibility and low immunogenicity.⁶⁸ As a drug carrier, alginate can prolong the half-life of therapeutic agents and enhance their solubility. Furthermore, alginate undergoes facile degradation under oxidative conditions, such as in aqueous solutions containing sodium periodate or hydrogen peroxide, suggesting its potential as an oxidation-responsive drug delivery vehicle. Consequently, it is frequently incorporated into scaffolds designed for drug release.⁶⁹ Li et al. developed an injectable self-healing hydrogel from alginate dialdehyde (ADA) and gelatin to overcome the limitation of conventional hydrogels, which require injection within a specific gelation period to avoid needle clogging or structural damage from high injection forces. After sodium alginate (SA) was prepared to form alginate solution, sodium periodate was added, and pure alginate dialdehyde (ADA) was obtained after filtration and dialysis.¹⁰ The ADA solution was dissolved in borax or PBS, and then added to gelatin solution. Periodate oxidized alginate and gelatin undergo self-crosslinking in the presence of borax, creating the hydrogels that demonstrated biocompatibility and biodegradability.⁶⁸ This hydrogel also reduces the risk of foreign body reactions in addition to enabling better handling in clinics.

Along with several natural polymers used in hydrogels for tissue engineering approaches, such as collagen, hyaluronan, alginate, chitosan or fibrin, an additional wide range of synthetic hydrogels have shown suitable physical and chemical properties for regenerative medicine applications; these materials include poly ethylene glycol (PEG), poly(vinyl alcohol) (PVA), poly(propylene fumarate) (PPF), cellulose derivatives, Pluronic F-127 and polypeptides.⁷⁰ These synthetic hydrogels are often used in combination with natural polymers. A HA-PEG hydrogel was synthesized by Zheng et al. via the crosslinking of acrylated 4-arm PEG and thiolated ECM-mimetic component HA,¹³ which showed a 3D porous network structure and viscoelastic properties, exhibiting excellent biocompatibility.⁷¹ Similarly, in order to construct porous hydrogels to effectively enhance the exchange of oxygen and nutrients to improve cell survival, diffusion, migration, and proliferation, Wang et al. designed a novel hydrogel scaffold with a bubble-like porous structure from hydroxyethyl chitosan (HECS) and cellulose (CEL) by a combination of chemical crosslinking, particle-leaching using silicon dioxide particles as porogen and freeze-drying method.⁷² For alike considerations, Wu prepared the porous hydrogel scaffold based on an aqueous two-phase emulsion, consisting of GelMA and poly (ethylene oxide) (PEO) solution, which led to the formation of highly interconnected micropores upon subsequent photocrosslinking and leaching procedures.¹² The fabrication of the porous hydrogel reduces the disturbance of the nutrient exchange obstacle inside the hydrogel scaffolds and the inferior vascularization of the regenerated bone tissue.

We further found that the hydrogel system based on the aqueous two-phase emulsion template showed better overall performance than the particle-leaching and chemical crosslinking methods. This approach combined photocrosslinking with aqueous emulsion leaching, allowing efficient and solvent-free fabrication under mild conditions, and the porogen could be completely removed by soaking in pure water. As reported by Wu et al., in the PEO-GelMA system, adding 0.2% (w/v) LAP photoinitiator enabled full curing under 365 nm UV light at 20 mW cm⁻² within 60 seconds, followed by complete removal of PEO after soaking in pure water for 30 minutes. The entire process does not require any organic solvents, making it suitable for continuous production and Good Manufacturing Practice (GMP)-compliant sterile scale-up.¹² In contrast, the HECS/CEL method described by Wang et al. involves epichlorohydrin crosslinking, SiO₂ templating, and long acid–base treatments, which are complex and unsuitable for large-scale manufacturing.⁷² In addition, although the particle-leaching method can also produce macroporous structures, it depends on the packing of particles, often leading to uneven pore size distribution and high batch-to-batch variability. By comparison, the PEO-GelMA system reported by Wu et al. can stably form interconnected pores of 100–200 μm with high porosity and good consistency. At the same time, this system has clear advantages in biosafety.¹² Both PEO and GelMA are FDA-approved biocompatible materials, whereas epichlorohydrin is considered a potential carcinogen. Therefore, we believe that the photocrosslinking combined with emulsion method achieves the best balance among scalability, reproducibility, and clinical safety, and represents the most promising and clinically translatable strategy for fabricating porous hydrogel scaffolds for bone tissue engineering.

Currently, hydrogels have gained considerable attention in the field of bone regeneration as attractive tissue-engineered scaffolds. Different methods of preparations have been designed to address the existing limitations. Researchers should analyze the requirements of the experiment in the actual study, and choose the appropriate hydrogel composition and construction design.

2.3 The combination of bone powder and hydrogel

As mentioned above, the use of hydrogel as a carrier of bone powder has become one of the current research topics in bone transplantation surgery. As known, the physical and biological factors of a material can influence osteogenesis.⁷³ The stiffness, degradation rate, porosity, pore size, and overall biological performance of hydrogel scaffolds can be modulated by adjusting the prepolymer concentration and the cross-linking time to mimic the native tissue microenvironment and meet the necessary requirements and cues that promote bone formation.⁷³

When researchers prepared the hydrogel with bone powder, they mixed the treated bone powder such as DBM or DCB with the hydrogel. The mixed pre-gel liquid was put in a warm water bath until GelMA dissolved. The mixture obtained limited adhesion capacity and formed the bone powder-laden hydrogel material with a certain form.^8,66 The concentration ratio of the prepolymer can affect the stiffness and other biological properties of the hydrogel. When delivering DBM with thermogelling chitosan, Tian et al. found that the optimal DBM/hydrogel ratio was 2 : 1.⁷⁴ After photocuring in the mold, the bone powder-laden hydrogel is usually prepared into the specified size stent using a biopsy punch.⁹

For further combining the osteogenic ingredients, some groups also add growth factors and bone regeneration related cells to bone powder-laden hydrogel. Wu et al. designed DBM particles, and poly(lactic-co-glycolic acid) (PLGA)/VEGF microspheres to construct the biomimetic osteogenic/angiogenic microenvironments in porous hydrogel scaffolds.¹² Li et al. adopted a dual delivery of DBM powder and hypoxia-pretreated BMSCs, the latter expressing osteocalcin (OCN) and VEGF to facilitate bone regeneration.¹⁰ These methods further model the biological microenvironment of living bone regeneration and promote the success of bone transplantation.

For alloplast bone powders, their characteristics differ significantly from xenobiotic bone powder. In previous studies, HA particles were physically mixed with collagen solutions, resulting in physical or chemical crosslinking of the solution to form composite hydrogels. However, the inorganic bioceramic phase cannot be uniformly dispersed within the hydrogel, and unevenly dispersed particles and aggregated particles in the organic matrix may lead to poor mechanical strength.⁷⁵

To address the challenges of particle aggregation and uneven dispersion in hydrogel composites, researchers have developed a variety of strategies to improve the distribution and integration of bone powder within the gel matrix: (1) surface chemical coupling to integrate particles with the gel matrix. An example is that vinyl/silane modification of HA and photointegration into GelMA produces a chemically integrated composite that in Tong et al. increased storage modulus ∼19.43-fold to 79.2 kPa and achieved ∼61.3% breaking-load strength and ∼73.1% BV/TV vs. native cranium in a rabbit model, demonstrating that surface modification improves both dispersion-driven mechanics and long-term bone regeneration;⁷⁶ (2) particle-size and additive optimization: embedding nano-HA together with bioactive nanosilicates (SN) into GelMA markedly increased cell spreading area, produced large fold-increases in osteogenic gene expression (ALP/RUNX2/OCN/OPN ∼25/35/16/13× at day7 and ∼44/75/36/23× at day14) and led to robust in vivo healing (≈65% defect fill by new bone at 8 weeks), showing the increase of uniformity, mechanics and osteogenesis;⁷⁷ (3) biomimetic coatings and processing: polydopamine (PDA)-assisted HA coatings dramatically improved particle surface wettability and coating robustness and yielded more new bone by micro-CT at 8 weeks in rabbit implants, indicating PDA or similar interfacial layers are effective dispersion strategies during fabrication and in vivo service.⁷⁸

Building upon these dispersion strategies, in situ mineralization provides a more straightforward and biomimetic approach that directly addresses particle aggregation while enhancing the mechanical and biological integration of the hydrogel composite. In situ mineralization is a straightforward and effective technique capable of mimicking natural mineralization, leading to the uniform distribution of nanoparticles within a hydrogel matrix and the formation of uniform hydrogel composites of HA with excellent interactions between the organic and inorganic components. Furthermore, the in situ synthesis of HA particles in the presence of collagen fibers represents a biomimetic approach that can result in the formation of Co/HA composites resembling bone. HA precursors are added to a neutralized collagen solution with a pH of 8–9 and a temperature below 10 °C for 2 hours to synthesize HA in situ. Subsequently, the collagen/HA solution is incubated at 23–25 °C for either 10 minutes or 16–20 hours to form a Co/HA hydrogel (Fig. 1).⁶⁰

Fig. 1 Schematic illustration of the construction of bone powder-laden hydrogel.

