Near‐Infrared Responsive Property and Nanozyme Effect‐Mediated 3D‐Printed Gradient Bioactive Scaffold for Intelligent Immunomodulation in Regeneration of Large Segmental Bone Defects

ABSTRACT

Regeneration of large segmental bone defects remains a formidable challenge, mainly due to the lack of a supportive microenvironment for the host's intrinsic repair cascade. Although 3D‐printed porous calcium phosphate ceramics have osteoconductive properties, their inability to actively regulate key immune processes and compromised mechanical strength caused by porosity greatly limit therapeutic efficacy. Herein, an “osteoimmunology”‐informed design of a near‐infrared (NIR)‐responsive 3D‐printed gradient scaffold is proposed for intelligent bone regeneration by temporally modulating macrophage phenotypic polarization. Fabricated via digital light processing (DLP) printing combined with secondary sintering, this scaffold has a core–shell‐like gradient structure (dense interior, porous exterior), achieving a balanced combination of mechanical robustness and bioactivity. Manganese iron Prussian blue analogue (MnFePBA) nanozymes were incorporated onto the scaffold surface. Under NIR irradiation, MnFePBA not only scavenges excessive reactive oxygen species (ROS) but also triggers controlled release of Mn^{2 +}, synergistically driving macrophages to polarize from proinflammatory M1 to proregenerative M2 phenotype. In vitro and in vivo studies confirm that the scaffold alleviates post‐implantation oxidative stress, fosters a favorable immune microenvironment, and significantly enhances bone repair, establishing a new paradigm for developing next‐generation smart bone grafts via active immunomodulation.

Mountains represent the brain, the sun symbolizes Near‐Infrared Radiation, and the river stands for blood. Neutrophils, macrophages, and various ions flow in the water. Nuwa holds materials representing the “Sky‐Mending Stones” to address the skull defect, and all kinds of ions work together to promote the repair of the defective area.

1. Introduction

The regeneration of large segmental bone defects caused by severe trauma, malignant tumors, or congenital deformities represents a core challenge in the field of bone tissue engineering [1, 2]. Due to the severe compromise of skeletal structural integrity, the body struggles to autonomously establish the reparative microenvironment necessary to support the “hematoma formation–immune regulation–callus formation–bone remodeling” cascade [3, 4, 5]. Consequently, reliance on biomaterials to act as bridging mediators is essential for driving the repair process. Porous calcium phosphate (CaP) ceramics have garnered considerable attention in the field of bone tissue engineering, primarily owing to their excellent osteoinductivity [6, 7, 8]. Their stable phases primarily include hydroxyapatite (HA), tricalcium phosphate (TCP), and biphasic calcium phosphate (BCP) [9, 10]. The microporous structure retained within the material matrix via incomplete sintering not only facilitates the regulation of active factor recruitment and secretion but also, through its calcium‐phosphorus microenvironment, enables the directional induction of osteogenic differentiation in stem cells and osteoprogenitor cells, thereby providing critical support for the repair of large segmental bone defects [11, 12]. When combined with 3D printing technology, this approach further enables the customization of CaP ceramics for patients’ personalized defect sites, an advantage that has attracted widespread interest in clinical practice [13, 14]. However, the relatively homogeneous structure and composition of porous CaP ceramics limit their capacity to actively modulate immune responses [15]. When coupled with the frequent mechanical inadequacy arising from their porous architecture, these factors introduce substantial uncertainty into the later stages of the repair process [16].

Extensive research indicates that the key to the clinical success of implants lies in achieving favorable osseointegration, a process predominantly regulated by the immune system [17, 18]. The concept of “Osteoimmunology” is specifically used to describe the complex mutual regulatory relationships between bone remodeling cells and immune cells [19]. As the core cells of immune regulation, macrophages participate in the entire process from the initiation of inflammatory responses to new bone formation, and their different subtypes can mediate distinctly different immune responses [20]. During the initial implantation phase, pro‐inflammatory M1 macrophages are rapidly recruited to the defect site, where they enhance osteoclast activity and promote their differentiation [21]. Simultaneously, they generate reactive oxygen species (ROS) via pathways including NADPH oxidase (NOX) and the mitochondrial respiratory chain to clear necrotic debris and implant materials [22]. Appropriate pro‐inflammatory signaling is crucial for initiating vascular endothelial growth factor secretion and promoting angiogenesis [23]. However, if left unregulated, excessive inflammation can inhibit osteoblast differentiation and activate osteoclasts, ultimately leading to osseointegration failure [24]. Anti‐inflammatory/pro‐repair M2 macrophages promote the recruitment and differentiation of osteoblasts on the implant surface by secreting cytokines and growth factors such as BMP‐2, TGF‐β, PDGF, and IL‐10, thereby facilitating the formation of new woven bone [25, 26, 27]. Huang et al. [28] coated MXene Ti₃C₂ nanosheets on the surface of 3D‐printed ceramic scaffolds to achieve a bone repair scaffold with near‐infrared (NIR) light‐responsive osteoimmunomodulatory properties. Therefore, achieving an orderly, temporally coordinated transition from M1 to M2 macrophages during osseointegration is essential for shifting from the inflammatory phase to the regenerative phase [29, 30].

Prussian blue (PB) is an agent approved by the U.S. Food and Drug Administration (FDA) for the treatment of radiation exposure, known for its good biocompatibility, safety, and high photothermal conversion efficiency [31, 32]. Upon photothermal excitation, intensified vibration of the ─CN─ groups within its crystal lattice can lead to decreased stability and subsequent release of coordinated iron (Fe) ions [33]. Meanwhile, the weakly acidic microenvironment at the bone defect induces further vibration of the ─CN─ bonds and breaks the coordination bonds, thereby exposing active sites and enhancing the nanozyme activity. In addition, the released Mn²⁺/Fe²⁺ participates in the Fenton reaction, which improves the reactive oxygen species (ROS) scavenging capability [34, 35, 36, 37]. Incorporating manganese (Mn) ions into the Prussian blue lattice via a site‐substitution strategy allows for the construction of the nanoparticles named manganese‐iron prussian blue analogue (MnFePBA) [38]. This material not only retains the photothermal properties and nanozyme activity of Prussian blue but also introduces new functions mediated by Mn ions [39, 40]. Mn, an essential trace element in the human body, plays a vital role in regulating macrophage phenotype switching and clearing excess ROS [41, 42]. Based on this, MnFePBA, through the controlled release of manganese ions, holds promise for precisely modulating the immune microenvironment during osseointegration [43, 44]. This approach aims to achieve an orderly, temporally controlled switch from pro‐inflammatory M1 to reparative M2 macrophages, ultimately facilitating a smooth transition from the inflammatory stage to the regenerative stage.

Addressing the challenge of repairing large segmental bone defects, this study pioneers the design of a 3D‐printed bioactive gradient porous scaffold with active immunomicroenvironment regulation capability. After implantation, the scaffold enables the smooth polarization of macrophages from the pro‐inflammatory phenotype to the pro‐regenerative phenotype, thereby facilitating the sequential and intelligent regeneration of large segmental bone defects (Scheme 1). This intelligent regulatory system incorporates two core design elements: (1) Integration of DLP (Digital Light Processing) printing technology and a secondary sintering process to construct a gradient structure characterized by “dense interior and porous exterior”. (2) Construction of manganese‐iron Prussian blue analog (MnFePBA) nanoparticles via a site‐substitution strategy, where manganese (Mn) ions are doped into the Prussian blue lattice, followed by their efficient deposition onto the surface of the 3D gradient scaffold. Under NIR light excitation, MnFePBA enables the responsive release of Mn ions to actively regulate the local immunomicroenvironment. Results demonstrate that the nanozyme activity of MnFePBA can effectively scavenge reactive oxygen species (ROS) in the defect region and alleviate oxidative stress. Furthermore, the NIR irradiation‐triggered controlled release of Mn ions further promotes the polarization of macrophages from the M1 to M2 subtype, thereby realizing the intelligent regulation of the inflammatory microenvironment. In vitro and in vivo experimental results confirm that the constructed NIR‐responsive and nanozyme effect‐mediated 3D‐printed gradient bioactive scaffold exhibits significant osteogenic capacity. By actively modulating the immunomicroenvironment, this study provides an efficient and intelligent sequential repair strategy for the treatment of large segmental bone defects.

