Skip to main content
NCBI home page
ACS Omega logoLink to ACS Omega
. 2025 Aug 25;10(35):39494–39503. doi: 10.1021/acsomega.4c11634

Strategic Uniaxial Macropore Sizing for Honeycomb-Structured Scaffolds for Enhanced Bone Reconstruction in Segmental Bone Defects

Keigo Shibahara †,‡, Koichiro Hayashi †,*, Yasuharu Nakashima ‡, Kunio Ishikawa †
PMCID: PMC12423895  PMID: 40949199

Abstract

Segmental bone defects often present as new irregularly growing bones that fail to bridge gaps. Optimizing the scaffold design for direct bone growth can enhance bone defect reconstruction. We evaluated three carbonate apatite honeycomb-structured scaffolds with uniaxial macropores of 280, 440, and 640 μm (labeled P280, P440, and P640, respectively) implanted into critical-sized ulnar defects in rabbits. The scaffold performance was assessed at 4 and 12 weeks postimplantation. P280 formed small volumes of oriented bone, which decreased the bone strength. P440 formed large volumes of oriented bone and restored bone strength. Meanwhile, P640 formed abundant volumes of disorganized bone and failed to enhance bone strength. The 440 μm macropores effectively guided new bone and enhanced the bone strength. These findings suggest that the macropore size is crucial for designing scaffolds for effective segmental bone defect reconstruction.

graphic file with name ao4c11634_0010.jpg

graphic file with name ao4c11634_0008.jpg

Introduction

Treating segmental defects in long bones is challenging because of deficiencies in cells, nutrition, and mechanical stability. , The early stage regeneration of newly formed bones is restricted to the edges of the bone stumps because bone-related cells are supplied by the periosteum and endosteum. − The newly formed bones then grow irregularly without guidance from the opposite bone stump. Failure to bridge bone defects leads to nonunion. In contrast, successful bone bridging restores mechanical bone stability and promotes bone remodeling. Therefore, one strategy for treating segmental bone defects involves promoting bone formation, minimizing irregular bone growth, and guiding the bone stumps.

Autografts and allografts are clinically used to treat segmental bone defects. Although autografts are the gold standard for bone grafting, harvesting procedures lead to pain and infection at the harvest site. Furthermore, autografts are prone to heterotopic ossification owing to their inability to orient bone growth. Meanwhile, allografts carry the risk of transmissible diseases. Allografts have poor mechanical structure that does not maintain bone growth in the defects. The risks associated with autografts and allografts can be avoided using synthetic bone scaffolds. The scaffold matrix determines the biocompatibility, bioresorbability, and mechanical properties, , while the scaffold structure controls the direction of tissue growth and affects the mechanical strength of the scaffold based on the pore architecture. − Although synthetic bone scaffolds are a promising option for treating segmental bone defects, they require further optimization of the scaffold material and structure because of their lower osteoconductivity than that of autografts and allografts. ,

Calcium phosphates are commonly used in synthetic bone scaffolds because their chemical components are similar to those in humans. − Although the typical calcium phosphates are hydroxyapatite (HAp: Ca10(PO4)6(OH)2) and beta-calcium phosphate (β-TCP: Ca3(PO4)2), their resorbabilities may not be optimal for bone regeneration. − HAp is rarely resorbed, thereby preventing bone formation at the material site. − β-TCP is resorbed spontaneously, potentially causing the scaffold to disappear before full bone formation at the defect site. − Meanwhile, carbonate apatite (CAp: Ca10‑a(PO4)6‑b(CO3)2‑c), another calcium phosphate material, can synchronize the rate of material resorption and bone formation because it is resorbed by osteoclasts. − Furthermore, previous studies have shown that scaffolds composed of CAp exhibited superior osteoconductivities to those of HAp and β-TCP. − Therefore, CAp is suitable for enhancing the osteoconductivity of synthetic bone scaffolds.

Scaffold structure critically affects bone formation and orientation. − A previous study using a segmental bone defect model revealed that a honeycomb (HC)-structured scaffold with square cells that can be regarded as uniaxial macropores promoted bone formation and reconstruction compared to three-dimensional pore-structured scaffolds in segmental bone defects. Uniaxial macropores parallel to the bone axis effectively guide bone ingrowth along the direction of the bone axis while preventing fibrous tissue invasion compared to multidirectional macropores. Furthermore, the uniaxial macropores smaller than 730 μm formed oriented bones and enhanced the strength of scaffolds. Although smaller uniaxial macropores form more oriented bones, excessively small uniaxial macropores may decrease the volume of tissue owing to spatial restrictions. Specifically, the uniaxial macropores smaller than 281 μm inhibit osteogenesis. Thus, a threshold for the uniaxial macropore size to maximize bone orientation and volume should fall between 281 and 730 μm. However, the optimal macropore size remains unclear.

To evaluate the effect of uniaxial macropore size in HC-structured scaffolds, varying the macropore size while keeping the cell wall (i.e., strut) thickness constant is essential. Previously, although not in segmental bone defects, we assessed the impact of macropore size while maintaining a uniform strut thickness. For example, in an evaluation where HC-structured scaffolds were implanted into bone defects in the rabbit femur condyle, the strut thickness was set at approximately 100 μm, while the macropore sizes were varied (100, 200, and 300 μm)35. The results showed that the rate of bone formation was greatest with a macropore size of 300 μm. Furthermore, in another study where HC-structured scaffolds with a uniform strut thickness of approximately 180 μm and macropore sizes of 230, 460, and 630 μm were implanted onto the rabbit calvaria, the 230 μm macropores prevented soft tissue invasion and promoted vertical bone ingrowth. In contrast, macropores of 460 and 630 μm did not prevent soft tissue infiltration, which limited vertical bone ingrowth.