Multiple studies report marked mechanical gains after in situ mineralization. For example, Shuai et al. discovered a poly-L-lactic acid (PLLA)/Combining graphene oxide (GO)-HA composite showing ∼53.7% and ∼98.8% increases in compressive strength and modulus, respectively, and observed that mineral crystals distributed uniformly along collagen/polymer fibrils, better mimicking the hierarchical structure of bone and enhancing long-term mechanical stability.⁷⁹ This is because the in situ mineralization process facilitates direct nucleation and growth within the hydrogel, thereby creating a seamless integration of the components at the molecular scale. This integration significantly enhances compressive strength, modulus, and homogeneity, while reducing sample-to-sample variation.⁸⁰ In contrast, scaffolds prepared by simple physical mixing tend to have a more heterogeneous particle distribution and weaker interfacial bonding. Consequently, they possess lower initial mechanical strength and modulus and are more prone to mechanical deterioration and local instability under long-term incubation conditions. As highlighted in several reviews and experimental studies, this is because such physically blended systems rely on secondary interactions (e.g., van der Waals forces or mechanical embedding), which provide inferior mechanical performance and higher sample-to-sample variability. For example, Charles et al. studied HA-coated PLLA fiber self-reinforced composites and found that while HA increases flexural modulus, the extent of enhancement strongly depends on HA loading and interface treatment; lower loadings or weak fiber–matrix bonding result in smaller gains, and samples show more variability in mechanical properties across batches.⁸¹ Regarding the comparative osteogenic potential, the two approaches also demonstrate clear differences. Bakhtiarimoghadam et al. reported that in situ mineralization within a collagen scaffold was significantly more effective in promoting osseointegration and osteogenic differentiation than simple physical mixing of the components.⁸² A similar conclusion was reached by Liang et al., who developed a biomimetic hydrogel through the in situ mineralization of bonelike nano-hydroxyapatite within decellularized periosteum matrices. Their micro-CT quantification revealed a significantly increased bone volume fraction in the mineralized group over the non-mineralized control, corroborating their finding.⁸³ Collectively, these findings indicate that, compared with physically mixed bone powder, the in situ mineralization strategy leads to superior mechanical strength, homogeneity, and long-term stability. Beyond these properties, it also enhances osteogenic differentiation in vitro and in vivo, resulting in accelerated and denser bone regeneration.

Beyond mineralization strategies, functionalization of collagen scaffolds with bioactive molecules represents another promising approach. For instance, Song et al. developed a DNA-crosslinked collagen scaffold (DNA-Col) that not only enhances mechanical strength but also promotes bone regeneration through immunomodulatory effects. The DNA-Col scaffold recruits regulatory T cells (Tregs) via the TLR4-p38-PGC-1α pathway, creating an anti-inflammatory microenvironment conducive to osteogenesis.⁸⁴

3 The characteristics and advantages of bone powder-laden hydrogel

The biological and mechanical performance of fabricated bone scaffolds are typically dictated by the interplay of characteristics classified into three broad groups: (1) inherent material characteristics, (2) surface characteristics, and (3) scaffold or architectural characteristics.⁸⁵ In addition, cell experiments are often used to observe cell proliferation, expression and differentiation related to bone regeneration. Bone powder hydrogel scaffolds are usually characterized by the following aspects:⁸⁶

3.1 Surface roughness/topography

The surface roughness of scaffolds directly influences the attachment capability of host cells. A certain degree of roughness on the pore walls favors cell adhesion and migration, and it also directs the body to respond and modulates crosstalk among different cellular components. Surface topography can be observed via scanning electron microscope (SEM). Under SEM, GelMA hydrogels exhibit a uniform and highly interconnected network with numerous regular pores.²¹ Pores are generated when aqueous solvents freeze to form ice crystals, leading to phase separation between the solvent and the composite. This porous structure is crucial for ensuring cell proliferation in tissue engineering applications by preserving sufficient space and facilitating nutrient transport, promoting oxygen exchange, and enabling the release of bioactive additives.⁸⁷ Wang et al. incorporated BC powders with an average particle size of approximately 0.8 μm into a polyurethane/polyacrylamide/gelatin (PPG) hydrogel via 3D printing. The BC/PPG bioinks, containing 10%, 20%, and 25% BC by weight, were crosslinked using ultraviolet light (365 nm, 1.5 W cm⁻²) and 1-ethyl-(3-dimethyl aminopropyl)-carbon-diimide hydrochloride (EDC)/N-hydroxy succinimide (NHS) (200 mM each). After 7 days of immersion in simulated body fluid (SBF), all BC/PPG scaffolds induced the deposition of hydroxyapatite on their surfaces, as confirmed by X-ray Diffraction (XRD). Following 5 days of culture with extracts from the BC (20%)/PPG scaffold, BMSCs exhibited the highest cell viability, indicating that the incorporation of bioactive ceramic powders roughened the surface of the hydrogel scaffolds, which in turn helped regulate cell behavior and promote cell adhesion and proliferation.²⁰ The dispersion of bone powder within the hydrogel can also be observed under SEM. Song et al. discovered that as the solid content of hydroxyapatite increases, more hydroxyapatite particles are uniformly dispersed throughout the hydrogel.²² Notably, Qin et al. developed an electroconductive hydrogel (GMPA) that not only presents a favorable microporous structure under SEM but also responds to endogenous electric fields, effectively promoting neurite outgrowth and neuropeptide release, thereby synergistically enhancing osseointegration and osseoperception around dental implants.⁸⁸

3.2 Pore size and porosity

The ability of scaffolds to promote osteogenesis on their surfaces and within their pores depends on their porosity, which is primarily determined by pore size, volume, and interconnectivity. The pore sizes both inside and outside the material can be observed via SEM, and porosity tests can be conducted on scaffolds using the ethanol displacement method or by employing a porosimeter to obtain specific porosity values. Closely related to porosity is permeability, which represents the ability of liquids to flow through the scaffold and is crucial for the transport of nutrients or metabolites within bone tissue. Assessment of mass delivery of porous hydrogels performed by rhodamine dye penetration experiments by Wu et al. showed that larger dye-permeable areas were observed with increasing micropores mediated by porous structure and particle occupancy effect.¹² Permeability values were further analyzed using computational fluid dynamics (CFD) analysis, which can also be employed to evaluate permeability.²¹

Pore sizes of ≥300 μm are conducive to the formation of new bone and capillaries, as larger pore sizes provide a greater surface area for cell adhesion and proliferation. Materials with high porosity facilitate nutrient transport and promote tissue repair.²³ Liu et al. prepared NP/SA composite scaffolds with varying NP-to-SA ratios (0 : 1, 1 : 2, 1 : 1, 2 : 1, w/w) using 3D printing. The porosity was measured via the ethanol displacement method. Results showed that as the NP content increased, the porosity decreased significantly. SEM analysis revealed that the pore size of the scaffolds also decreased with increasing NP content, ranging from approximately 400–700 μm. This reduction in both porosity and pore size was attributed to the occupation of the internal hydrogel space by NP particles, which limits pore formation.¹⁶ Li et al. assessed the pore size and porosity of injectable ECM hydrogels loaded with biphasic calcium phosphate (BCP) powder through SEM and mercury intrusion method in their study with the Ventura team. The experiment used 30 mg mL⁻¹ of acellular porcine dermal ECM, with 0%, 10%, and 15% (w/v) BCP powder added to prepare the hydrogel samples. SEM images indicated that while the addition of BCP did not damage the porous connectivity, it effectively regulated the pore structure through physical filling, thereby affecting its mechanical properties and osteogenic induction potential.¹⁷ However, Li also noted that some researchers’ data are inconsistent with theirs, potentially due to differences in the powder added. Generally, scaffolds with a porosity range of 89% to 93% are favorable for cell migration and blood vessel formation.⁸⁹ Notably, in a recent study by Zhao et al., a biomimetic bilayer piezoelectric scaffold was designed with a pore size of 100–350 μm in the periosteum-like layer and 100–200 μm in the bone-like layer, which effectively supported cell infiltration, nutrient exchange, and neuro-vascularized bone regeneration under ultrasound stimulation.⁹⁰