SCHEME 1.

Fabrication and Application of Functionalized Dual‐Enhanced MnFePBA&GBC Integrated Bone Repair Platform with Intelligent Response in Promoting Post‐Surgical Implantation Bone Regeneration.

2. Results

2.1. Synthesis, Optimization and Characterization of MnFePBA

Briefly, a series of MnFePBA samples with varying Mn feeding ratios (designated as MnFePBA 0.6#, 0.9#, and 1.2#) were synthesized via a hydrothermal single‐precursor approach. This process leveraged the reducibility of polyvinylpyrrolidone (PVP) in an acidic aqueous medium, where K₃[Fe(CN)₆] underwent dissociation, with Mn(CH₃COO)₂·4H₂O added at 0.015 g/0.6 mol, 0.0225 g/0.9 mol, and 0.003 g/1.2 mol, respectively [45]. For comparative purposes, Prussian blue (PB) was prepared using the identical protocol but without Mn(CH₃COO)₂·4H₂O (Figure 1a). This synthetic route offered the advantages of simplicity, high efficiency, and no requirement for excess metal solutions [46]. Scanning electron microscopy (SEM) was employed to characterize the surface micromorphology of PB and MnFePBA x# (x = 0.6/0.9/1.2). Results indicated that MnFePBA x# shared a uniform cubic crystal structure with PB. However, elevated Mn feeding ratios induced unstable surface collapse, whereas MnFePBA 0.6# maintained a homogeneous and intact morphology (Figure 1b). Intriguingly, increasing the reaction temperature triggered a gradual morphological transition of MnFePBA 0.6# from cubic to spherical (Figure S1), presumably driven by thermodynamic minimization of surface energy [47]. Particle size analysis revealed a positive correlation between manganese (Mn) ion feeding ratio and particle size of MnFePBA x# (Figure 1c). MnFePBA 0.6# exhibited a mean nanoparticle size of ∼412 nm, while MnFePBA 0.9# and 1.2# increased to ∼780 and ∼1060 nm, respectively. Pure PB, in contrast, had a smaller size of ∼280 nm. Larger Mn atoms substituting for smaller Fe atoms can promote the grain growth of PB, while excessive doping can even induce lattice distortion (Figure S2) [48]. Concomitantly, the solution and powder color shifted progressively from dark blue to light blue with increasing manganese (Mn) ion feeding ratio (Figure 1d and Figure S3). This color change stems from electron transfer between ion (Fe) ions of different valences in PB and lattice ─CN─ groups (responsible for PB's dark blue hue), which is disrupted upon Fe substitution by a manganese (Mn) ion. Energy‐dispersive X‐ray spectroscopy (EDS) further confirmed the uniform distribution of Mn, Fe, C, and N across MnFePBA 0.6# (Figure 1e and Figure S4), validating the feasibility of Mn incorporation into the PB framework via the proposed synthesis strategy. Transmission electron microscopy (TEM) and selected area electron diffraction (SAED) analyses demonstrated that both PB and MnFePBA 0.6# possessed a face‐centered cubic (Fm‐3m) crystal structure (Figure 1f) [49]. X‐ray diffraction (XRD) patterns of MnFePBA 0.6#/0.9#/1.2# were consistent with the standard PB reference (PDF#52‐1907) (Figure 1g), confirming successful manganese (Mn) ion doping into the PB lattice. Notably, the main diffraction peaks of MnFePBA ((200), (220), (400), (420)) shifted to lower angles relative to standard PB, attributed to increased interplanar spacing caused by lattice substitution of Fe with larger manganese ions [50, 51]. Fourier transform infrared (FTIR) spectroscopy (Figure 1h) revealed distinct differences in the position and intensity of ─CN─ stretching vibration peaks among MnFePBA samples with different feeding ratios, arising from altered ─CN─ coordination environments [52, 53]. Overall, since MnFePBA 0.6# still retains the stable microstructure and morphology of PB, it is selected as the functional particle for subsequent studies (referred to as MnFePBA).

FIGURE 1.

Morphology and structure of MnFePBA. (a) Schematic illustration of the preparation method and simulated structure. (b) SEM images, (c) Zeta potential particle size distribution curves (n = 3), and (d) photographs showing color changes of aqueous solutions of PB and MnFePBA prepared with different Mn feeding ratios. (e) EDS spectrum of MnFePBA #0.6. (f) TEM images and corresponding SAED patterns of PB and MnFePBA #0.6. (g) XRD patterns and (h) FTIR spectra of MnFePBA prepared with different Mn feeding ratios.

2.2. Preparation and Characterization of 3D‐Printed CaP‐Based Gradient Bioactive Ceramics (GBC)

First, building on our previous research, a high‐precision green body of CaSO₄/hydroxyapatite (HAP) ceramic was fabricated via digital light processing (DLP) 3D printing [54, 55]. Subsequent gradient sintering produced a reinforced CaP‐based ceramic scaffold, in which HAP whiskers exhibited oriented growth. This oriented growth of HAP whiskers has been confirmed to effectively enhance the mechanical properties of CaP‐based ceramics. In tissue engineering, a rich microporous structure is critical for tissue regeneration, yet porous structures often significantly compromise the mechanical properties of ceramics, presenting an inherent contradiction. This study resolved this issue through an ingenious design. Specifically, after fully impregnating the surface of the aforementioned ceramic scaffold with a nano‐hydroxyapatite (nHAP) slurry, secondary sintering was performed at 1100°C. During this process, the nHAP on the scaffold surface participated in sintering for the first time and formed a highly porous outer layer due to its low density. In contrast, the inner part of the ceramic was further densified with secondary sintering, leading to improved mechanical properties. Ultimately, a gradient structure with “a dense interior and a porous exterior” was formed. To achieve a uniformly porous surface structure, we further investigated the effects of different media and slurry solid contents [56, 57]. Meanwhile, the effects of dispersion medium and nHAP concentration on the morphology of the outer layer were investigated. It was found that when using ethylene glycol as the medium and a 0.5 g·mL⁻¹ nHAP suspension, the sintered samples exhibited fewer cracks, better dispersion, and more uniform surface coating. This condition was adopted for subsequent studies (Figure S5). The scaffold coated once (10HAP) showed visible nHAP particles but incomplete coverage of the internal layer, with large underlying HAP grains still exposed. In contrast, scaffolds coated three times (30HAP) and six times (60HAP) were fully covered by a crack‐free, loose porous nHAP layer while retaining the hierarchical structure of the DLP‐printed internal scaffold, a morphology favorable for cell adhesion and migration (Figure 2c). Additionally, the scaffold still retains the macroporous structure with 3D printing characteristics, where the macropore size is about 400 µm, and the BET surface area is 6.685 m² g⁻¹. The resulting dual‐reinforced CaP‐based bioactive porous scaffolds exhibited well‐defined contours and clear pores(Figure S6). The thickness of the external loose nHAP layer increased with coating cycles, reaching 22.12 µm for the 60HAP scaffold. Both 30HAP and 60HAP achieved 100% surface coverage by the loose nHAP layer. ImageJ analysis showed that surface porosity increased to 22.17% with increasing coating cycles (Figure 2b). Nitrogen adsorption‐desorption isotherm (BET) analysis further demonstrated a significant increase in specific surface area with more external coating cycles, alongside corresponding changes in pore size distribution (Figure 2e). To verify the stability of the external nHAP layer, the scaffolds were ultrasonicated in anhydrous ethanol for 1, 2, 20, and 120 min. Post‐ultrasonication characterization showed that the majority of the nHAP layer was retained, with the loose porous structure intact and no exposure of large internal grains. These results confirm the good stability of the nHAP coating prepared by this method (Figure S8).