Based on the findings from our previous studies, in this study, we have designed HC-structured scaffolds with an average strut thickness of <180 μm and macropore sizes of 280, 440, and 640 μm (designated as P280, P440, and P640, respectively). These scaffolds were implanted into segmental bone defects in rabbits to evaluate the effects of macropore size on bone formation, including bone volume, orientation, and mechanical strength.

Methods

Fabrication of P280, P440, and P640

P280, P440, and P640 were fabricated using extrusion molding as previously described. − A mixture of CaCO3 powder and a methylcellulose-based binder was prepared for extrusion. An extruder (V-30-II, Universe Co., Ltd., Saga, Japan) pushed out the mixture through dies with a slit thickness of 200 μm and pitch of 500, 700, and 900 μm. For debinding, the HC green bodies were heated up to 650 °C with a rate of 0.03 °C/min and kept at 650 °C for 24 h under CO2 flow. The debinded scaffolds were immersed in a mixture of 1 mol/L Na2HPO4 (Fujifilm Wako Pure Chemical Co., Ltd., Tokyo, Japan, Osaka, Japan) and 0.5 mol/L carbonate ion (NaHCO3 and Na2CO3; Fujifilm Wako Pure Chemical Co., Ltd.) at 80 °C for 1 week. The immersion caused a chemical conversion from CaCO3 to CAp via a dissolution–precipitation reaction. For implantation, the scaffolds were cut into 10 mm × 7 mm × 4 mm pieces using a mechanical polisher (Olympus Terumo Biomaterials Co., Ltd., Tokyo, Japan).

Physicochemical Analyses

X-ray diffraction (XRD) patterns of P280, P440, P640, and commercial CAp (Cytrans Granules, GC Co., Ltd., Tokyo, Japan) were recorded using a diffractometer (D8 Advance, Bruker AXS GmbH, Karlsruhe, Germany). Fourier transform infrared (FTIR) spectra of P280, P440, P640, and commercial HAp (HAP-100, Taihei Chemical Industries Co., Ltd., Nara, Japan) were recorded using a spectrometer (FT/IR-6200; JASCO Co., Ltd., Tokyo, Japan). The carbonate content of P280, P440, and P640 was measured using a CHN coder (MT-6; Yanako Analytical Instruments Co., Ltd., Kyoto, Japan; n = 3 per group). Macroscale (>100 μm) morphologies were evaluated by images from microcomputer tomography (μ-CT; ScanXmate-L080T/L090T, Comscan Techno Co., Ltd., Kanagawa, Japan) and scanning electron microscopy (SEM; S3400N, Hitachi High-Technologies Co., Ltd., Tokyo, Japan). The aperture side lengths and strut sizes were measured using the SEM images (n = 10 per group). The macroscale-specific surface areas of P280, P440, and P640 were determined using the representative volume element (RVE) method. Specifically, the macroscale-specific surface area was obtained by dividing the internal wall surface area by the unit cell volume. Given a unit cell with an average aperture side length of pore L, average strut thickness t, and scaffold height H, the internal wall surface area of one macropore and the unit cell volume were 4LH and (L+t) 2 H, respectively. The macroscale-specific surface area was calculated using the following equation:

Macroscale‐specific surface area⁡(μm−1)=4LH/(L+t)2H=4L/(L+t)2 1

The specific surface area determined using the RVE method was originally expressed in μm– 1. Therefore, it was converted to m2/g using the theoretical density value of HAp (3.16 g/cm3). The theoretical density of HAp was used because the theoretical density of CAp is not constant. However, because the densities of CAp and HAp are generally considered to be similar, this approximation is reasonable.

Microscale (<100 μm) morphologies were evaluated by mercury injection porosimetry (MIP; AutoPore 9420, Shimadzu Co., Ltd., Kyoto, Japan) tests and SEM images. The micro- and nanoscale specific surface area of P280, P440, and P640 were determined using MIP. The total porosities of P280, P440, and P640 were measured using the theoretical density of HAp (n = 8 per group). The mechanical strength of P280, P440, and P640 was evaluated by measuring their compressive strengths and Young’s moduli. A universal testing machine (Autograph AGS-J; Shimadzu Co., Ltd.) was used to load the scaffolds parallel to the pore direction at a crosshead speed of 1 mm/min. The loading forces were recorded when the scaffold fractured (n = 8 per group). The Young’s moduli of P280, P440, and P640 were determined from the linear regions of the stress–strain curves.

Animal Experiments

All the animal experiments were approved by the Animal Care and Use Committee of Kyushu University (approval no. A23–007–0). A total of 24 Japanese white rabbits (18 weeks of age, male, 2.9–3.4 kg, Japan SLC, Co., Ltd., Shizuoka, Japan) were used in this study. For general anesthesia, an intramuscular injection of ketamine (30 mg/kg) and xylazine (5.0 mg/kg) was administered. The forearm fur was cut using a shaver (Akiyama-Seisakusyo Co. Ltd., Tokyo, Japan). Lidocaine (6 mg/kg) was administered subcutaneously to induce local anesthesia. Before the surgery, gentamicin (4 mg/kg) was administered to prevent wound infection. A 10% povidone-iodine liquid was sprayed on the surgical site for skin disinfection (Meiji Seika Pharma Co., Ltd., Tokyo, Japan). The ulna was exposed using a surgical knife and cut into a 10 mm-long defect using a bone saw (sagittal blade; Zimmer Biomet Co., Ltd., Tokyo, Japan) and a 10 mm-long guide block. − ,− A polyethylene terephthalate (PET) sheet was placed on the radius to eliminate its effects on bone regeneration. ,,− After fixing the stainless-steel plate and screws, the scaffold was implanted in the defect and looped onto a plate with 4–0 nylon (MANI Co., Ltd., Tochigi, Japan). The muscles and skin were tightly sutured using 4–0 nylon. Both the forelimbs were used for surgery. The forearms were not immobilized after surgery. An overdose of the anesthetic was used to sacrifice the rabbits at 4 and 12 weeks postimplantation (PI4w and PI12w, respectively). For the radiographic and histomorphometric analyses, the specimens were immersed in a 10% formalin solution to fix the tissues.