3.3 Mechanical properties

The stent should have mechanical properties similar to the host bone in the defect site, and the mechanical properties of the material are also an important parameter when designing a bone tissue engineering stent. Harder matrices that closely match bone matrix properties are generally considered more suitable for effective bone regeneration.⁹¹ The mechanical properties of bone powder-laden hydrogels are often tested through compressive tests and tensile loading–unloading tests, yielding compression stress–strain curves and tensile stress–strain curves, respectively, from which the compression modulus and elastic modulus are calculated. Universal testing machines or dynamic mechanical analyzers are utilized for these measurements. Generally, the incorporation of bone powder into hydrogels enhances their compressive strength and toughness. Miao et al. demonstrated that the addition of HA endows hydrogels with greater stress, as the rigid HA layer partially dissipated deformation energy. Furthermore, among various hydrogels, HA can also compensate for the weakened strength of the hydrogels.^23,92 Li et al. evaluated the mechanical properties of nano-pearl powder (NPP)/chitosan–hyaluronic acid porous composite scaffolds using a universal material testing machine under compression. Cylindrical samples (25 mm in diameter, 10 mm in height) were compressed at a rate of 5 mm min⁻¹. Results indicated a significant enhancement in compressive strength with increasing NPP content from 0% to 25% (p < 0.05). Specifically, scaffolds containing 25% NPP exhibited the highest mechanical resistance, outperforming the NPP-free control group. This improvement was attributed to the homogeneous dispersion of NPP particles within the polymer matrix and their physical interaction with the chitosan–hyaluronic acid network, which collectively enhanced structural integrity and supported the scaffold's potential for load-bearing bone regeneration applications.¹⁴ Wu et al. demonstrated that incorporating 5% w/v DBM particles into 10% w/v GelMA hydrogels significantly enhanced the storage modulus from 2.1 kPa to 4.8 kPa, reduced the phase angle from 0.35 to 0.22, and shortened the gelation time from 8 s to 5 s, indicating improved network rigidity due to DBM. Compression tests further confirmed that the composite hydrogel achieved a compressive modulus of (125 ± 8) kPa, representing an approximately 84% increase compared with pure GelMA.¹² Overall, the increase in stent aperture indicates higher permeability and lower mechanical strength and should reach a compromise between permeability and mechanical properties.

3.4 Degree of gelation and the rheological properties

In the preparation of composite hydrogel scaffolds, the composite hydrogel needs to undergo gelation from a free-flowing liquid state (sol) to a non-flowing solid state (gel). The method for assessing turbidimetric gelation involves using a spectrophotometer to measure absorbance, thereby obtaining the time to half gelation (t_1/2) and the gelation rate (S).¹⁷ The degree of gelation from sol to gel in composite hydrogels can further be confirmed through the tube inversion test. Additionally, some researchers determine the degree of gelation by measuring viscosity, such as through the shear rate–viscosity test or directly using viscometers to obtain the viscosity and shear modulus of composite hydrogels.²² Using this method, Kang et al. systematically confirmed the self-gelling ability of DBM-laden hydrogels (hDBM) at 37 °C. The gelation process, characterized via tube inversion and turbidimetric analysis, showed a complete gelation time of 18.03 ± 0.76 minutes. Rheological evaluation revealed a distinct shear-thinning behavior, with viscosity decreasing from 1.25 Pa to 0.017 Pa as the shear rate increased, which is critical for injectability. Furthermore, SEM imaging confirmed that the hDBM possessed a fibrous, porous extracellular matrix structure. The incorporation of granular β-TCP (gTCP) with its interconnected multi-channel architecture provided a macroporous scaffold within the hydrogel. These specific data on gelation kinetics, rheology, and microstructure robustly support the material's advantages as an injectable bone graft.¹¹ The use of a rotary rheometer to assess the viscoelasticity of the hydrogel, namely the rheological properties, is also the method used by many research groups. During testing, an oscillatory amplitude scan at a constant frequency is employed, with strain increasing from 0.01% to 100%, to measure the storage modulus and loss modulus. The results indicate that hydrogels are mechanically responsive systems, with their rheological properties fluctuating in response to changes in stress or strain. When the storage modulus is greater than the loss modulus, all hydrogels exhibit elastic behavior, whereas when the reverse is true, all hydrogels exhibit liquid-like behavior. As oscillatory strain increases, storage modulus begins to decrease, and loss modulus begins to increase, indicating deformation of the hydrogel network. Beyond the transition point (where the storage modulus equals the loss modulus), the hydrogels transition from solid-like to liquid-like behavior.^24,93

3.5 The ability of swelling

The determination of swelling capacity involves assessing the amount of fluid that a material can absorb over a given period. The swelling capacity of bone powder-laden hydrogels is measured by immersing them in various liquids such as PBS, artificial saliva, SBF, Ringer's solution, and distilled water. High swelling ratios of materials often lead to scaffold collapse, causing the loss of the scaffold's original structure. Such high swelling can result in implant loosening and exert adverse pressure on bone tissue and the surrounding microenvironment.¹⁶ Furthermore, the swelling characteristics of materials can also affect cellular infiltration in vivo and the ability of the material to allow the diffusion of nutrients and waste within the hydrogel matrix.⁹⁴ The anti-swelling capacity of bone powder-laden hydrogels is beneficial for avoiding postoperative deformation and reducing pressure on surrounding tissues.⁹⁵ In the absence of bone powder, there are free spaces between hydrogel chains where fluids can easily penetrate. However, after loading with bone powder, these free spaces are filled, thereby limiting the possibility of fluid adsorption. Miao et al. systematically evaluated the swelling behavior of mineralized GelMA/SA hydrogels. Lyophilized samples were immersed in physiological saline at 37 °C and weighed at 1, 2, 3, 6, 12, and 24 hours to calculate the swelling ratio (S% = (M₁ − M₀)/M₀ × 100%). Results showed that the equilibrium swelling ratio of pure GelMA hydrogel was 919.08 ± 9.66%. With the addition of 1% alginate, the swelling ratio decreased to 817.31 ± 26.52%. After three days of sequential mineralization (alternating immersion in 166.7 mM CaCl₂ and 100 mM K₂HPO₄ solutions), the swelling ratio of GelMA/SA hydrogels further decreased significantly to approximately 400%. This reduction was attributed to the incorporation of HA, which occupies pore spaces within the polymer network, and the enhanced crosslinking density from alginate–Ca²⁺ interactions, collectively limiting water absorption capacity.²³ Similarly, Gao et al. systematically evaluated the swelling behavior of DCB incorporated into GelMA hydrogels. Cylindrical scaffolds were immersed in deionized water and PBS at 37 °C until equilibrium swelling was reached, and the swelling ratio was calculated as (W_t − W₄)/W₄ × 100%. Results indicated that as the DCB content increased from 0% to 50%, the swelling ratio in deionized water decreased significantly from approximately 600% to 200%, with a similar trend observed in PBS. This reduction was attributed to the physical occupancy of DCB particles within the hydrogel network pores and enhanced structural density through DCB–GelMA interactions, collectively limiting water absorption capacity.⁸

3.6 Phase and chemical structure

The testing of the hydrogel phase and microstructure involves utilizing XRD to detect the crystalline phase of the hydrogel and identifying the chemical composition of the hybrid hydrogel within a certain range through Fourier Transform Infrared Spectroscopy (FTIR). XRD is employed to assess the crystalline structure of the hydrogel composites and can also be aided by methods such as Cu-Kα radiation.¹⁸ On the X-ray diffraction pattern, different diffraction peaks can be observed, each associated with the crystal pattern of the respective components. For instance, the diffraction pattern of collagen displays a broad peak, indicating its amorphous structure, while the crystalline structure of HA resembles the inorganic component of bone.¹⁹

The FTIR method is frequently used to confirm the components and characteristic functional groups in mineralized hydrogels, applied to hydrogels synthesized under established optimized conditions.⁹⁶ In unmineralized hydrogels, the characteristic absorption peaks of GelMA are present in all hydrogels. These include a broad peak at 3301 cm⁻¹ (common signal for O–H and N–H stretching), 3068 cm⁻¹ (N–H), 2933 cm⁻¹ (saturated C–H stretching), 1631 cm⁻¹ (amide I, vibration of the C=O bonds), 1534 cm⁻¹ (amide II, N–H bending), and 1241 cm⁻¹ (amide III, C–N stretching, N–H bending, and glycine C–H₂).⁶² The addition of other substances, such as bone powder, shifts these absorption peaks, indicating the formation of bone powder grafts or other loaded molecules within the hydrogel.²³

Furthermore, elemental mapping analysis can also be used to detect the formation of loaded molecules within the hydrogel. For example, when HA is used as a bone powder graft, the uniform distribution of C, O, P, and Ca elements can confirm the successful coating of HA nanoparticles onto the scaffold.⁷⁷ Energy-Dispersive Spectroscopy (EDS) can also calculate the molar ratio of specific atoms, compared with standard values, to prove that no other impurities have been introduced during the formation of bone powder-laden hydrogels and surface modification.⁹⁷