FIGURE 2.

Preparation and physical properties of dual‐reinforced HA‐based scaffolds. (a) Schematic illustration of the preparation process. (b) Surface porosity and coverage of the loose nHA coating layer in HAP scaffolds with different external coating cycles. Magnified SEM images of the scaffold (c) surface and (d) cross‐section. (e) Specific surface area and surface pore size distribution analysis. (f) Schematic diagram of reinforcement regulation mechanism. (g and h) visual comparison of mechanical reinforcement for dual‐reinforced scaffolds. (i) Compressive strength of dual‐reinforced vs. unreinforced scaffolds (n = 3) (^* p ≤ 0.05, ^** p ≤ 0.01, ^*** p ≤ 0.01). (j) Stress–strain curves of dual‐reinforced scaffolds and conventional whisker‐reinforced scaffolds.

Aim to further verify the effect of this fabrication process on the compressive performance of the CaP‐based bioactive porous ceramics (Figure 2f), reinforced ceramics with a porosity of 65% (prepared with different coating cycles: 3 and 6 cycles, denoted as 30HAP and 60HAP hereafter) (Figure 2g). All scaffolds exhibited brittle fracture, with stress dropping rapidly after reaching the yield strength (Figure 2j). Among the 65% porous ceramics, the 30HAP scaffolds (after two sintering steps) showed a compressive strength of ∼13.1 MPa, and the 60HAP scaffolds achieved an even higher compressive strength of 16.4 MPa. They showed a 150%–200% strength increase compared with the original homogeneous structure scaffolds. Notably, compared with pure HA porous scaffolds (compressive strength ∼2 MPa), the dual‐sintered porous scaffolds exhibited a 600%–700% strength enhancement (Figure 2i) [58, 59]. The significant enhancement in the mechanical properties of the coated scaffolds can be fundamentally attributed to the microstructural evolution driven by the secondary sintering process. To maintain high bioactivity, pure HAP scaffolds are typically not fully densified during initial fabrication, leaving behind numerous micropores that inevitably act as stress concentration points under mechanical loading. However, in this study, the gradient porous scaffolds underwent an additional sintering step after coating. This secondary sintering provided the thermodynamic driving force for localized densification, effectively healing and reducing the micro‐defects within the ceramic matrix. (Figure S7)

As mentioned earlier, enhancing the mechanical properties of CaP‐based bioactive porous ceramic scaffolds via this sintering process provides a feasible solution to the clinical bottlenecks of CaP materials.

2.3. Characterization of CaP‐Based Gradient Bioactive Ceramics (GBC) Loaded With MnFePBA

PB and its analogues exhibit strong absorption of NIR light in the range of 600–900 nm due to charge‐transfer transitions, and can convert light energy into thermal energy. Results from ultraviolet–visible–near‐infrared (UV–vis–NIR) spectroscopy revealed that the synthesized MnFePBA showed a high absorption peak in the NIR region (Figure 3b), which confirms its effective photothermal conversion capability. To further clarify the photothermal conversion efficiency of MnFePBA, the temperature evolution of MnFePBA #0.6 aqueous solutions (0.25, 0.5, 1.0, and 2.0 mg·mL⁻¹, hereafter designated 0.25 MnFePBA, 0.5 MnFePBA, 1.0 MnFePBA, and 2.0 MnFePBA) was monitored under 808 nm NIR irradiation at different power densities (1.01, 0.80, and 0.69 W·cm⁻²), with real‐time temperature profiles recorded specifically at 0.80 W·cm⁻² (Figure 3a). Under NIR exposure, the 0.25 MnFePBA solution temperature rose from 24°C to over 48°C within 10 min. Notably, the temperature elevation became more pronounced with increasing MnFePBA concentration: after 10 min of irradiation, the 2.0 MnFePBA solution reached 50.5°C. In contrast, pure water exhibited a mere ∼2°C temperature increase under identical conditions (Figure 3a,c). Similarly, increasing the power of NIR further enhances the photothermal conversion efficiency. When the power is 1.01 W, the temperature of the solution can reach 60°C, which indicates that the synthesized MnFePBA possesses controllable photothermal conversion performance (Figure S9).

FIGURE 3.

Responsive Ion Release and In Vitro Biocompatibility Verification of MnFePBA&GBC. Different particles in water under 0.8 W cm⁻² NIR irradiation: (a) Infrared thermal images. (b) UV–vis–NIR spectra. (c) Temperature‐rise curves. (d) Preparation of MnFePBA&GBC. (e) Magnified SEM images of the surface and (f–i) Manganese (Mn) and iron (Fe) ion release behaviors of MnFePBA&GBC with different loadings under NIR “on/off” conditions (n = 3). Assessing MC3T3‐E1 cell proliferation on different scaffolds. (j) Typical CLSM images of live‐stained MC3T3‐E1 cells cultured on different scaffolds. (k)Results of CCK‐8 assay (n = 3) (^* p ≤ 0.05, ^** p ≤ 0.01, ^*** p ≤ 0.01). (l) IF staining of adherent‐stained BMSCs seeded on 2.0 MnFePBA‐HA scaffolds.

In this work, the abundant microporous structure on the GBC surface facilitates the adsorption and anchoring of MnFePBA, thereby enabling the material to exert its metal‐like nanozyme activity. Specifically, an increase in temperature enhances the vibration of ─CN─ groups within its crystal lattice, which in turn destabilizes the bound manganese (Mn) ions and iron (Fe) ions [60]. SEM observations confirmed that MnFePBA particles were uniformly embedded within the surface micropores of the GBC(Figure 3d,e). We investigated the use of the photothermal effect of MnFePBA to achieve intelligent and controllable release of Mn ions, which is expected to modulate inflammatory responses. The weakly acidic inflammatory microenvironment impairs the stability of the metal network formed by metal ions, which can also synergistically accelerate the release of ions from the PB lattice (Figure 3i). Therefore, we constructed a bioactive ceramic scaffold with intelligent immunomodulatory functions using MnFePBA&GBC, aiming to realize photothermally and pH‐responsive Mn ion release under the weakly acidic inflammatory microenvironment. We verified the temperature variation of MnFePBA&GBC scaffolds with different loading concentrations under irradiation at various power densities for 10 min. A positive correlation was observed between the temperature elevation and nanoparticle concentration, indicating that the photothermal effect of the scaffold was enhanced with increasing MnFePBA content. We then measured the ion release amounts of two groups of 2.0 MnFePBA&GBC (the concentration with the highest temperature rise) soaked in PBS: one group was exposed to 808 nm NIR irradiation at 1.01 W·cm⁻² for 10 min on the first day (LASER ON group), while the other group was left untreated (LASER OFF group). With the increase in MnFePBA concentration, the cumulative release amounts of Mn ions and Fe ions gradually increased and reached saturation on the third day. Notably, the Mn ion release in the LASER ON group was significantly higher than that in the LASER OFF group. To further confirm that the photothermal effect promotes the release efficiency of Mn ions, we also measured the Mn ion release of scaffolds loaded with double the amount of MnFePBA (4.0 MnFePBA&GBC) without NIR irradiation. Before release saturation (Day 1, Day 3), the Mn²⁺ release from 4.0 MnFePBA&GBC was approximately twice that from 2.0 MnFePBA&GBC. Meanwhile, the Mn²⁺ release level of the NIR‐irradiated 2.0 MnFePBA&GBC group was significantly higher than 1.5 times that of the non‐irradiated 4.0 MnFePBA&GBC group (i.e., 0.5 times higher than the latter). Thus, NIR irradiation can serve as a “switch” for regulating ion release. Specifically, the cumulative Mn ion release increased from 1.3 to 1.8 mg on the third day (Figure 3f–h). These results demonstrate that loading MnFePBA nanoparticles with photothermal conversion capability provides an effective strategy for endowing the scaffold with intelligent ion release functionality, and there is an obvious stimulus‐response relationship between the photothermal effect and ion release efficiency.