Radiographical and Histomorphometric Analyses

Radiographs were obtained using HA-60 (HITEX Co., Ltd., Osaka, Japan). The μ-CT images were obtained using ScanXmate-L080T/L090T (Comscan Techno Co., Ltd., Kanagawa, Japan). The bones and materials in the defects were evaluated using the TRI/3D-BON-FCS64 software (RATOC System Engineering Co., Ltd., Tokyo, Japan). The volumes of newly formed bones and materials were identified by selecting the range of interest (ROI) that most accurately reflected the shape of each object (Figure S1A and S1B). Residual material was calculated by dividing the material volume after implantation by the material volume before implantation. The resorption rate of the material was calculated by subtracting the residual material percentage after each implantation period from a value of 100%, considering the residual material ratio immediately after implantation as 100%. Hematoxylin-eosin (HE)-stained images were prepared to evaluate the morphology of bones and blood vessels at PI4w and PI12w. All specimens were cut along the bone axis. Histological images were obtained using a BZ-X digital analyzer (Keyence Co., Ltd., Osaka, Japan). Twelve Japanese white rabbits were used for these analyses (n = 4 per group at PI4w and PI12w, respectively).

Bone Orientation Analyses

Bone orientation was evaluated using 10-fold magnified H&E-stained images and the ImageJ plug-in FibrilTool as previously indicated. ,, The range of interest squares (ROIs) with a side length of approximately 100 μm and an area of approximately 0.01 mm2 were created in the newly formed bone area. The orientation angles and anisotropic degrees of the newly formed bones were measured using ROIs. The ROI numbers for P280, P440, and P640 were 144, 125, and 134 at PI4w and 110, 109, and 146 at PI12w. The orientation angle of the newly formed bone is the angle relative to the bone axis. The degree of anisotropy of the newly formed bone ranged from 0 (no orientation) to 1 (perfect orientation).

Four-Point Bending Tests

The bones implanted with P280, P440, and P640 for biomechanical analyses were prepared without plate fixation and with scaffold implantation in the same 10 mm segmental defects. Plate removal is generally accompanied by the removal of some parts of new bone, which can affect bone mechanics. At PI12w, both forearm bones were harvested. Prior to the test, the radius of the forearm was cut to eliminate biomechanics. Forearm bones with radius cuts were subjected to four-point bending tests according to a previously reported method. , The setup consisted of a 10 mm distance load point and 30 mm distance supporting points. A universal testing machine (Autograph AGS-J; Shimadzu Co., Ltd.) was used to apply forces perpendicular to the bone axes of the specimens at a crosshead speed of 1 mm/min. The specimens were loaded until fracture. The maximum measured force was considered as the flexural strength of the bones. Twelve white Japanese rabbits were used for these analyses. Three forearm bones implanted with P280, P400, and P640 (one sample from each group) were excluded from the analyses because they were fractured prior to testing. Seven samples were analyzed for each group. The original forearm bone with a radial cut was used as a control (n = 8). The restoration percentage of flexural strength was calculated by dividing the average flexural strength of the radius-cut forearm bone implanted with the scaffolds by that of the forearm bone implanted with a radius cut.

Statistical Analyses

One-way analysis of variance and Tukey’s test were used for statistical analyses, and p < 0.05 indicated statistical significance.

Results

Fabrication and Characterization of P280, P400, and P640

The XRD patterns of P280, P400, and P640 showed apatite diffraction, as indicated by commercial HAp (Figure A). , The FTIR spectrum of P280, P400, and P640 showed phosphate (1027–960 and 605–565 cm–1) and carbonate bands (1476, 1416, and 873 cm–1) but not hydroxyl bands (630 cm–1) (Figure B). , The double carbonate bands at 1476 and 1416 cm–1 in the ν region indicated A- and B-type carbonates, respectively. , These findings indicate that P280, P400, and P640 were composed of AB-type carbonate apatite. The carbonate contents of P280, P400, and P640 were in the range of 10.9%–12.8%.

1.

1

Open in a new tab

(A) XRD patterns and (B) FTIR spectra of P280, P440, and P640. The XRD pattern of CAp and the FTIR spectrum of HAp were used as references.

P280, P400, and P640 possessed uniaxial square macropores (Figure A–C). Moreover, the average aperture side lengths and strut thicknesses increased with increasing macropore size, as shown in Table . Statistically significant differences in strut thickness were observed among the groups (P280:169 ± 3 μm; P440:174 ± 6 μm; P640:179 ± 7 μm; p < 0.05). The macroscale-specific surface areas of P280, P440, and P640 determined using the RVE method were 5.6 × 10–3, 4.7 × 10–3, and 3.8 × 10–3 μm-1, respectively. When converted to different units, these values correspond to 1.8 × 10–3, 1.5 × 10–3, and 1.2 × 10–3 m2/g, respectively. The total porosity of the scaffolds increased with the aperture side lengths of the macropores, while the compressive strengths decreased.

2.

2

Open in a new tab

Macroscale morphological images of P280, P440, and P640: (A–C) 3D μ-CT images and SEM images of the (D–F) pore apertures and (G–I) struts. The white and black double-headed arrows indicate aperture side length of pores and strut thicknesses.