3.7 Surface hydrophilicity

The hydrophilicity of scaffolds can be evaluated by measuring the contact angle at the solid–liquid interface. The contact angle is measured by directly adding deionized water to the surface of the scaffold under certain conditions.⁹⁸ Li et al. evaluated the surface hydrophilicity of NPP/chitosan–hyaluronic acid composite scaffolds using an optical tensiometer. A 5 μL droplet of deionized water was placed on the scaffold film surface, and the contact angle was measured after 5 seconds. Results showed that as the NPP content increased from 0% to 25%, the water contact angle significantly decreased from approximately 110° to 65° (p < 0.05), indicating enhanced hydrophilicity. This improvement was attributed to the increased surface roughness induced by NPP incorporation, which promotes water spreading and facilitates cell adhesion and migration.¹⁴ Researchers using NP as a bone powder graft also propose that NP, a natural composite material primarily sourced from marine mollusks, is renowned for its excellent hydrophilicity.⁹⁹ Consequently, an increase in NP content in NP-laden hydrogel scaffolds leads to enhanced hydrophilicity. The improvement in hydrophilicity can promote cell adhesion, migration, and body fluid infiltration.²¹

From the aforementioned scaffold characterization tests, we can elucidate the properties of the bone powder-laden hydrogel and summarize its advantages. Among various bone substitutes, bone powder exhibits excellent biocompatibility and osteoconductivity, and its morphology contributes to superior bone defect healing effects. However, when used alone, the pressing issues are its high brittleness and low toughness, leading to poor reliability and resistance to damage.¹⁰⁰ DBM powder promotes bone regeneration due to its osteoinductivity resulting from the inclusion of bone morphogenetic proteins in the matrix and it possessing the maximum surface area for interaction with target cells. However, the use of DBM powder alone remains challenging because of its unstable osteoinductivity due to inactivation of growth factors during the preparation process, lack of bone regeneration cells, and difficulty in handling.¹⁰ Various forms of bone substitutes, including isolated particles, powders, and bone blocks, exhibit poor handling properties and spatial maintenance characteristics when used alone. Mobile isolated particles with high stiffness can elicit local inflammatory responses, thereby hindering bone regeneration. Compared with bone powder alone, hydrogels possess a strong ability to aggregate and encapsulate isolated bone powder particles, adhering firmly to the host bone defect as a single unit after photocuring.

There are also problems for hydrogels that do not bind to bone powder and to similar specific representations that limit their use in bone tissue engineering. Hydrogels provide a suitable environment for cell growth due to their porosity and water affinity, and blend well with surrounding tissue and reduce the inflammatory response. Their plasticity, viscoelasticity and stretchability enable scaffolds to combine with other biomaterials, enhancing scaffolds’ mechanical and physical properties.¹³ Furthermore, raw materials for hydrogels are abundant and can be tailored for desired shapes, with controlled degradation, porosity, and release profiles by adjusting crosslinking. Additionally, hydrogels are absorbable and integrate well with tissues, avoiding surgical removal and reducing inflammation.¹⁰¹ These advantages allow the various hydrogels to be widely used in bone tissue engineering to carry various classes of molecules to bone defect sites.⁶⁰ However, hydrogels alone are limited by their rapid hydrolysis, susceptibility to enzymatic degradation, excessive swelling, and low mechanical strength, which may hinder their osteogenic capabilities. Incorporating inorganic materials into polymer hydrogels may be a suitable option to overcome these issues and enhance their bioperformance.¹⁰² The loading of bone powder not only exhibits its unique osteoconductive properties and bioactivity but also modifies the mechanical strength and other characteristics of hydrogels. For instance, the incorporation of HA enhances collagen-based hydrogels with excellent durability, bioactivity, collagen induction, elastin regeneration, and fibroblast activity stimulation.¹⁰³ Liu et al. also found that adding NPs to hydrogel scaffolds not only increases their stiffness but also modulates their degradation rate to align with the rate of bone regeneration.¹⁶

Furthermore, poor dispersion of bone powder materials in composites may lead to unstable mechanical properties and ineffective bone repair outcomes. Ghorbani et al. demonstrated that the surface modification of hydroxyapatite with silane coupling agent, which can react with inorganic materials, can improve the dispersion of hydroxyapatite in the composites.¹⁰⁴ Regarding the morphology of bone powder, nanoparticles exhibit greater potential to promote the mechanical and biological properties of scaffolds. Nanoparticles can serve as “adhesion-promoting” agents, significantly enhancing cell adhesion, proliferation, and other related functions. Moreover, nanoscale HA precipitates possess higher dispersion and structural stability, which will benefit the scaffolds by enhancing uniformity and stability, thereby exhibiting higher mechanical strength.¹⁰⁵

In summary, well-dispersed bone powder-laden hydrogel composites demonstrate enhanced mechanical properties and adjustable biological characteristics, promoting osteogenic differentiation (Fig. 2). Bone powder-laden hydrogel composites can maximize the performance of individual components.¹⁰⁶

Fig. 2 Schematic illustration of the characteristics of bone powder-loaded hydrogel. The bone powder-loaded hydrogel exhibits numerous distinctive features, demonstrating excellent properties.

3.8 Immunomodulatory and angiogenic effects

Beyond their excellent mechanical and physicochemical properties, bone powder-laden hydrogels play pivotal roles in immune regulation and angiogenesis, both of which are critical for in vivo bone regeneration. Incorporating DCB derived from α-1,3-galactosyltransferase knockout pigs into GelMA effectively reduces xenogeneic immune rejection. Removal of α-1,3-galactose antigens markedly attenuates hyperacute immune responses. In a rat calvarial defect model, DCB–GelMA scaffolds exhibited minimal inflammatory cell infiltration as shown by HE staining, while human MSCs (hMSCs) seeded on these scaffolds displayed robust upregulation of osteogenic genes such as RUNX2, bone sialoprotein (BSP), and OPN. These findings suggest that the composite not only provides excellent immunocompatibility but may also indirectly shape the local immune microenvironment by enhancing osteogenic differentiation, thereby promoting angiogenesis and bone repair.³⁰

Similarly, combining rat BMSC-derived small extracellular vesicles (rBMSC-sEVs) with a PEG/HA hydrogel and Bio-Oss® produced sEVs/PEG/HA-Bio scaffolds with low immunogenicity. After 12 weeks of implantation in a rat calvarial defect, HE staining revealed no obvious inflammatory cell infiltration across all groups, while immunohistochemical analyses demonstrated significantly higher CD34-positive cell counts and increased OCN and BMP-2 expression. These data indicated that the scaffold effectively promoted neovascularization and exhibited strong osteoinductive capacity, likely by modulating the local immune milieu.¹³

Further evidence supports the cooperative effects of bone powder-laden hydrogel composites on immune modulation and angiogenesis. For example, an injectable PEG/hydroxypropyl methyl cellulose (HPMC)-based hydrogel containing equine cortical bone powder and demineralized bone matrix promoted cartilage-like tissue formation and vascular invasion in a rabbit femoral defect model. Immunohistochemical staining confirmed type II collagen deposition and active marrow regeneration, suggesting that the degradation products of the hydrogel modulated local mechanical and immune signals, guiding mesenchymal stem cells toward chondrogenesis and enhancing endochondral ossification with concomitant angiogenesis.⁷⁰

Moreover, bone powder-laden hydrogels can facilitate vascularization by sustaining or releasing pro-angiogenic factors. Incorporating platelet-rich fibrin into a NP/SA scaffold significantly upregulated VEGF and type I collagen (Col I) expression at 4 and 8 weeks post-implantation. This VEGF release, together with early immune-cell-mediated inflammatory modulation, accelerated new vessel formation and bone regeneration, leading to increased bone volume fraction and bone mineral density.¹⁶ Likewise, injectable porcine dermis-derived decellularized ECM hydrogels containing BCP retained bioactive molecules such as VEGF and BMP-2, as confirmed by ELISA, thereby providing sustained angiogenic and osteogenic cues within the defect site.

Taken together, these findings demonstrate that bone powder-laden hydrogels not only supply osteoconductive matrices but also actively modulate host immunity and stimulate angiogenesis. Through mechanisms including xenoantigen removal, delivery of pro-angiogenic growth factors, and creation of a favorable immune microenvironment, these composites synergistically enhance bone defect healing beyond the capabilities of hydrogels or bone powder alone.