2.4. In Vitro Biocompatibility Evaluation of MnFePBA&GBC

To determine the optimal MnFePBA concentration in MnFePBA&GBC scaffolds, in vitro biocompatibility was evaluated using mouse Embryonic Preosteoblast Cells (MC3T3‐E1) (Figure 3j). Cell proliferation on different scaffolds was assessed via the Cell Counting Kit‐8 (CCK‐8) assay. At all detection time points, the optical density (OD) values of scaffolds loaded with 0.25, 0.5, 1.0, 2.0, and 4.0 mg·mL⁻¹ MnFePBA were comparable to those of the blank HA control. However, OD values decreased significantly when MnFePBA concentration exceeded 4.0 mg·mL⁻¹, indicating that MnFePBA loading at ≤ 4.0 mg·mL⁻¹ exhibits no cytotoxicity. Laser scanning confocal microscopy (CLSM) further compared MC3T3‐E1 proliferation on 0.25, 0.5, 1.0, and 2.0 mg·mL⁻¹ MnFePBA&GBC scaffolds versus pure HA One day post‐seeding, cells remained sparsely distributed and spherical without spreading. By day 3, cells had fully spread, revealing the scaffold morphology, with notably higher cell densities observed on 2.0 and 0.5 mg·mL⁻¹ MnFePBA&GBC. By day 7, cell numbers increased significantly with uniform distribution across the scaffold surface, accompanied by distinct cell pseudopodia and minimal cell death. Among all groups, 2.0 and 0.5 mg·mL⁻¹ MnFePBA&GBC exhibited higher cell proliferation levels. To characterize cell distribution on 2.0 mg·mL⁻¹ MnFePBA&GBC scaffolds, CLSM imaging was performed using rhodamine‐phalloidin (red) for cytoskeletal labeling and anti‐vinculin antibodies (green) for vinculin visualization (Figure 3l). Bone marrow stromal cells (BMSCs) displayed normal spreading patterns on 2.0 mg·mL⁻¹ MnFePBA&GBC, with broad spreading and reduced protrusions. This behavior is attributed to the modified surface topography conferred by the external nHAP layer, which increases surface roughness and potentially promotes BMSCs' spreading and osteogenic activity [61].

2.5. In Vitro Nanoenzymatic Activity of MnFePBA&GBC for ROS Scavenging and Immunomodulation

Superoxide anion inhibition rates were quantified to assess SOD‐mimetic activity (Figure 4b). The SOD‐mimetic activity of MnFePBA&GBC exhibited a positive correlation with its loading concentration. Furthermore, NIR irradiation significantly enhanced this activity at equivalent concentrations: within the biocompatible concentration range, the inhibition rate of MnFePBA&GBC increased from ∼30% to over 60%. At the same concentration, NIR‐mediated controlled ion release from the scaffold further improved the inhibition rate by approximately 10% [62]. To verify hydroxyl radical (•OH)‐scavenging capacity, we measured relative •OH levels in samples incubated with HA and x MnFePBA&GBC (x = 0.5/2.0) (Figure 4c) [63]. •OH was generated via the Fenton reaction and detected using terephthalic acid. After 2 h of incubation, •OH levels in MnFePBA&GBC groups were significantly lower than in the HA group, confirming effective scavenging of both hydrogen peroxide (H₂O₂) and •OH by MnFePBA&GBC. TEM was used to visualize organelles and cellular ultrastructure in lipopolysaccharide (LPS)‐stimulated model cells (with excessive ROS production) following MnFePBA phagocytosis (Figure 4d) [64]. No overt disruption or damage was observed in the endoplasmic reticulum or mitochondria, with all organelles maintaining normal morphology. These findings confirm that MnFePBA possesses robust ROS‐scavenging activity. Subsequently, we evaluated intracellular ROS dynamics using the fluorescent probe 2′,7′‐dichlorodihydrofluorescein diacetate (DCFH‐DA), with confocal CLSM utilized to visualize ROS scavenging efficacy in vitro (Figure 4e) [65]. Compared to the HA group, both the 2.0 MnFePBA&GBC and 0.5 MnFePBA&GBC groups displayed markedly reduced green fluorescence intensity. Notably, upon NIR irradiation, the fluorescence signals were further attenuated in the 2.0 MnFePBA&GBC ON and 0.5 MnFePBA&GBC ON groups. Leveraging the unique photothermal properties of MnFePBA, NIR excitation triggers enhanced manganese (Mn) ion release—this modulates the acute inflammatory phase post‐implantation by driving macrophage polarization from the pro‐inflammatory M1 to anti‐inflammatory M2 phenotype, thereby potentiating ROS scavenging (Figure 4k).

FIGURE 4.

In vitro Immunomodulatory Mechanism of MnFePBA&GBC. (a) Schematic illustration of the nanoenzymatic activity of MnFePBA&GBC. MnFePBA at different concentrations under NIR ON/OFF conditions: (b) Superoxide dismutase‐mimetic catalytic activity (n = 3). (c) Hydroxyl radical‐scavenging capacity (n = 3). (d) Ultrastructural images of cells uptake. (e) CLSM images of in vitro ROS scavenging efficacy. (f) Flow cytometry analysis plots. transcriptome analysis of BMSCs. (g) Volcano plot of differentially expressed genes (DEGs). Gray dots indicate genes with no significant difference in expression, purple dots represent significantly upregulated genes, and green dots represent significantly downregulated genes. (h) Clustered heatmaps of DEGs. (i and j) Circle plot of GO results based on 20 significantly downregulated enriched pathways and 10 typical immune‐related enriched pathways in BP, CC, and MF processes. (k) Schematic diagram of the regulatory effect of MnFePBA&GBC on RAW 264.7.

To quantitatively validate MnFePBA&GBC‐induced macrophage polarization, flow cytometric analysis was performed on RAW 264.7 cells stained with anti‐CD86 (M1‐specific marker) and anti‐CD206 (M2‐specific marker) antibodies (Figure 4f). In the LPS‐stimulated positive control (M1‐polarized), CD86+ and CD206+ cells accounted for 53.8% and 44.5%, respectively, while the HA group showed comparable proportions (51.4% CD86+, 46.1% CD206+). Following treatment with 0.5 MnFePBA&GBC or 2.0 MnFePBA&GBC, the cell population exhibited a distinct M1‐to‐M2 phenotypic shift: the 0.5 MnFePBA&GBC group showed 49.9% CD86+ cells and a significant increase in CD206+ cells to 49.3%, whereas the 2.0 MnFePBA&GBC group displayed 51.6% CD86+ and 47.3% CD206+ cells. These results confirm that MnFePBA&GBC effectively mitigates M1 macrophage accumulation in LPS‐induced inflammatory models. We further explored the NIR‐modulated, spatiotemporally controllable immunoregulatory capacity of MnFePBA&GBC. After NIR‐mediated controlled ion release, M2 macrophage proportions were further augmented in irradiated groups: the 0.5 MnFePBA&GBC ON group showed a reduction in CD86+ cells to 47.1% and a notable increase in CD206+ cells to 52%. These findings demonstrate that NIR irradiation reinforces M2 polarization, facilitating efficient inflammation resolution.