1. Structural Parameters of P280, P440, and P640 ,

Parameters P280 P440 P640
Aperture side length of pores (μm) 278 ± 5 444 ± 7 638 ±8
Strut thickness (μm) 169 ± 3 174 ± 6 179 ± 7
Total porosity (%) 69 ± 1 74 ± 1 80 ± 1
Compressive strength (MPa) 30 ± 6 21 ± 3 14 ± 3
Young’s modulus (GPa) 15 ± 1 14 ± 1 14 ± 1
Open in a new tab
a

Values are presented as the means ± standard deviation.

b

p < 0.05.

High-magnification SEM images of the P280, P400, and P640 showed micropores <10 μm between the CAp crystals (Figure A–C). The MIP test showed that micropores were distributed under 10 μm (Figure D–F). The micropore volumes were 0.23–0.24 cm3/g (Figure G–I). The micro- and nanoscale specific surface areas of P280, P440, and P640 determined from MIP were 1.7, 2.3, and 2.1 m2/g, respectively

3.

3

Open in a new tab

Microscale morphological images of P280, P440, and P640. (A–C) low-magnification SEM images and (D–F) high-magnification SEM images. The yellow arrows indicate the micropores. The yellow dotted outlines indicate micropores. (G–L) mercury intrusion porosimetry tests. (G–I) Pore-size distribution vs pore diameter and (J–L) cumulative pore volume vs pore diameter in the range below 100 μm.

Radiographic and Histomorphometric Evaluation

A one-sided gap between P280 and the bone stump persisted at PI4w (Figure A). The scaffold edge on the gap side exhibited poor osseointegration (Figure B,C). In contrast, P440 and P640 showed osseointegration and bony callus incorporation (Figure D–I). At PI12w, P280, P440, and P640 showed osseointegration between the scaffolds and bone stumps (Figure J, M, and P). Although P280 showed no bony bridging (Figure K,L), P440 and P640 showed bony bridging between the gaps (Figure N, O, Q, and R).

4.

4

Open in a new tab

Radiographic and histomorphometric images of P280, P400, and P640 at (A–I) PI4w and (J–R) PI12w. The left, center, and right columns show the radiographs, cross-sectional μ-CT images, and HE-stained images, respectively. The yellow arrowhead indicates osseointegration between the scaffold edge and the bone stump. The blue arrowhead indicates the remaining gap between the edge of the scaffold and the bone stump. The red regions in the μ-CT images indicate bone formation interior to the scaffold. The red arrowhead indicates bony bridging through the macropores.

At PI4w, the volumes of newly formed bones interior in P440 (25.1 ± 3.5 mm3) and P640 (24.0 ± 4.1 mm3) were significantly higher than that in P280 (15.1 ± 2.0 mm3, p = 0.01 and 0.03, Figure A). At PI12w, the volumes in P440 (75.4 ± 2.3 mm3) and P640 (76.1 ± 5.3 mm3) remained significantly higher than in P280 (63.4 ± 3.4 mm3, p = 0.002 and 0.001, respectively, Figure A). No significant difference was observed in blood vessel thickness within newly formed bones among P280, P440, and P640 at PI4w and PI12w (51.1–54.1 μm at PI4w and 52.1–53.4 μm at PI12, Figures B and S2). Although the material volumes in the defects were different among the scaffolds (Figure C), no significant difference was observed in the residual materials at PI4w and PI12w (Figure D, 87.2%–90.7% at PI4w and 16.3%–21.0% at PI12w). The resorption rates of the materials were similar among P280, P440, and P640 at PI4w and PI12w (Figure D, 4.3%–4.8% at PI4w and 66.7%–73.7% at PI12w).

5.

5

Open in a new tab

Radiographical and histomorphometric findings for P280, P440, and P640 at PI4w and PI12w: (A) volumes of newly formed bones interior of the scaffolds (mm3); (B) blood vessel thicknesses within newly formed bones (μm); (C) material volumes (mm3); (D) residual materials (%); (E) resorption rates of the materials (%). Significant differences are indicated by different characteristics (p < 0.05).

Bone Orientation Evaluation

At PI4w, although most of the newly formed bones within the channels of P280 and P440 were well-oriented along the channel directions, some newly formed bones of P640 were poorly oriented (Figure A–C). Similar trends in the bone orientations of P280, P440, and P640 were observed at PI12w (Figure D–F). At PI4w, the orientation angles of P280 (11.5 ± 2.1°) and P440 (10.3 ± 2.0°) were significantly lower than that of P640 at PI4w (19.4 ± 7.0°, p = 0.001 and 0.0001, Figure G). At PI12w, although the orientation angles of P440 show no statistically significant difference from those of P280 (11.5 ± 2.4°) and P440 (14.3 ± 3.9°) (p = 0.93 and 0.08), that of P280 was significantly lower than that of P640 (23.8 ± 11.6°, p = 0.002, Figure G). The anisotropic degrees of P280 (0.18 ± 0.04) and P440 (0.17 ± 0.02) at PI4w were significantly higher than that of P640 at PI4w (0.13 ± 0.04, p = 0.000003 and 0.004, Figure H). The anisotropic degrees of P280 (0.20 ± 0.02) and P440 (0.20 ± 0.03) at PI12w remained significantly higher than in P640 (0.14 ± 0.03, p = 0.000003 and 0.0000001, respectively, Figure H).

6.

6

Open in a new tab

Newly formed bone images of P280, P440, and P640 at (A–F) PI4w and (D–F) PI12w. The left and right columns show HE-stained images in the channels and bone-oriented images in the channels, respectively. NB and M indicate the newly formed bones and materials, respectively. (G) Orientation angles of newly formed bones within channels and (H) anisotropic degrees of newly formed bones within channels.