Building on these characteristics, the physicochemical characteristics of bone powder-laden hydrogel scaffolds, including pore structure, surface roughness, mechanical properties, and water absorption, interact with each other and collectively determine their biological functions. First, the pore size, porosity, and interconnectivity of the scaffold directly influence cell migration, proliferation, and nutrient exchange.^12,90 Pores ≥300 μm favor new bone formation and vascularization, while highly porous scaffolds facilitate the transport of nutrients and signaling molecules, thereby enhancing the activity of osteogenic cells and endothelial cells. In other words, the pore structure improves the microenvironment and indirectly promotes vascularization and osteogenic differentiation. Second, surface roughness and microtopography can enhance cell adhesion and spreading, which in turn upregulates the expression of osteogenic genes such as RUNX2, ALP, and OCN.^20,22 This physical signal regulates MSC differentiation through mechanosensing mechanisms, thereby promoting bone regeneration. In addition, appropriate scaffold stiffness and elastic modulus can influence osteogenic differentiation of bMSCs via mechanical cues.^14,23 For example, bone powder-reinforced hydrogels increase compressive strength and storage modulus, providing stable mechanical support under load, preventing collapse, and facilitating mechanotransduction that guides cells toward osteogenic differentiation. Moreover, moderate water absorption ensures fluid mobility and factor diffusion within the scaffold, supporting cellular metabolism and the release of angiogenic factors such as VEGF, thereby creating a microenvironment conducive to vascularization.^8,23 Finally, uniformly dispersed nano-bone powder not only enhances mechanical performance but also provides abundant osteoconductive cues and a favorable local microenvironment to promote osteogenic differentiation and vascularization.^16,104 Nanoscale HA particles significantly improve cell adhesion, proliferation, and osteogenic gene expression, while also modulating the microenvironment to support vascularization. In summary, pore structure and water absorption promote vascularization and nutrient exchange, surface roughness and bone powder dispersion enhance osteogenic cell adhesion and gene expression, and mechanical properties provide a stable microenvironment and mechanical signals. These features work synergistically to ultimately promote bone defect repair and vascularization.

4. The functions and applications of bone powder-laden hydrogel

The tissue engineering scaffolds based on different materials have been widely used in bone regeneration. Currently, the concept of using hydrogel scaffolds to promote bone regeneration is still in the preclinical development and animal testing stages. To date, no synthetic hydrogel product has been approved for the regeneration market.¹⁰⁷ The functionality of bone powder-laden hydrogels can be verified through cell experiments and animal studies.

4.1 In vitro experiments in osteogenesis-related cells

To investigate the osteoinductive capability and biocompatibility of bone powder-laden hydrogel scaffolds in vitro, researchers quantitatively assess several factors, including cell viability, cell proliferation, cell morphology, cell migration and recruitment, and osteogenic gene expression in cells co-cultured with the scaffolds. The cells typically used for these experiments are bone marrow mesenchymal stem cells with osteogenic potential, sourced from humans, rabbits, and rats. Ideal scaffolds in vitro are expected to promote cell proliferation, differentiation, migration, and the expression of osteogenesis-related genes.

Cell viability under scaffold culture has been measured using a live/dead cell staining kit. The composite of bone powder-laden hydrogel scaffolds and BMSCs was incubated for 5 days, with live/dead cell counting performed on Day 1, Day 3, and Day 5. In brief, the hydrogel/BMSC composite was rinsed with sterile PBS and then incubated with a working solution containing calcein acetoxymethyl ester (calcein-AM) (which stains live cells green) and ethidium homodimer-1 (which stains dead cells red) for 30 minutes at 37 °C. The composite was subsequently examined under a fluorescence microscope using 488 nm and 568 nm excitation.¹⁰ A large proportion of the MSCs stained green and remained viable, tightly adhering to the scaffold surface, demonstrating that the scaffold effectively supports MSCs attachment without exhibiting cytotoxicity.²¹

There are several quantitative methods for measuring cell proliferation, with the Cell Counting Kit-8 (CCK-8) and methylthiazolyldiphenyl-tetrazolium bromide (MTT) assays being commonly used. The CCK-8 assay is employed to quantitatively assess the cell viability and proliferative activity of MSCs on different scaffolds. Typically, a 10% CCK-8 solution is added, and the samples are incubated in a 5% CO₂ environment at 37 °C for 2 hours, followed by measurement at a wavelength of 450 nm using a microplate reader. Additionally, AO/EB fluorescent staining has been used to further observe cell viability.⁸ As the cell culture time increased, the number of MSCs grew significantly, with no detectable differences between the control and experimental groups at the same time points, indicating that the prepared scaffolds exhibited satisfactory cell compatibility.²¹ The MTT assay for measuring cell proliferation involves adding MTT solution at the designated time, followed by incubation to form formazan crystals. After 4 hours, the MTT solution is removed, and 200 μl of dimethyl sulfoxide (DMSO) is added to dissolve the formazan crystals. The optical density (OD) at 492 nm is then measured using a microplate reader.^10,16 According to the ISO10993-5:2009.46 standard, cell viability in all scaffold groups exceeds 80%, demonstrating that the composite scaffolds are non-toxic and exhibit excellent biocompatibility.¹⁰⁸

To investigate cellular morphology and distribution, in addition to directly utilizing SEM images to display cellular morphology and their podia on the scaffold, an alternative procedure is to stain with 4′,6-diamidino-2-phenylindole (DAPI) reagent. After rinsing off the staining solution, the scaffold can be photographed under a fluorescence microscope.¹⁴ Tetramethyl rhodamine isothiocynate (TRITC)-phalloidin reagent can also be used simultaneously with DAPI or independently to demonstrate cell adhesion and elongation, allowing for the observation of cellular pseudopodia growing on the substrate.^12,18

The influence of scaffolds on the migratory activity of osteogenesis-related cells can be quantitatively assessed by Transwell assay. The permeability characteristics of hydrogels in vitro cellular states can also be observed in the Transwell model, demonstrating the hydrogels’ recruitment capability in vivo.¹⁰⁹ Transwell assay is frequently employed to measure endothelial cell recruitment, confirming that scaffolds support neovascularization. Zhang et al. invented the HA hydrogel scaffold loaded with manganese carbonyl (MnCO) nanosheets. A Fenton-like reaction occurs between MnCO and endogenous hydrogen peroxide produced at the implant-tissue interface, resulting in the sustained release of carbon monoxide and Mn²⁺. This process notably attenuates the inflammatory response by enhancing the M2 phenotype expression in macrophages, which subsequently secrete vascular endothelial growth factor to promote blood vessel formation.²¹ In their cell experiments, Transwell assay was used to determine the chemotaxis effect of stents on human umbilical vein endothelial cells (HUVEC).

To investigate the impact of bone powder-laden hydrogels on osteogenesis-related cell proliferation and osteogenic differentiation in vitro, the osteogenic activity of the scaffolds can be evaluated through ALP staining, alizarin red S (ARS) staining, western blot (WB) and quantitative Reverse Transcription Polymerase Chain Reaction (qRT-PCR). Differentiation and mineralization levels of osteogenesis-related cells were analyzed via alizarin red staining at Day 7 or Day 14, with ALP staining conducted at Day 7. Both alizarin red and ALP staining were performed according to the manufacturer's protocols, with samples imaged under a microscope.¹³ As a well-known biomarker of osteogenic differentiation, ALP plays a crucial role in bone mineralization by stimulating pyrophosphate hydrolysis.¹¹⁰ Ideal bone powder-laden hydrogel scaffolds exhibited positive ALP staining results at all time points in osteogenic induction medium. ARS staining analyzes the mineralization stage of osteogenic differentiated cells by staining for red calcium nodules. In Fig. 2, the bone powder-laden hydrogel group exhibited intensified red coloration compared with the control group over the culture period, indicating that the scaffolds more effectively promoted osteogenic mineralization of cells.²¹ In addition, WB analysis and qRT-PCR were performed to reflect the osteogenic gene expression. For example, Col I, OCN, OPN, RUNX-2, and ALP are the most frequently tested osteo-specific gene expression markers. The effect of the scaffold in promoting osteogenesis was assessed by the extent of gene expression (Fig. 3).⁹