Transcriptomic profiling was conducted on bone marrow stromal cells (BMSCs) co‐cultured with MnFePBA&GBC or HA for 14 days. Compared to the HA group, the NIR‐irradiated 2.0 MnFePBA&GBC group exhibited 368 differentially expressed genes (DEGs), with 288 significantly downregulated and 80 significantly upregulated (Figure 4g). All tested metrics showed statistically significant differences between the HA and 2.0 MnFePBA&GBC groups (Figure 4h). Gene Ontology (GO) enrichment analysis revealed that NIR/photothermally triggered ion release from 2.0 MnFePBA&GBC primarily impacted cell surface receptor signaling pathways, immune response processes (biological process [BP]), signal receptor binding (molecular function [MF]), and membrane‐associated cellular components (cellular component [CC]). Specifically, the 2.0 MnFePBA&GBC group significantly suppressed cytokine‐mediated immune responses (Figure 4i) while enhancing cytokine‐cytokine receptor interactions (Figure 4j). Collectively, these data confirm that 2.0 MnFePBA&GBC implantation effectively mitigates implant‐associated immune inflammation by modulating key immune response pathways.

2.6. In Vitro Osteogenic Performance Evaluation of MnFePBA&GBC

To investigate whether the mechanical enhancement strategy of MnFePBA&GBC scaffolds impacts their osteoinductivity, bone BMSCs were employed to assess the osteogenic performance of the scaffolds (Figure 5a). BMSCs were co‐cultured with scaffolds of varying formulations, including 0.5 MnFePBA&GBC ON, 0.5 MnFePBA&GBC OFF, 2.0 MnFePBA&GBC ON, 2.0 MnFePBA&GBC OFF, and pure hydroxyapatite (HA). SEM observations revealed that BMSCs fully spread and exhibited robust adhesion on the surface of all scaffold groups (Figure S10), confirming the good biocompatibility of the scaffolds. Subsequently, BMSCs were seeded onto the scaffold surface and subjected to osteogenic induction to evaluate the influence of the outer active layer on osteogenic differentiation. Alkaline phosphatase (ALP) activity, a key marker for early‐to‐middle stage osteogenic differentiation, was quantified on days 7 and 14 (Figure 5b). All MnFePBA&GBC scaffold groups demonstrated significantly higher ALP activity compared to the HAP control group. Moreover, a positive dose‐dependent relationship was observed between MnFePBA content and ALP activity, with higher MnFePBA loading resulting in enhanced ALP expression. To further assess the middle‐to‐late stage osteogenic differentiation of BMSCs, Alizarin Red S staining was performed to visualize calcium nodule formation (Figure 5c). On day 7 of co‐culture, minimal calcium nodule formation was detected across all groups due to insufficient differentiation time. However, after 14 days of induction, prominent red‐stained calcium nodules were observed in all scaffold groups. Notably, the 0.5 MnFePBA&GBC and 2.0 MnFePBA&GBC groups exhibited a greater number of calcium nodules distributed over a larger area compared to the HA group. This finding suggests that MnFePBA&GBC scaffolds possess superior osteogenic‐inducing capacity, which may be attributed to the key factors: the outer nano‐hydroxyapatite (nHAP) layer compensates for the reduced inherent bioactivity caused by double sintering [66, 67]. Immunofluorescence staining was further utilized to analyze the expression of osteogenic‐related proteins, including bone morphogenetic protein‐2 (BMP‐2) and type I collagen (COL‐1) (Figure 5d,e). BMP‐2, a member of the transforming growth factor‐β (TGF‐β) superfamily, functions as a critical osteogenic inducer and essential transcription factor for osteogenic differentiation [68]. At day 14, the fluorescence intensity of BMP‐2 in MnFePBA&GBC groups were significantly higher than that in the HAP group, with a gradual increase observed as MnFePBA concentration increased. COL‐1, a major extracellular matrix component synthesized and secreted by osteoblasts, exhibited a similar expression pattern to BMP‐2. The addition of MnFePBA significantly upregulated COL‐1 expression, which may be partially attributed to the catalase‐like activity of MnFePBA, which scavenges H₂O₂ and mitigates its inhibitory effect on osteogenesis [69]. To explore the underlying molecular mechanisms, transcriptome analysis was performed on BMSCs cultured for 14 days. Gene Ontology (GO) enrichment analysis was conducted on upregulated genes (UPs) from three categories: molecular function, biological process, and cellular component (Figure 5f,g). Compared with the HA group, five key pathways related to osteogenic calcification and calcium ion deposition were significantly enriched in MnFePBA&GBC groups. This result further confirms that the incorporation of MnFePBA (via ion channel doping) and the porous nHAP surface synergistically promote BMSC osteogenic differentiation.

FIGURE 5.

Osteogenic Differentiation‐Inducing Effect of MnFePBA&GBC on BMSCs. (a) Schematic diagram showing the design in vitro. (b) ARS and (c) ALP staining of BMSCs in different treatment groups after 7 and 14 days of stimulation with osteogenic medium. (d) IF staining of BMP‐2 and COL‐1 expression of BMSCs stimulated by osteogenic medium for 14 days with different treatments(blue represents the cell nucleus, red represents the cytoskeleton, and green represents the positive expression of BMP‐2 and COL‐1). (e) Statistical graph of expression levels quantified by ImageJ. (f) transcriptomic analysis of BMSCs treated with 2.0 MnFePBA&GBC and HA for 14 Days (n = 3) (^* p ≤ 0.05, ^** p ≤ 0.01, ^*** p ≤ 0.01). (g) Based on 20 significantly downregulated enriched pathways in BP, CC, and MF processes.

It is important to note that in all aforementioned osteogenic performance assays, cells were co‐cultured with scaffold extracts rather than in direct contact with the scaffolds themselves. This experimental design indicates that the trace ion release from MnFePBA&GBC scaffolds (rather than direct cell‐scaffold interaction) is the primary mechanism driving enhanced osteogenic differentiation.

2.7. Osteoregenerative Efficacy of MnFePBA&GBC in a Rat In Vivo Calvarial Defect Model

To evaluate the immunomodulatory and osteogenic potential of MnFePBA&GBC, a rat model with critical‐sized calvarial defects was established to assess the bone regeneration efficacy of the MnFePBA&GBC scaffold (Figure S11). This experimental design was based on the unique properties of MnFePBA&GBC, including its ability to release manganese (Mn) and iron (Fe) ions, and additional calcium (Ca) ions from the outer porous nano‐layer. Briefly, 30 8‐week‐old Sprague‐Dawley (SD) rats were anesthetized, and two bilateral critical‐sized defects (φ = 5 mm) were created on the calvaria. HA scaffolds and 2.0 MnFePBA&GBC scaffolds (φ = 5 mm) were randomly implanted into the left and right defect sites, respectively, followed by NIR irradiation of the implantation regions. Rats were euthanized at 1, 4, and 8 weeks postoperatively, and the implanted scaffolds, along with adjacent tissues, were harvested for subsequent analyses (Figure S10).

Macroscopic examination at 4 weeks post‐implantation indicated significant new tissue formation on both the 2.0 MnFePBA&GBC and pure HA scaffolds. Compared to the HA group, the 2.0 MnFePBA&GBC were fully encapsulated by the newly formed tissue, rendering the original scaffold structure nearly indistinguishable. Representative images documenting the surgical implantation and the post‐repair outcomes are available in the Supporting Information (Figure S12). Micro‐computed tomography (Micro‐CT) was employed to assess calvarial defect reconstruction following implantation of different scaffolds (Figure 6a). Initial observations revealed that the volume of new bone (NB, red arrows) in the dual‐enhanced 2.0 MnFePBA&GBC and 0.5 MnFePBA&GBC groups was significantly greater than that in the HA group and blank control group. Quantitative analysis of bone volume fraction (BV/TV) demonstrated that the 2.0 MnFePBA&GBC group (52.42967% ± 0.22443% and 57.74375% ± 1.62081%) and 0.5 MnFePBA&GBC group (50.79667% ± 1.18068% and 51.46263% ± 2.38936%) exhibited significantly higher values at both time points compared to the HA group (31.804% ± 5.86453% and 40.65154% ± 2.26026%) and blank control group (5.93633% ± 2.45333% and 12.71024% ± 0.55983%). Meanwhile, the bone surface area/bone volume ratio (BS/BV) of MnFePBA&GBC showed a decreasing trend. At 4 and 8 weeks, the BS/BV of 2.0 MnFePBA&GBC was significantly lower than that of the HA group. The new bone mineral density (BMD‐BV) increased with the increase in trabecular thickness (Tb.Th) and trabecular connectivity density (Tb.Conn. D), as well as the decrease in trabecular separation (Tb.Sp). These results indicate that the newly formed bone tissue in the MnFePBA&GBC group was more fully developed, with thicker trabeculae, and showed a gradual maturation process. (Figure 6b–h) 3D reconstruction images further corroborated these findings, showing that the MnFePBA&GBC groups displayed more robust new bone ingrowth and a denser overall structure. Notably, one month post‐implantation, a distinct boundary was evident between the scaffold and host tissue, with residual gaps observed between portions of the newly formed bone and the scaffold. By week 8, new bone formation had increased substantially, the interfacial gaps had been eliminated, and the newly formed bone exhibited a more compact and mature morphology.