Four-Point Bending Tests

The flexural strengths of bone with radius cut and bones implanted with P280, P440, and P640 were 133.5 ± 14.4, 38.8 ± 11.2, 89.3 ± 30.5, 68.6 ± 19.8 N, and, respectively (Figure ). The bones implanted with P440 and P640 showed significantly higher flexural strength than those implanted with P280 (p = 0.0004 and 0.048, respectively). The flexural strength of bones with radial cuts was significantly higher than that of bones implanted with scaffolds (p < 0.001). The restoration percentages of flexural strengths for P280, P440, and P640 were 29.0%, 66.9%, and 51.4%, respectively.

7.

7

Open in a new tab

Four-point bending test of bone with a radius cut and bones implanted with P280, P440, and P640 at PI12w. The left and right Y-axes show the flexural strengths of the bones and restoration percentages of the flexural strengths, respectively. Significant differences are indicated by different characteristics (p < 0.05).

Discussion

Uniaxial 280 μm macropores formed oriented bones, whereas bone volumes were small, resulting in failure to bridge the defects and lower strengths of scaffold-implanted bones. For uniaxial 440 μm macropores, abundant oriented bones were formed and strengths of the scaffold-implanted bones were restored. Lastly, uniaxial 640 μm macropores formed abundant bones whereas the bone orientations were poor, subsequently failing to enhance the strength of scaffold-implanted bones. Thus, the size of uniaxial macropores should be between 280 and 640 μm to maximize both bone volumes and orientations. Consequently, both restorations strengthened the scaffold-implanted long bones. This is the first study to explore the optimal pore size from the perspectives of bone volume, orientation, and strength in segmental bone defects.

Although statistical analysis revealed significant differences in strut thickness among the P280, P440, and P640 scaffolds (p < 0.05), the absolute differences were minimal (less than 10 μm). Given the substantial variation in macropore sizes (278–638 μm), these minor differences in strut thickness were considered negligible. Therefore, strut thickness was regarded as approximately consistent across groups, in accordance with our scaffold design strategy.

The compressive strengths of P280, P440, and P640 scaffolds were 30 ± 6 MPa, 21 ± 3 MPa, and 14 ± 3 MPa, respectively, and decreased with increasing pore size (P280:278 ± 5 μm; P440:444 ± 7 μm; P640:638 ± 8 μm). This trend primarily reflects the corresponding increases in total porosity (P280:69 ± 1%; P440:74 ± 1%; P640:80 ± 1%), in line with previous reports., , Although strut thickness also increased slightly with pore size (P280:169 ± 3 μm; P440:174 ± 6 μm; P640:179 ± 7 μm), the small magnitude of these differences (<10 μm) suggests that porosity was the dominant factor influencing compressive strength. The relatively low porosity of P280 conferred high initial scaffold strength but restricted bone ingrowth, resulting in limited new bone formation and consequently low flexural strength of the reconstructed bone (38.8 ± 11.2 N). In contrast, P440 exhibited moderate porosity and compressive strength, providing sufficient mechanical support while enabling oriented bone formation, thereby achieving the highest flexural strength among the groups (89.3 ± 30.5 N). Although P640 facilitated substantial bone formation due to its high porosity, the reduced initial scaffold strength likely impaired the formation of well-oriented bone, resulting in a lower flexural strength (68.6 ± 19.8 N) compared to P440.

Previous studies have explored the optimal macropore size for bone formation by evaluating differences in bone volume. Li et al. implanted polycaprolactone/polyethylene glycol/hydroxyapatite bioactive scaffolds with different uniform macropore sizes (210, 386, and 582 μm) into critical-sized (5 mm diameter) mandibular bone defects of rats. The 582 μm macropores promoted bone ingrowth more than those of 210 and 386 μm. Qin et al. implanted magnesium-substituted calcium silicate scaffolds with uniform macropore sizes (464, 590, and 703 μm) into critically sized (10 × 6 × 4 mm) mandibular defects in rabbits. The 590 μm macropores formed more new bones than that of 703 μm. Compared to these previous studies, our research uniquely considered not only bone volumes, but also orientations and strengths. We demonstrated how changes in the uniaxial macropore size affect these factors, providing valuable insights for scaffold design. Uniaxial 280 μm macropores impose strong spatial constraints, enhancing bone orientation while reducing bone volume, suggesting a new minimum for uniaxial macropore size in segmental bone defects.

Uniaxial 640 μm macropores disorganized the bones compared to 440 μm macropores. Despite weaker spatial constraints, the 640 μm macropores did not increase bone volumes. Thus, reducing spatial constraints increases the number of disorganized bones, which subsequently obstructs bone ingrowth from the bone stumps. Previous studies using segmental bone defects have shown that the orientation of soft tissue affects the direction of bone ingrowth. , Petersen et al. demonstrated the effect of collagen orientation on the bone ingrowth direction using porcine collagen scaffolds with parallel macropores, both parallel and perpendicular to the bone axis. Parallel macropores supported bone ingrowth in the same direction, whereas macropores both parallel and perpendicular to the bone axis did not support bone ingrowth because of the disorganized soft tissue invasion from the perpendicular macropores. Similarly, in our previous study, we investigated the effects of macropore orientation on the ingrowth directions of bone and fibrous tissue using CAp scaffolds with parallel macropores, both parallel and perpendicular to the bone axis. Parallel macropores oriented the bone and fibrous tissue in the same direction and successfully increased the bone volume. In contrast, macropores, both parallel and perpendicular to the bone axis, allow fibrous tissue to invade perpendicularly, leading to disorganized bones and decreased bone volume. Therefore, the findings of the present and previous studies suggest that the presence of any tissue, even the bone itself, that loses orientation relative to the macropore direction obstructs bone ingrowth in the macropore direction.