Fig. 3 The functions and applications of bone powder-laden hydrogel were tested in vitro and in vivo. (A) In vitro experiments in MSCs. (B) In vivo studies in rabbits and rats models with bone defects. (C) Live/dead staining of MSCs cultured on bone powder-laden hydrogel scaffolds. (D) ALP staining and ARS staining of each group of MSCs,¹⁵ reproduced from ref. 15 with permission from Wiley, Pearl Powder Hybrid Bioactive Scaffolds from Microfluidic 3D Printing for Bone Regeneration, copyright 2023. (E) CCK-8 assay quantification for 1 to 7 days post seeding. (F) Osteogenic differentiation tendency of MSCs cultured on bone powder-laden hydrogel scaffolds based on evaluation of RUNX2, OPN, BSP and ALP expressions (n = 4; *indicates significant differences, p < 0.05; ** indicates highly significant differences, p < 0.01),⁸ reproduced from ref. 8 with permission from Royal Society of Chemistry, Hydrogel composite scaffolds with an attenuated immunogenicity component for bone tissue engineering applications, copyright 2021. (G) Quantitative micro-CT analysis of bone mineral density and (H) new bone volume fraction of regenerated rat cranial bone defects from each group (n = 3; *indicates significant differences, p < 0.05; ** indicates highly significant differences, p < 0.01). (I) Representative images of HE staining, Masson's trichrome staining, BMP-2 staining, as well as immunohistochemical staining for bone regeneration markers OCN and angiogenesis markers CD34 of the bone sections at 12 weeks post-implantation. Scale bar is 500 μm and scale bar for magnified images is 200 μm,¹³ reproduced from ref. 13 with permission from IOP Publishing Ltd, Incorporation of small extracellular vesicles in PEG/HA-Bio-Oss® hydrogel composite scaffold for bone regeneration, copyright 2024. (J) 3D micro-CT images of regenerated rat cranial bone defects from each group implanted with bone powder-laden hydrogel scaffolds,⁸ reproduced from ref. 8 with permission from Royal Society of Chemistry, Hydrogel composite scaffolds with an attenuated immunogenicity component for bone tissue engineering applications, copyright 2021.

4.2 In vivo studies in bone defect animal models

In vitro evaluations are frequently employed to appraise the biocompatibility, mineralization, and cytotoxicity of biomaterials. Yet, they fall short of fully mirroring the intricate nature of the in vivo setting, notably due to the lack of immune or inflammatory reactions. Furthermore, the toxicity levels of materials are often exaggerated in vitro, and investigations typically confine themselves to acute toxicity effects owing to the limited lifespan of cultured cells. Additionally, in vitro studies are commonly performed in a 2D setting that fails to emulate the 3D environment of tissues and organs. Consequently, the clinical relevance of in vitro evaluations of biomaterials remains questionable.¹¹¹ For the preclinical stage of development, in vivo studies in animal models are able to provide important data and knowledge.

In selecting experimental models, various factors such as animal species, animal age, defect size, method of defect creation, duration of the study, and fixation techniques can all influence bone repair following implantation. Most importantly, the chosen animals should exhibit physiological and pathological conditions similar to those in humans. Among the animal models commonly used for bone powder-laden hydrogel implantation experiments, the most prevalent bone defect models are those in rats and rabbits, with studies on cranial and femoral defects receiving more attention than those on other body parts. The calvarial defect is frequently utilized as a non-load-bearing model to investigate scaffolds that possess inferior mechanical properties compared with bone. Conversely, the osteogenic potential of scaffolds with properties akin to bone is typically assessed in load-bearing long bone defect models, such as those involving the femur, tibia, radius, and humerus.¹¹²

After surgical implantation of bone powder-laden hydrogel scaffolds in animals, imaging analysis (X-ray and micro-CT), histological and immunohistochemical analysis, and gene expression analysis are typically conducted sequentially. Under X-ray, if no obvious fractures are observed in the defect area after hydrogel implantation, this indicates that the composite hydrogel could be used for bone tissue substitution and load-bearing applications. Micro-CT further allows for the analysis of the structure of newly formed bone in the femoral defect area, with results demonstrating that the bone powder-laden hydrogel exhibits superior osteogenic performance compared with single-component formulations.²⁴ Histological and immunohistochemical analyses are employed to evaluate the details of bone formation, and can not only identify structures such as trabeculae and the quality of bone integration, but also detect the presence of fibrous tissue, collagenous tissue, and osteogenesis-related cells. The expression of osteogenic genes is assessed through staining intensity.¹⁷ Additionally, polymerase chain reaction (PCR) results can also demonstrate the activity level of gene expression (Fig. 3).

4.3 In vitro evaluation of immune responses

To further assess the immunocompatibility of bone powder-laden hydrogel scaffolds, it is essential to incorporate assays targeting immune cell responses in addition to osteogenesis-related in vitro evaluations. Human peripheral blood mononuclear cells (PBMCs) can serve as a representative in vitro immune model. In this approach, PBMCs are co-cultured with scaffold extracts for 24–48 h, followed by quantitative analysis of inflammatory cytokine secretion using ELISA. Specifically, the levels of tumor necrosis TNF-α, IL-1β, and IL-6 are measured. A significantly lower cytokine release in the scaffold-treated group compared with the positive control (e.g., lipopolysaccharide treatment) would indicate reduced immunogenicity, consistent with previous findings on α-1,3-galactosyltransferase knockout porcine bone tissue used for xenotransplantation.³⁰

Furthermore, T-cell activation can be examined by flow cytometry. PBMCs seeded at 1 × 10⁶ cells per mL in 24-well plates are cultured with scaffold extracts for 16–48 h, then stained with antibodies against CD3, CD4, CD69, and CD25 for gated analysis. Comparable proportions of CD4⁺,CD69⁺ and CD4⁺,CD25⁺ cells between scaffold-treated groups and negative controls, along with a significant reduction relative to anti-CD3/CD28-stimulated positive controls, would suggest that the scaffolds effectively suppress aberrant T-cell activation.

In addition, molecular pathways underlying macrophage responses can be investigated in PBMC- or THP-1-derived macrophages by monitoring phosphorylation of key proteins in the NF-κB and MAPK signaling pathways, including p-p65, p-ERK, and p-p38. Samples are collected at 0, 0.5, 1, and 2 h post-treatment, and phosphorylation-to-total protein ratios are quantified by WB. A significant reduction in pathway activation in the scaffold-treated groups would further substantiate their immunoregulatory potential.

Taken together, these immune-related in vitro assays not only provide evidence of the low stimulatory effect of bone powder-laden hydrogels on immune cells but also offer mechanistic insights that support the favorable tissue integration and minimal rejection observed in vivo.³⁰

5. Limitations and future prospects

The goal of bone tissue engineering is to develop functional strategies for the most challenging bone-related clinical problems. Even though the development and functionality of bone powder-loaded hydrogels have been evaluated through cellular and animal experiments, there are still various issues that limit their clinical application (Table 2).

Table 2 Challenges and future directions of bone powder-laden hydrogel scaffolds

Existing challenges	Future directions
Abbreviations: SPC, statistical process control; PAT, process analytical technology.
Poor mechanical and biological properties of bone powder	1) Select appropriate acids during the demineralization process.
	2) Avoid the use of chemical substances such as hydrogen peroxide that affect osteoinductivity.
	3) Incorporate a certain proportion of metal ions into bone powder composites to improve mechanical properties and biocompatibility.
	4) Consider differences in biomolecule content among different animal breeds when selecting bone powder.
Weak mechanical properties and limited function of the hydrogels	1) Use double network hydrogel.
	2) Introduce 3D printing technology
	3) Introduce nanoparticle composite technology
	4) Optimize synthesis process of double-network hydrogels to reduce multi-step complexity and toxic crosslinkers.
Limited delivery of composite hydrogel scaffolds	1) Develop injectable hydrogel scaffold.
	2) Prepare targeting scaffold based on microgels.
	3) Use “smart” hydrogel composite scaffolds.
	4) Develop nanogel scaffold.
	5) Employ pH-responsive or other stimulus-responsive “smart” hydrogel systems for personalized drug release and improved targeting.
Negative effects produced by the loading of bone powder	1) The construction of tubular framework composite hydrogel scaffold.
	2) Esterification, polymer grafting, and coupling agent modification.
	3) Utilize silane coupling agents or other surface treatments to improve interfacial bonding and reduce particle aggregation.
Scale-up, commercialization, and risk control challenges	1) Implement strict donor and material stratification with biochemical quantification of osteoinductive proteins.
	2) Apply SPC and PAT to ensure process consistency.
	3) Adopt advanced sterilization methods such as supercritical CO2 to balance biosafety and bioactivity preservation.
	4) Incorporate hybrid material strategies and automation to reduce costs.
	5) Conduct standardized multi-batch and biosafety validations to meet regulatory requirements.