FIGURE 6.

In vivo bone regeneration enhanced by MnFePBA&GBC in rat cranial defect model. (a) Computed Tomography (CT) evaluation of bone tissue regeneration in different materials, including 3D reconstructions, single‐layer scanning images, and implant images after repair (S: Scaffold, NB: New Bone). (b–h) Micro CT–rendered parameters of ingrown bone at different time points (n = 4) (^* p ≤ 0.05, ^** p ≤ 0.01, ^*** p ≤ 0.01). (i) Hematoxylin and Eosin (H&E) and (j) Masson's trichrome staining at 4 Weeks (Upper Panels) and 8 Weeks (Lower Panels) post‐implantation.

Hematoxylin and eosin (H&E) staining and Masson's trichrome staining were performed on paraffin‐embedded tissue sections to evaluate the repair efficacy of the tissue‐engineered scaffolds (Figure 6i,j). The groups are arranged from left to right as follows: 2.0 MnFePBA&GBC, 0.5 MnFePBA&GBC, HA, and blank control. Soft tissue infiltration was observed within the porous architecture of all three scaffold groups. Compared to the HA and blank control groups, the MnFePBA&GBC groups showed a significantly larger area of peri‐implant new bone formation (red arrows). In both H&E and Masson's staining, bone tissue was observed to grow along the inner walls of the scaffold pores. At week 4 post‐implantation, all scaffold groups exhibited bone formation with a thickness exceeding 20 µm, while the GBC groups further displayed sheet‐like bone formation exceeding 200 µm in thickness. Importantly, new bone formation increased in a MnFePBA concentration‐dependent manner, with extensive large‐area new bone formation observed in the higher‐concentration groups. At two months post‐implantation, histological analysis revealed that all implant groups exhibited sheet‐like bone formation exceeding 100 µm in thickness, with the 2.0 MnFePBA&GBC and 0.5 MnFePBA&GBC groups showing significantly greater bone formation volume than the HA group.

2.8. In Vivo Osteogenic Mechanism and Anti‐Inflammatory Regulatory Effect of MnFePBA&GBC

The expression of osteogenic markers (OPN, BMP‐2, VEGF) was further analyzed by immunofluorescence. (Figure 7b) The results demonstrated notably higher signal intensity in the MnFePBA&GBC groups than in the HA group at every time point, unequivocally confirming their enhanced osteogenic potential. Consistent with in vivo observations, the osteogenic efficacy of MnFePBA&GBC showed a positive correlation with MnFePBA loading concentration, with the 2.0 MnFePBA&GBC group exhibiting the strongest fluorescence intensity across all markers. This finding supports the hypothesis that the enhanced new bone formation in the MnFePBA&GBC groups is attributed to the sustained release of Mn ions and the underlying immunomodulatory mechanism.

FIGURE 7.

In Vivo Regulatory Effects of MnFePBA&GBC on Immunity and Macrophage Polarization in a Rat Calvarial Defect Model. (a) Schematic illustration of the in vivo bone regeneration assessments. (b) IF staining of OPN (red), BMP‐2 (green), and VEGF (yellow) from tissue sections of different scaffolds. (c) IF Staining of CD206 (red) and iNOS (red). (d) Tissue photographs of 2.0 MnFePBA&HA and HA implants in calvarial defects at 4 weeks post‐In vvo implantation. (e,f) qPCR analysis showing the gene expression levels of CD206 and iNOS, and (g) Western Blot (WB) analysis of CD206, iNOS, and CCR7 in scaffolds and their surrounding tissues 1 week after in vivo implantation. (n = 3) (^* p ≤ 0.05, ^** p ≤ 0.01, ^*** p ≤ 0.01).

To assess the impact of the MnFePBA&GBC scaffold on macrophage polarization within the early inflammatory microenvironment in vivo, a comparative analysis was conducted between the 2.0 MnFePBA&GBC and HA groups one week post‐implantation. (Figure 7a) At this stage, considered a crucial window of inflammation, a propitious immune response was a prerequisite for the subsequent replacement of the fibrous tissue by mature bone. Immunofluorescence staining was employed to detect the expression of iNOS (a canonical M1 pro‐inflammatory marker) and CD206 (a classic M2 anti‐inflammatory marker) in the harvested specimens (Figure 7c) [70, 71, 72]. Following NIR irradiation, the 2.0 MnFePBA&GBC group exhibited a marked upregulation of CD206 expression in and around the scaffold, accompanied by a significant downregulation of iNOS expression. From the appearance of the retrieved scaffolds, the interior was fully filled with translucent fibrous connective tissue, while the scaffold morphology remained clearly visible (Figure 7d). These findings preliminarily confirm that MnFePBA&GBC exerts effective immunomodulatory activity in the early post‐implantation phase. Subsequently, quantitative polymerase chain reaction (qPCR) (Figure 7e,f) and Western blot (WB) (Figure 7g and Figure S13) analyses confirmed that the sustained release of Mn ions from MnFePBA&GBC significantly upregulated the M2 macrophage marker CD206 compared to the HA control. Conversely, the expression levels of M1 macrophage markers (iNOS and CCR7) were significantly downregulated in the 2.0 MnFePBA&GBC group compared to the HA group [73]. Collectively, these results confirm that the MnFePBA coupled with the outer layer design synergistically enhances bone defect repair efficacy. These results demonstrate that the synergistic integration of MnFePBA&GBC's immunomodulatory properties, antioxidant activity, and osteoinductivity effectively promotes regeneration at bone defect sites [74].

3. Discussion

In summary, we synthesized MnFePBA nanoparticles via a highly efficient method and immobilized them onto a double‐layered HA substrate, constructing an intelligent‐responsive, functionalized dual‐enhanced MnFePBA&GBC integrated bone repair platform. This platform exhibits multifunctional bioactive properties, including stimuli‐responsive release, non‐invasiveness, immunomodulation, and antioxidation, which enable it to dynamically match the body's spatiotemporal immunomodulatory demands in vivo. Furthermore, it synergistically cooperates with the superior osteogenic performance of the dual‐enhanced HA scaffold to facilitate the regeneration of bone defect sites. Specifically, our design concept is anchored in three core strategies: 1) Immobilization of MnFePBA with nanozyme activity and NIR responsiveness allows for the controlled release of optimal doses of manganese (Mn) ions, which not only exert potent immunomodulatory effects but also scavenge excessive ROS in the implantation microenvironment, thereby establishing a favorable niche for the differentiation of BMSCs. 2) The double‐layered architecture provides sustained additional calcium (Ca) ion release, further accelerating the bone regeneration cascade. 3) Double sintering of the CaP matrix, an optimized sintering protocol, renders the inner layer highly dense, significantly enhancing the mechanical robustness of the platform. Collectively, this platform represents a promising avenue for the development of advanced functionalized bone implants.