Our previous study demonstrated that the segmental bone defect model used in this study does not undergo spontaneous healing without scaffold implantation, even at 4 and 12 weeks postimplantation (PI4w and PI12w, respectively) (Figure S3). Specifically, the volumes of newly formed bone in the defect without scaffold implantation were 2.8 ± 0.4 mm3 at PI4w and 2.9 ± 0.6 mm3 at PI12w (Figure S3G). In contrast, the CAp HC-structured scaffolds (P280, P440, and P640) facilitated bone regeneration volumes that were at least 5.3-fold (PI4w) and 9.0-fold (PI12w) greater than those observed in the no-scaffold group.

Furthermore, a previous study using the same segmental defect model evaluated the performance of a commercially available β-TCP 3D porous scaffold (Figure S4A, B). At PI4w, although new bone formation was observed around the scaffold periphery, notable gaps persisted between the scaffold edge and bone stump (Figure S4C–E). By PI12w, bone formation occurred at one edge of the scaffold; however, minimal bone penetration into the scaffold interior was observed (Figure S4F–H). In some cases, one bone stump obstructed the bone cavity due to scaffold interference. The volumes of newly formed bone inside the β-TCP scaffold were 1.3 ± 0.5 mm3 at PI4w and 3.4 ± 0.6 mm3 at PI12w (Figure S3G). At PI4w, the bone volume inside the β-TCP scaffold was significantly lower than that of the no-scaffold group. At PI12w, there was no statistically significant difference between the two groups (Figures S3G and S4I). These results collectively suggest that the CAp HC-structured scaffolds exhibit superior bone regeneration capacity compared to both no-scaffold and β-TCP scaffold conditions, supporting their potential for effective clinical bone reconstruction.

Resorption rates of materials depend on specific surface area, which is affected by micro- and nanoscale pores rather than macroscale pores. − P280, P440, and P640 have different macropore volumes but similar volumes of micro- and nanoscale pores (0.23–0.24 cm3/g). The macroscale specific surface area of P280, P440, and P640 (1.2 × 10–3–1.8 × 10–3 m2/g) is 1/1000 of the micro- and nanoscale specific surface area (1.7–2.3 m2/g). Thus, the total specific surface area, including macro-, micro-, and nanoscale regions, shows no apparent difference among P280, P440, and P640. Consequently, no significant difference in the resorption rate of material is observed among P280, P440, and P640. Meanwhile, the differences in bone volume, orientation, and mechanical strength were evident. These findings, therefore, indicate that macroscale pores affect bone reconstruction.

A limitation of this study is that we did not investigate the cellular and molecular mechanismssuch as osteoblast differentiation, osteoclast activity, or extracellular matrix secretionthrough which uniaxial macropore size affects bone regeneration. Although our findings clearly demonstrate the influence of pore size on bone volume and orientation, the underlying signaling pathways remain unexplored. Previous studies have indicated that scaffold pore size can modulate cellular behavior and gene expression. For instance, Li et al. demonstrated that a 3D-printed porous bioactive ceramic scaffold with 582 μm pores enhanced osteoblast proliferation and vascularization compared to scaffolds with 210 or 386 μm pores, as evidenced by increased numbers of osteoprotegerin- and CD31-positive cells. Similarly, Barba et al. reported that the architecture of calcium phosphate scaffolds can influence the expression of osteogenic markers such as BMP-2, ALP, collagen I, osteonectin, osteopontin, and osteocalcin. Future studies are planned to include in vitro experiments and molecular analyses to elucidate the signaling pathways involved and to strengthen the mechanistic understanding of our current findings.

Conclusions

We fabricated three types of CAp HC-structured scaffolds with different macropore sizes (280, 440, and 640 μm). In segmental bone defects, uniaxial 280 μm macropores formed small volumes of oriented bone, resulting in failed bony bridging and reduced strength. Uniaxial 440 μm macropores formed abundant oriented bones and restored the strength, while 640 μm macropores bridged the gaps by the abundant bones but did not further enhance the strength because of poor bone orientation. Thus, a uniaxial macropore size of 440 μm is suitable for efficiently reconstructing segmental bone defects. Our results contribute to the structural design and development of HC scaffolds for the reconstruction of segmental bone defects.

Supplementary Material

ao4c11634_si_001.pdf (1.2MB, pdf)

Acknowledgments

This study was supported by the Japan Agency for Medical Research and Development (grant numbers JP24ym0126098h0003 and JP24ym0126811j0003 to KH) and the Japan Society for the Promotion of Science (grant numbers JP23K18593 and JP23K25208 to KH).

Glossary

Abbreviations

HAp

hydroxyapatite

β-TCP

beta-calcium phosphate

CAp

carbonate apatite

HC

honeycomb

μ-CT

microcomputer tomography

SEM

scanning electron microscopy

XRD

X-ray diffraction

FTIR

Fourier transform infrared spectroscopy

ROI

range of interest

The Supporting Information is available free of charge at https://pubs.acs.org/doi/10.1021/acsomega.4c11634.

  • Identification processes for newly formed bones and materials in μ-CT analyses and BV images within the macropores of P280, P440, and P640 at PI4w and PI12w (PDF)

The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. K.S.: Formal analysis, Investigation, Data Curation, Methodology, WritingOriginal Draft. K.H.: conceptualization, supervision, methodology, validation, resources, writing–review and editing, and funding acquisition. Y.N. and K.I.: Validation and Project administration.