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5.1 Challenges for the laden bone powder

Bone powder has been commonly utilized in bone tissue engineering due to its excellent biocompatibility, elemental biomechanics, low antigenicity, and potential osteoinductivity.¹¹³ However, it also has several difficulties in its preparation and application. Several bone powder materials incorporated into hydrogels have been developed for the treatment of bone defects and voids. However, variations in the source of the bone powder may lead to differences in performance. For instance, compared with BC, substituting them with particles derived from natural bone may simplify cell–substrate recognition and matrix deposition, as they provide a natural microstructural scaffold, which is a well-established requirement for appropriate matrix deposition.¹¹⁴ Since bone tissue engineering aims to provide bone substitutes that mimic the structure and composition of native bone as closely as possible, animal-derived bone powder is currently considered to offer distinct advantages over other types of bone powder.¹¹⁵ We therefore infer that animal-derived bone powder with greater similarity to human bone in both structure and composition may promote superior bone regeneration outcomes in vivo, although this hypothesis requires validation through future clinical studies. Furthermore, researchers have observed that the efficacy of hydrogel-based grafts loaded with bone powder may also vary depending on the source species or even different breeds within the same species. Zhao et al. found that three composite scaffolds composed of hydrogel and DBPs from three different sources [Yeonsan Ogye (YO), Cornish cross, and Pekin duck] exhibited differing capabilities in promoting bone regeneration-related proliferation and differentiation. Although the structural and compositional properties of these bone powders were demonstrated to be highly similar, the observed differences were potentially attributable to their varying melanin content.⁹ In particular, the YO melanin is reported to promote osteoblast differentiation, inhibit osteoclast differentiation, strengthen the bone, and increase ALP activity and bone mineralization.¹¹⁶ Consequently, future research may focus not only on the further development of animal-derived bone powders, but also on differences in biomolecule content among different animal breeds, particularly when their skeletal structures and compositions are similar.

DBMs have been widely used for their osteoinductive properties which benefit interaction with target cells and surrounding bone tissue at the graft site.⁷⁴ However, numerous experiments have mentioned that demineralization and irradiation can diminish the osteoinductive properties of DBMs. Researchers have examined allogeneic bone grafts demineralized by HCl and EDTA, noting that HCl-demineralized bone exhibited slightly more positive bone formation outcomes compared with EDTA, although the specific mechanisms remain unclear.¹¹⁷ In addition to the processes of demineralization and irradiation, chemical substances such as hydrogen peroxide also impact osteoinductivity. The organic components of DBMs primarily consist of two parts: collagen and non-collagenous proteins, with the latter almost entirely responsible for all the biological functions of DBMs, including osteoinductivity. BMP-2, as one of the non-collagenous proteins in DBM, is believed to promote osteogenic differentiation of stem cells and the formation of new bone, and is closely related to the reduction in osteoinductivity.^118,119 It is speculated that chemical substances such as hydrogen peroxide may lead to a decrease in osteoinductivity by impairing the function of BMP, and should therefore be avoided in bone regeneration experiments.

Mechanical properties and biocompatibility are two properties that are of concern in bone powder. Fadeeva prepared brushite cement based on powders of Zinc (Zn) substituted β-TCP and found it promising in view of its significant antibacterial activity, improved mechanical strength and enhanced cell viability.¹²⁰ Similarly, MgZn/HA composites were prepared using powder metallurgy (PM) by Lu et al. They proved that the mechanical properties and biocompatibility of the composite scaffold were significantly affected by the HA proportion, which showed improved porosity, increased mechanical properties, and enhanced corrosion resistance with a high MgZn content.¹²¹ It must be acknowledged that the biosafety and long-term stability of bone powder-laden hydrogels must be thoroughly verified before clinical translation. This indicates the focus of future research. To sum up, the loading of metal ions with a moderate proportion in the bone powder mixture is an idea for improving the mechanical properties and biocompatibility of the composite scaffolds.

5.2 Challenges for hydrogel preparation

Hydrogels can be prepared to possess simultaneous encapsulation and delivery capabilities, as well as the ability to mimic tissue mechanical properties and integrate with adjacent tissues, thereby exhibiting immense potential in biomedical applications.¹²² Despite this, many conventional hydrogels exhibit weak mechanical properties and limited functions, making the problem one of the most challenging issues for gel researchers. To date, several hydrogels with exceptional properties have been developed (Fig. 4).

Fig. 4 Several novel hydrogels targeted at bone defects in bone tissue engineering. These strategies have potential for the design of bone powder-laden hydrogels.

Double-network (DN) hydrogels, a type featuring an interpenetrating polymer network (IPN) structure, exhibit stronger mechanical properties and enhanced stability when compared with conventional single-network hydrogels.¹²³ This unique network topology is typically synthesized through a two-step sequential free-radical polymerization process, utilizing a polyelectrolyte as the initial network hydrogel and a neutral polymer as the subsequent network.¹²⁴ The use of DN hydrogels is an effective strategy for enhancing the mechanical properties of hydrogels. The mechanical performance of these DN hydrogels can be modulated by altering the content and structure of either network. Owing to their robust characteristics, DN hydrogels hold potential for various applications, including drug delivery carriers, tissue engineering scaffolds, and intelligent devices.¹²⁵ However, current methods for synthesizing DN hydrogels face limitations such as multi-step procedures, complex preparation conditions, and the use of toxic chemical crosslinkers.¹²⁶ These issues restrict their large-scale production and necessitate further optimization of the fabrication process for practical in vivo applications, including drug delivery and bone defect repair.¹²⁷

In addition, some current advanced technologies are also being utilized in the fabrication of hydrogels, such as 3D printing technology and nanoparticle composite technology. These advancements offer new possibilities for overcoming the shortcomings of traditional materials in bone repair, including through the rational design and manufacture of hydrogels with desired structures and properties, encompassing good mechanical performance and high biological activity.¹²⁸

5.3 Challenges for the combination of components

Although it has been demonstrated that the incorporation of bone powder enables the composite hydrogels to exhibit enhanced mechanical properties, osteogenic differentiation, and adjustable biological characteristics, thereby maximizing the performance of individual components, the loading of bone powder bears corresponding adverse effects. For instance, as the concentration of bone powder increases, the storage modulus of the scaffolds also rises. However, Gao et al. found that when the DCB content reaches 50%, the storage modulus decreases. This phenomenon is attributed to the aggregation of DCB particles at excessively high concentrations, leading to an uneven distribution of DCB particles within the scaffolds, and consequently, the 50% DCB scaffolds exhibit a lower storage modulus.²² Additionally, with the incorporation of HA, the open porosity of GelMA decreases.¹²⁹ Hou et al. indicated that the HA in these studies predominantly exists in the forms of micro-/nanorods, wires, and particles, which scarcely contribute to either the pores or their interconnectivity. Therefore, they constructed a tubular framework composite hydrogel scaffold using hydroxyapatite microtubes (HAMT) based on GelMA. The tubular HAMT enhanced the mechanical properties of the scaffold and interconnected adjacent pores on the scaffold, while simultaneously roughening the surface of the composite scaffold. Consequently, this composite scaffold exhibited superior adhesive properties, mechanical performance, gas permeability, and signal transmission capabilities.¹⁰⁶ This approach merits consideration by scholars for application in improving the morphology and dispersion of bone powder within composite materials.

On the other hand, the compatibility and interfacial bonding strength between inorganic bioceramics and hydrogels are rather poor. There are numerous approaches for addressing this issue, most of which involve surface treatment of the bioceramics, including esterification, polymer grafting, and coupling agent modification.¹⁰⁰ Among them, coupling agent modification is the most widely used method to improve the interfacial bonding between bioceramics and biopolymers, such as silanes, titanates, and zirconates. This method can enhance the bonding properties between bioceramics and biopolymers, improving the degree of interfacial bonding between the two phases.¹³⁰ Öner et al. utilized silane coupling agent (3-methacryloxypropyltrimethoxysilane) as a “bridge” between HA and HA-laden polymers to enhance the interfacial adhesion strength of the composite material. Silanization of HA is an effective strategy to reduce the surface energy of the particles, thereby overcoming agglomeration of particles within the polymer.¹³¹

5.4 Challenges for the form and the delivery of the composite scaffold

In the delivery of composite hydrogels, the clinical promise of new hydrogels continues to evolve rapidly, driven by the growing need to address emerging clinical problems.¹³² Injectable hydrogel scaffolds possess advantages over traditional implantable scaffolds in minimally invasive surgical procedures, as they can be administered to various irregular bone defects with minimal discomfort while conforming closely to the application site.¹⁰⁷ Various biomaterials can serve as ideal candidates for the development of injectable hydrogels, such as hyaluronic acid, chondroitin sulfate, chitosan, collagen, and polypeptides, among others.¹²⁸ Based on this, scholars have developed numerous injectable smart targeted hydrogels with diverse properties.