Materials characterization confirmed the successful synthesis of MnFePBA nanoparticles with stable micromorphology and crystalline structure. These nanoparticles exhibited superior NIR‐triggered controlled ion release behavior, coupled with SOD‐like and CAT‐like activities. Post‐implantation, they enabled NIR‐mediated adaptive ROS scavenging, effectively preventing ROS‐induced inhibition of osteogenic differentiation and mitochondrial damage. Concomitantly, the controlled release of Mn ions efficiently activated Mn‐SOD enzymes and suppressed the secretion of inflammasome‐related factors, significantly protecting the osteogenic process from ROS‐mediated damage and inflammatory stimuli, thereby promoting calcium deposition and further facilitating osteogenic differentiation. Leveraging the NIR sensitivity of MnFePBA, its immobilization onto the HA substrate enabled precise catalytic regulation of ROS and targeted immunomodulation at the implantation site. Subsequent characterization of the cross‐sectional and surface morphologies of the hierarchical integrated platform validated the scientific rationale of the design and the structural stability of the platform. Our experimental results and mechanistic analyses highlight the excellent biocompatibility and physiological stability of MnFePBA&GBC, as well as its capacity for phase‐specific regulation of inflammatory responses to enhance bone regeneration therapy. Additionally, the outer layer design not only compensates for sintering‐induced bioactivity loss via additional Ca ion release but also incorporates nano‐sized HA, which significantly increases the surface roughness and specific surface area of the integrated platform. This nanostructured design exhibits high biocompatibility, perfectly mimics the nanostructure of natural bone, and establishes a microenvironment conducive to cell infiltration and adhesion, thereby further promoting osteogenic differentiation. The supplementary Ca ion release further upregulates the expression of osteogenic markers (e.g., BMP‐2, COL‐1, OPN), accelerating the bone regeneration process. The development of this integrated implant scaffold—endowed with high mechanical strength, excellent biocompatibility, nanozyme activity, and immunomodulatory capabilities—achieves multi‐aspect synergy, offering a novel strategy for on‐demand regulation and accelerated repair of bone defect sites.

Future research will focus on long‐term in vivo biocompatibility assessment of MnFePBA&GBC and its application in diverse preclinical bone disease models. This study utilized a conventional rat cranial defect model to evaluate the bone repair efficacy of the integrated scaffold platform. Although the rat model is well recognized as a classic preclinical small‐animal model for studying intramembranous ossification and can reliably reflect the in vivo osteogenic behavior of scaffolds, it still has inherent physiological differences from humans. Our long‐term research aim is to further validate its biological performance in large animal bone defect models. Accordingly, future work will adopt large animal models, including sheep and beagles, whose skeletal structure and physiological characteristics closely resemble those of humans, to further assess the mechanical stability, long‐term biosafety, and clinical translation potential of this gradient scaffold. In addition, the present study only characterized the compressive strength of the scaffolds, which represents a major limitation of this work. For the repair of large segmental bone defects, particularly those of load‐bearing bones (e.g., femur and tibia), scaffolds are required to exhibit not only favorable compressive strength but also excellent tensile strength, flexural strength, fatigue resistance, and dynamic mechanical properties. The mechanical data obtained herein merely reflect the static compressive performance of the scaffolds and thus cannot comprehensively assess their mechanical behavior under the dynamic loading conditions of the in vivo physiological microenvironment. Furthermore, this study only determined the in vitro mechanical properties of the scaffolds; their mechanical performance post in vivo implantation (e.g., degradation rate and mechanical stability over time) remains to be further investigated. As indicated in the Discussion section, these aforementioned research directions will be the focus of our subsequent work, and the relevant findings will be reported in future manuscripts. Furthermore, owing to the unique photothermal effect of MnFePBA&GBC, whereby it absorbs NIR radiation to generate heat, we aim to explore its potential application in photothermal therapy (PTT) for bone tumor microenvironments (TMEs). We hypothesize that the platform's unique ROS‐scavenging and immunomodulatory properties are well‐suited for the repair of tumor‐associated bone defects. Additionally, the ─CN─ bonds in MnFePBA exhibit poor stability in acidic environments; thus, the acidic microenvironment of bone tumors can be exploited to achieve environment‐responsive release of MnFePBA&GBC, while addressing the limitation of poor tissue penetration associated with light‐driven responsiveness. Therefore, future studies are warranted to validate the application potential of MnFePBA&GBC in more complex clinical scenarios.

4. Conclusion

In this study, we provide a promising strategy for the development of NIR‐responsive, intelligently regulated, integrated dual‐enhanced implant materials capable of modulating the tissue immune microenvironment. It also offers a solution to the long‐standing bottleneck of simultaneously enhancing the bioactivity and mechanical properties of 3D‐printed CaP ceramics, providing valuable insights for the clinical translation of CaP‐based ceramic implants.

5. Experimental Section

5.1. Materials

Hydroxyapatite powder (HAp) and nano hydroxyapatite (nHA) slurry were purchased from Sichuan Baiamon Bioactive Materials Co., Ltd. (Sichuan, China). K₃[Fe(CN)₆]₃·H₂O, Mn(CH₃COO)₂·4H₂O, calcium sulfate dihydrate (CaSO₄·2H₂O) and 2′,7′‐dichlorodihydrofluoresceindiacetate (DCFH‐DA) were purchased from Aladdin Industrial Co.(Shanghai, China). Lipopolysaccharide (LPS) was purchased from Sigma–Aldrich Chemical Co. (USA). Polyvinylpyrrolidone (PVP) was purchased from Beijing Solarbio Science & Technology Co., Ltd. (Beijing, China).

5.2. Preparation of Manganese‐Iron Prussian Blue Nanoparticles

First, 1 g of polyvinylpyrrolidone (PVP) was dispersed in 27 mL ultrapure water and stirred violently until clear. Then, HCl (270 µL, 1 mol L⁻¹) solution, K₃[Fe(CN)₆] (0.264 g), and Mn(CH₃COO)₂·4H₂O (0.015, 0.0225, and 0.003 g) were added in turn, and then placed in an oil bath at 80°C for 20 h. After that, the round‐bottom flask was aged at room temperature for 2 h, and centrifuged. They were washed (three times each with absolute ethanol and ultrapure water), and the MnFePBA NPs were collected.

5.3. Preparation of Internal HA Scaffolds

The slurry used for Photopolymerization Printing was composed of premixed resin and HA powder, in which the HA content is 60%. The premixed resin was thoroughly mixed with 50 g of polyurethane acrylate, 15 g of polyethylene glycol diacrylate, and 1 g of BYK‐2155; the aforementioned mixture was added to a planetary ball mill and milled for 20 h at a speed of 250 rpm. Subsequently, 20% of CaSO₄·2H₂O was added to the milled slurry, and the mill continued for 2 h at a speed of 250 rpm. 2 g of TPO was finally added to the slurry and milled overnight. With the slurry obtained by the method just and the printed scaffold, the excess slurry was cleaned by ultrasonic and then put into the muffle furnace for gradient sintering, which followed a temperature profile: ramping at 5°C min⁻¹ to 300°C and holding for 6 h, further ramping at 5°C min⁻¹ to 600°C and holding for 2 h, subsequent ramping at 5°C min⁻¹ to 900°C and holding for 10 h, and finally ramping at 5°C min⁻¹ to 1150°C and holding for 2 h, followed by natural cooling to room temperature.

5.4. Preparation of Outer nHA Layer

The purchased nHA slurry was measured to have a concentration of 0.75 g·mL⁻¹. It was diluted 0.5‑ and 1‑fold with ethanol and ethylene glycol, respectively, to prepare nHAP suspensions of 0.5 and 0.375 g·mL⁻¹ for comparison. The scaffolds were immersed in each suspension for 20 min, then dried in an oven at 60°C for 30 min. The above immersion‐drying procedure was repeated 1, 3, 6, and 9 times, respectively. The scaffolds were again placed in a muffle furnace for the second sintering, following a temperature profile: heating to 1100°C and holding for 2 h, then letting cool naturally to room temperature.