Any funds used to support the research of the manuscript should be placed here (per journal style).

The authors declare no competing financial interest.

References

  1. Andrzejowski P., Giannoudis P. V.. The ‘diamond concept’ for long bone non-union management. J. Orthop. Traumatol. 2019;20(1):21. doi: 10.1186/s10195-019-0528-0. [DOI] [PMC free article] [PubMed] [Google Scholar]
  2. Baker C. E., Moore-Lotridge S. N., Hysong A. A., Posey S. L., Robinette J. P., Blum D. M., Benvenuti M. A., Cole H. A., Egawa S., Okawa A., Saito M., McCarthy J. R., Nyman J. S., Yuasa M., Schoenecker J. G.. Bone fracture acute phase responseA unifying theory of fracture repair: clinical and scientific implications. Clin. Rev. Bone Miner. Metab. 2018;16:142–158. doi: 10.1007/s12018-018-9256-x. [DOI] [PMC free article] [PubMed] [Google Scholar]
  3. Sandberg O. H., Aspenberg P.. Inter-trabecular bone formation: a specific mechanism for healing of cancellous bone. Acta Orthop. 2016;87:459–465. doi: 10.1080/17453674.2016.1205172. [DOI] [PMC free article] [PubMed] [Google Scholar]
  4. Lopes D., Martins-Cruz C., Oliveira M. B., Mano J. F.. Bone physiology as inspiration for tissue regenerative therapies. Biomaterials. 2018;185:240–275. doi: 10.1016/j.biomaterials.2018.09.028. [DOI] [PMC free article] [PubMed] [Google Scholar]
  5. Hoerth R. M., Seidt B. M., Shah M., Schwarz C., Willie B. M., Duda G. N., Fratzl P., Wagermaier W.. Mechanical and structural properties of bone in non-critical and critical healing in rat. Acta Biomater. 2014;10:4009–4019. doi: 10.1016/j.actbio.2014.06.003. [DOI] [PubMed] [Google Scholar]
  6. Feltri P., Solaro L., Di Martino A., Candrian C., Errani C., Filardo G.. Union, complication, reintervention and failure rates of surgical techniques for large diaphyseal defects: a systematic review and meta-analysis. Sci. Rep. 2022;12(1):9098. doi: 10.1038/s41598-022-12140-5. [DOI] [PMC free article] [PubMed] [Google Scholar]
  7. Schmidt A. H.. Autologous bone graft: is it still the gold standard? Injury. 2021;52(Supplement 2):S18–S22. doi: 10.1016/j.injury.2021.01.043. [DOI] [PubMed] [Google Scholar]
  8. Marchand L. S., Kellam P. J., Dekeyser G. J., Haller J. M., Rothberg D. L., Higgins T. F.. Transfusion after harvesting bone graft with RIA: practice changes reduced transfusion rate by more than half. Injury. 2023;54:110797. doi: 10.1016/j.injury.2023.05.028. [DOI] [PubMed] [Google Scholar]
  9. Steijvers E., Ghei A., Xia Z.. Manufacturing artificial bone allografts: a perspective. Biomater. Transl. 2022;3:65–80. doi: 10.12336/biomatertransl.2022.01.007. [DOI] [PMC free article] [PubMed] [Google Scholar]
  10. Tanvir M. A. H., Khaleque M. A., Kim G. H., Yoo W. Y., Kim Y. Y.. The role of bioceramics for bone regeneration: history, mechanisms, and future perspectives. Biomimetics. 2024;9:230. doi: 10.3390/biomimetics9040230. [DOI] [PMC free article] [PubMed] [Google Scholar]
  11. Dehiya, B. ; Sindhu, A. ; Kumar, P. ; Dehiya, B. S. . Bioceramics for Hard Tissue Engineering Applications: A Review, 2018. http://www.ripublication.com.
  12. Shekhawat D., Singh A., Banerjee M. K., Singh T., Patnaik A.. Bioceramic composites for orthopaedic applications: A comprehensive review of mechanical, biological, and microstructural properties. Ceram. Int. 2021;47:3013–3030. doi: 10.1016/j.ceramint.2020.09.214. [DOI] [Google Scholar]
  13. Vidal L., Kampleitner C., Krissian S., Brennan M. Á., Hoffmann O., Raymond Y., Maazouz Y., Ginebra M. P., Rosset P., Layrolle P.. Regeneration of segmental defects in metatarsus of sheep with vascularized and customized 3D-printed calcium phosphate scaffolds. Sci. Rep. 2020;10(1):7068. doi: 10.1038/s41598-020-63742-w. [DOI] [PMC free article] [PubMed] [Google Scholar]
  14. Maevskaia E., Ghayor C., Bhattacharya I., Guerrero J., Weber F. E.. TPMS microarchitectures for vertical bone augmentation and osteoconduction: an in vivo study. Materials. 2024;17:2533. doi: 10.3390/ma17112533. [DOI] [PMC free article] [PubMed] [Google Scholar]
  15. Jodati H., Yılmaz B., Evis Z.. A review of bioceramic porous scaffolds for hard tissue applications: effects of structural features. Ceram. Int. 2020;46:15725–15739. doi: 10.1016/j.ceramint.2020.03.192. [DOI] [Google Scholar]
  16. Lash N. J., Feller J. A., Batty L. M., Wasiak J., Richmond A. K.. Bone grafts and bone substitutes for opening-wedge osteotomies of the knee: A systematic review. Arthroscopy. 2015;31:720–730. doi: 10.1016/j.arthro.2014.09.011. [DOI] [PubMed] [Google Scholar]
  17. Slevin O., Ayeni O. R., Hinterwimmer S., Tischer T., Feucht M. J., Hirschmann M. T.. The role of bone void fillers in medial opening wedge high tibial osteotomy: a systematic review. Knee Surg Sports Traumatol Arthrosc. 2016;24:3584–3598. doi: 10.1007/s00167-016-4297-5. [DOI] [PubMed] [Google Scholar]