The traditional bulk composite hydrogels, which are commonly used as scaffolds in bone tissue engineering and carriers for bone powder delivery, exhibit several drawbacks, including large size, low specific surface area, lack of micropores, and slow degradation rates. These disadvantages lead to poor cellular infiltration and viability, as well as weak vascularization.¹³³ A promising solution to address these issues is the targeted carrier system based on microgels or nanogels. Microgels are colloidal particles with macromolecular networks, ranging in size from tens of nanometers to several micrometers.¹³⁴ Microgels possess various characteristics, including injectability, microporous structure, heterogeneity, and surfaces that can be modified.¹³⁵ On this basis, the incorporation of different monomers copolymerized with functional groups, the encapsulation of inorganic nanoparticles, or the addition of targeting ligands can endow the microgel system with intelligent properties.¹³⁴ Targeting carriers based on microgels can elevate local drug concentrations and augment the therapeutic efficacy. Nevertheless, their small size and rapid degradation impede their ability to achieve sustained long-term drug delivery. Nanogels are spherical hydrogels with a nanoscale (<100 nm) size fabricated through physically or chemically crosslinked aqueous networks. In terms of nanosize and surface area, nanogels exhibit broader applications in drug delivery systems compared with microgels.¹³⁶ Owing to their unique properties, nanogels have been utilized in a wide range of fields, including biochemistry, biomedicine, photonics, bionics, pollution control, and cosmetics.¹³⁷ Amiri et al. constructed nanocomposites using hydrogel beads, achieving promising applications in antibacterial activity and molecular delivery.¹³⁸ These novel hydrogel materials represent a potential avenue for improving bone powder-loaded hydrogel scaffolds.

“Smart” hydrogels have become a prominent research hotspot in recent investigations, and respond to external stimuli, such as pH, light, temperature.¹³⁹ Among these, pH-responsive hydrogel composite scaffolds, which react to environmental pH and ionic strength, play a significant role in promoting bone defect repair. The inflammatory phase of bone regeneration exhibits distinct pH variations, which pH-responsive composite scaffolds can exploit. The pH differences between normal tissues and lesion sites trigger sol–gel transitions in the hydrogels, thereby improving targeting precision and enabling personalized delivery strategies.^140,141 “Smart” hydrogels not only modify their properties in response to these factors but also actively release drugs that influence the same parameters, creating a self-regulating system where drug release kinetics are continuously adjusted by the evolving microenvironment while simultaneously controlling inflammation.¹⁴² Zhang et al. proposed smart drug depot hydrogels, which enable intelligent targeted drug release, achieving long-term therapeutic delivery through a single administration.¹³⁵ However, the practical applications of smart hydrogel-based drug delivery systems remain superficial, currently limited primarily to skin and the ocular surface.¹⁴³ Furthermore, numerous challenges, including those related to biocompatibility, toxicology, and biodegradation, must be addressed. Specifically, their safety for applications such as bone defect repair requires thorough validation.¹⁴⁴

5.5 Challenges for scale-up, commercialization, and risk control

Although bone powder-laden hydrogels demonstrate strong osteogenic potential in vitro and in animal models, several obstacles hinder their clinical translation and large-scale production. One of the major issues is the variability of raw materials. Multiple studies have shown that the donor source, anatomical site, and age significantly affect the protein content of DBM, especially bone morphogenetic proteins such as BMP-2 and BMP-7. For example, Urist and colleagues reported that the osteoinductive activity of DBM varied more than fivefold across donor sites, largely due to differences in BMP content.¹⁴⁵ More recently, Zhao et al. demonstrated that femoral DBM contained nearly twice as much BMP-2 as iliac-derived DBM and further validated guanidine hydrochloride (GuHCl) extraction as a reliable method for BMP quantification.¹⁴⁶ These findings suggest that strict donor stratification and biochemical screening are necessary. Limiting donor age (e.g., <50 years), choosing standardized anatomical sites such as femoral shafts, and conducting rapid pre-screening of protein, collagen, and ash content, together with quantitative assays such as ELISA or proteomics, could establish batch acceptance thresholds to reduce variability.

Another challenge is batch-to-batch consistency during processing. Even within the same commercial DBM product, significant variability in osteoinductivity has been documented.¹⁴⁷ To address this, statistical process control (SPC) and process analytical technology (PAT) should be integrated across the production pipeline, including demineralization, washing, sterilization, and particle size reduction. Raman spectroscopy and near-infrared (NIR) monitoring, for example, have been successfully used to track residual mineral content during acid demineralization, while online laser diffraction analysis can ensure that the particle size distribution remains within the 100–500 μm range optimal for scaffold integration.¹⁴⁸ Such approaches minimize uncontrolled variability and improve product reproducibility.

Biosafety is another critical concern. Xenogeneic or allogeneic bone powders may contain pathogens, including viruses or prions, and conventional high-dose γ irradiation (>25 kGy) is known to damage collagen ultrastructure and reduce BMP bioactivity by over 40%. In response, supercritical CO₂ (sCO₂) processing has emerged as a promising alternative. Preclinical studies show that sCO₂-sterilized DBM retains >90% of its collagen structure and osteoinductive activity compared with untreated controls, while effectively eliminating microbial contaminants.¹⁴⁹ When combined with mild oxidizing agents such as peracetic acid, this method achieves sterility assurance levels required by regulatory authorities without significant protein degradation. Post-processing assessments, including DNA quantification (<50 ng mg⁻¹ dry weight), endotoxin testing (<0.5 EU mg⁻¹), and qPCR viral screening, should serve as release criteria.

Cost and scalability also present barriers. Recombinant BMPs, while effective, remain prohibitively expensive, with costs exceeding $5000 per patient for spinal fusion. Hybrid strategies have been shown to reduce costs while preserving osteoinductivity, such as blending screened DBM with synthetic hydroxyapatite, recombinant collagen, or microencapsulated low-dose BMPs.¹⁵⁰ Additionally, the introduction of automated continuous or semi-continuous GMP-compliant production lines equipped with PAT monitoring can reduce labor costs and defect rates by up to 30%, enhancing scalability and economic feasibility.¹⁴⁸

Finally, regulatory approval requires rigorous validation. Multi-batch studies (≥3–5 independent lots) should demonstrate consistent osteogenic activity through in vitro assays (e.g., ALP induction >2-fold relative to control) and in vivo ectopic bone formation models. Post-sterilization bioactivity, biosafety evaluations (microbial culture, endotoxin, residual DNA), and mechanical testing (compressive strength ≥2 MPa, fatigue resistance >10⁴ cycles) are essential datasets for submission.¹⁵¹

In summary, the clinical translation and industrialization of bone powder-laden hydrogel scaffolds will require a multi-pronged strategy: (1) rigorous donor and material stratification with biochemical quantification of osteoinductive proteins; (2) application of SPC and PAT to ensure process consistency; (3) adoption of advanced sterilization methods such as sCO₂ to balance biosafety and bioactivity preservation; (4) incorporation of hybrid material strategies and automation to reduce costs; and (5) standardized multi-batch and biosafety validations to meet regulatory requirements. Together, these measures could bridge the gap between preclinical promise and clinical application.

It is undeniable that each type of bone graft biomaterial has its drawbacks, including those related to immune response, quantity, quality after the manufacturing process, and rapid resorption, among others, suggesting that no single material may be suitable for treating all conditions. Through preclinical experiments, scholars should design the optimal material for different types of conditions based on the inherent properties of the material, the type of bone defect to be treated, the operator's preference, associated costs, and patient acceptance.²⁵

6. Conclusions

In recent years, hydrogels loaded with bone powder have garnered considerable attention in bone tissue engineering for the repair of bone defects. Bone powder and hydrogels represent relatively ideal choices among bone graft materials and have been extensively studied by scholars. Furthermore, compared with the use of bone powder or hydrogel scaffolds alone, composite hydrogel scaffolds loaded with bone powder exhibit superior performance in physicochemical characterization, biological functionality, and in vitro and in vivo experiments. In this paper, several studies on composite materials loaded with bone powder hydrogels are summarized, all of which attempt to combine various types of bone powder with hydrogels in different ways and evaluate the characterization and results of these hydrogels in vitro and in vivo to verify their feasibility and advantages. This review summarizes the material selection and synthesis processes of bone powder-laden hydrogels, discusses their functions, advantages, and applications, and concludes by highlighting the current limitations and future prospects of these products, providing directions for future research and development of bone tissue engineering materials.

Author contributions

Xinyi Shen: writing – original draft, data curation, formal analysis, investigation, visualization. Danji Zhu: writing – review and editing, conceptualization, investigation. Haorui Hu: writing – review and editing, investigation, formal analysis. Lingkai Su: writing – review and editing, supervision. Gang Wu, Tim Forouzanfar: writing – review and editing. Guoli Yang: funding acquisition, supervision, resources. Zhiwei Jiang: methodology, project administration, supervision, resources.

Conflicts of interest

The authors declare no conflict of interest.

Data availability

No primary research results, software or code have been included and no new data were generated or analysed as part of this review.

Acknowledgements

This work was supported by the National Natural Science Foundation of China (Grant Nos: 81901051 and 82370991). The funders had no role in study design, data collection, and interpretation, or the decision to submit the work for publication.

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Footnote

† Authors contributed equally to this article.

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