5.5. Preparation of MnFePBA&GBC

The dispersed PBA was sonicated and mixed with the prepared nHA slurry at a concentration ratio of 0.5/1.0/2.0/4.0 mg mL⁻¹, and then the twice sintered HA scaffolds were immersed for 30 min, during which stirring was maintained to prevent nHA agglomeration. The resulting scaffolds were placed in an oven to dry at 40°C for 30 to 60 min to obtain the MnFePBA&GBC samples.

5.6. Materials Microstructure

The microstructure of the samples was observed and photographed via scanning electron microscopy (SEM; JSE5900LV, Japan). All ceramic samples were sputter‐coated with gold for 70–140 s before the analysis. The composition of the samples was analyzed using the EDS extension function of the scanning electron microscope. A transmission electron microscope (TEM, Tecnai G2 F20 S‐TWIN, America) was used to observe the internal morphology of MnFePBA.

5.7. Manganese (Mn) Ion Release

MnFePBA&GBC scaffolds of the same volume and porosity were placed in 5 mL Tris‐HCl buffer solution (0.1 m, pH 7.4) to create a system. Those solution systems were then incubated at 37°C in a thermostatic shaker and incubated at various time points (24, 72, and 168 h), and 2 mL of the supernatant was taken from each interval and supplemented with an equal amount of fresh solution. Next, inductively coupled plasma atomic emission spectrometry (ICP‐AES) was used to determine the concentration of manganese (Mn) ions and iron (Fe) ions, and strontium ions released from MnFePBA&GBC in vitro.

5.8. Compression Test

The compressive strength of the ceramic samples was analyzed using a universal testing machine according to the standard ISO13314:2011 (E). The test samples were cylinders with dimensions of φ 6.5 × 13 mm that were compressed at a rate of 1 mm min⁻¹ along the axis of the cylinder until failure. Each group consisted of at least three parallel samples. Stress–strain curves were calculated based on the load and displacement during the compression process.

5.9. XRD

After the MnFePBA prepared by different manganese (Mn) ion ratios was dispersed, the diffraction peaks of the material and the standard Prussian blue (PB) were analyzed by X‐Ray diffractometer (XRD, Philips X'Pert 1 X‐ray diffractometer, Netherlands). The samples were scanned at a rate of 0.05 ° s⁻¹ in the range of 10°–70° using a Cu target with a voltage of 30 kV and a current of 20 mA.

5.10. Cell Culture

MC3T3‐E1 cells were used for in vitro cytocompatibility experiments, and bone marrow mesenchymal stem cells (BMSCs) isolated from Sprague‐Dawley rats were used for in vitro osteogenesis assay. Both cells were cultured in α‐MEM supplemented with 10% fetal bovine serum and 1% penicillin/streptomycin and then incubated at 37°C in a 5% CO₂ atmosphere. The RAW 264.7 cell line cultured in a high‐glucose DMEM with 10% FBS and 1% penicillin/streptomycin was used in observing the ability of the substance to scavenge intracellular ROS and in an in vitro anti‐inflammatory assay.

5.11. In Vitro Proliferation and Differentiation

After seeding 5 × 10⁴ cells on the sterilized surface of different materials and incubating in an incubator at 37°C for 30 min, the cells were cultured with 1 mL complete medium, and the cytotoxicity and cell proliferation and cell morphology of the materials were respectively detected by CCK‐8 reagent kit and Flu‐orescein diacetate (FDA)/propidium iodide (PI) staining solution at 1, 3, and 7 days. Then, use confocal laser scanning microscopy to collect each picture. 5 × 10⁴ BMSCs were seeded on each material for normal culture, and osteogenic differentiation was carried out with osteogenic induction medium when the number of cells in the well plate reached 80%, and staining was performed with ALP activity kit on days 7 and 14 days after changing to osteogenic induction medium to measure the alkaline phosphatase (ALP) activity. Analogously, the calcification nodules of the cells were observed using Alizarin Red S (ARS) staining.

5.12. Evaluation of the Repolarization of Macrophages

To verify the polarization regulation effect of MnFePBA&GBC at different concentrations on macrophage with or without NIR laser irradiation, 2 × 10⁴ RAW 264.7 cells were seeded in 24‐well plates and cultured for 24 h, then 2 µg mL⁻¹ LPS was added to release inflammatory signals to simulate the inflammatory state of the bone defect. At the same time, different concentrations of MnFePBA&GBC were immersed in complete medium for 24 h, and half of the same concentration of MnFePBA&GBC was irradiated with 808 NIR laser for 10 min at 12 h of immersion. The medium of polarized RAW 264.7 was replaced by the material extract obtained above and cultured for 48 h. After the medium was removed and washed with PBS, the cells were collected, stained with antibodies PE‐CD206 and APC‐CD86, and M1/M2 analysis of the cells was performed by flow cytometry.

5.13. In Vitro ROS Scavenging Capacity

RAW 264.7 cells were cultured in 24‐well plates for 24 h to ensure that the number of cells in each well reached 5 × 10⁵. Then the culture medium supernatant obtained by soaking different materials for 24 h was added to each well plate for 48 h to observe the changes of intracellular ROS, and half of the same concentration was irradiated with 808 nm NIR laser for 10 min at 12 h of soaking. Then, according to the ROS reactive oxygen species detection kit, the cells were labeled with 2′,7′‐dichlorofluorescein diacetate (H2DCFH‐DA) to detect the ROS level, and a confocal laser scanning microscope was used to collect pictures. Similarly, Analysis of the intracellular ROS levels by flow cytometry.

5.14. Establishment of the Skull Defect Model

To further evaluate the immunomodulatory function and bone regeneration‐promoting ability of functionalized reinforced calcium phosphate ceramics in vivo, 30 Sprague Dawley rats were divided into four groups (blank control group, HA group, 0.5 MnFePBA&GBC group, and 2.0 MnFePBA&GBC group, randomized, with 10 parallel samples in each group. After successful anesthesia of the Sprague Dawley rats, the cranial skin was shaved and disinfected with 1% iodophor. The skin, soft tissue, and periosteum along the midline of the skull were incised layer by layer, and a 5 mm diameter bone defect was created on both sides of the skull using a drill. Except for the blank control group, sterilized materials were implanted into the defect sites. The surgical site was sutured and disinfected using surgical thread. 24 h later, the surgical site was irradiated with an 808 nm NIR laser for 10 min.SD rats were euthanized at 1, 4, and 8 weeks post‐implantation, respectively. The calvaria were harvested for radiological and pathological examinations. Micro‐computed tomography (Micro‐CT) scanning was performed using the vivaCT80 imaging system (SCANCO Medical AG, Switzerland). Three‐dimensional (3D) structures were reconstructed with Mimics software, and quantitative analysis of bone mineral density (BMD) and bone volume (BV) was conducted using CTan software.

This experiment was conducted under the approval and guidance of the Animal Management and Use Committee of Sichuan University (Ethics Registration Number: 20230310045).

5.15. Statistical Analysis

All data in this section were analyzed using GraphPad Version 8.0 (GraphPad Prism). Data were expressed as mean ± standard deviation. A Student's t‐test was used to assess differences between two groups, while one/two‐way analysis of variance (ANOVA) was applied for comparisons among three or more groups, with ^* p ≤ 0.05, ^** p ≤ 0.01, ^*** p ≤ 0.001, and ^**** p ≤ 0.0001 indicating statistically significant differences.

Conflicts of Interest

The authors declare no conflicts of interest.

Supporting information

Supporting File: smll73923‐sup‐0001‐SuppMat.docx.

Supporting File: smll73923‐sup‐0002‐DataFile.pdf

Data Availability Statement

The data that support the findings of this study are available from the corresponding author upon reasonable request.