  18. Koshino T., Murase T., Saito T.. Medial opening-wedge high tibial osteotomy with use of porous hydroxyapatite to treat medial compartment osteoarthritis of the knee. J. Bone Jt. Surg. Am. 2003;85:78–85. doi: 10.2106/00004623-200301000-00013. [DOI] [PubMed] [Google Scholar]
  19. Matsumine A., Myoui A., Kusuzaki K., Araki N., Seto M., Yoshikawa H., Uchida A.. Professor Uchida, A. H. Professor Yoshikawa, Calcium Hydroxyapatite Ceramic Implants in Bone Tumour Surgery. J. Bone Joint Surg. Br. 2004;86(5):719–25. doi: 10.1302/0301-620x.86b5.14242. [DOI] [PubMed] [Google Scholar]
  20. Kimura A., Seichi A., Inoue H., Hoshino Y.. Long-term results of double-door laminoplasty using hydroxyapatite spacers in patients with compressive cervical myelopathy. Eur. Spine J. 2011;20:1560–1566. doi: 10.1007/s00586-011-1724-7. [DOI] [PMC free article] [PubMed] [Google Scholar]
  21. Barba A., Diez-Escudero A., Maazouz Y., Rappe K., Espanol M., Montufar E. B., Bonany M., Sadowska J. M., Guillem-Marti J., Öhman-Mägi C., Persson C., Manzanares M. C., Franch J., Ginebra M. P.. Osteoinduction by foamed and 3D-printed calcium phosphate scaffolds: effect of nanostructure and pore architecture. ACS Appl. Mater. Interfaces. 2017;9:41722–41736. doi: 10.1021/acsami.7b14175. [DOI] [PubMed] [Google Scholar]
  22. Shibahara K., Hamai S., Akasaki Y., Nakashima Y.. Histological evaluation of the low-density region around beta-tricalcium phosphate scaffolds used in opening wedge high tibial osteotomy: A report of two cases. J. Orthop. Sci. 2023;28:1606–1610. doi: 10.1016/j.jos.2021.12.012. [DOI] [PubMed] [Google Scholar]
  23. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Honeycomb blocks composed of carbonate apatite, β-tricalcium phosphate, and hydroxyapatite for bone regeneration: effects of composition on biological responses. Mater. Today Bio. 2019;4:100031. doi: 10.1016/j.mtbio.2019.100031. [DOI] [PMC free article] [PubMed] [Google Scholar]
  24. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Granular honeycombs composed of carbonate apatite, hydroxyapatite, and β-tricalcium phosphate as bone graft substitutes: effects of composition on bone formation and maturation. ACS Appl. Bio Mater. 2020;3:1787–1795. doi: 10.1021/acsabm.0c00060. [DOI] [PubMed] [Google Scholar]
  25. Yotsova R., Peev S.. Biological properties and medical applications of carbonate apatite: A systematic review. Pharmaceutics. 2024;16:291. doi: 10.3390/pharmaceutics16020291. [DOI] [PMC free article] [PubMed] [Google Scholar]
  26. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Superiority of triply periodic minimal surface Gyroid structure to strut-based grid structure in both strength and bone regeneration. ACS Appl. Mater. Interfaces. 2023;15:34570–34577. doi: 10.1021/acsami.3c06263. [DOI] [PubMed] [Google Scholar]
  27. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Transformable carbonate apatite chains as a novel type of bone graft. Adv. Healthcare Mater. 2024;13(12):e2303245. doi: 10.1002/adhm.202303245. [DOI] [PubMed] [Google Scholar]
  28. Hayashi K., Yanagisawa T., Kishida R., Ishikawa K.. Effects of scaffold shape on bone regeneration: tiny shape differences affect the entire system. ACS Nano. 2022;16:11755–11768. doi: 10.1021/acsnano.2c03776. [DOI] [PMC free article] [PubMed] [Google Scholar]
  29. Hayashi K., Yanagisawa T., Shimabukuro M., Kishida R., Ishikawa K.. Granular honeycomb scaffolds composed of carbonate apatite for simultaneous intra- and inter-granular osteogenesis and angiogenesis. Mater. Today Bio. 2022;14:100247. doi: 10.1016/j.mtbio.2022.100247. [DOI] [PMC free article] [PubMed] [Google Scholar]
  30. Hayashi K., Yanagisawa T., Kishida R., Tsuchiya A., Ishikawa K.. Gear-shaped carbonate apatite granules with a hexagonal macropore for rapid bone regeneration. Comp. Struct. Biotechnol. J. 2023;21:2514–2523. doi: 10.1016/j.csbj.2023.03.053. [DOI] [PMC free article] [PubMed] [Google Scholar]
  31. Hayashi K., Kato N., Kato M., Ishikawa K.. Impacts of channel direction on bone tissue engineering in 3D-printed carbonate apatite scaffolds. Mater. Des. 2021;204:109686. doi: 10.1016/j.matdes.2021.109686. [DOI] [Google Scholar]
  32. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Effects of space dimensionality within scaffold for bone regeneration with large and oriented blood vessels. Materials. 2023;16:7518. doi: 10.3390/ma16247518. [DOI] [PMC free article] [PubMed] [Google Scholar]
  33. Shibahara K., Hayashi K., Nakashima Y., Ishikawa K.. Roles of pore architecture of artificial bone grafts in invasion competition between bone and fibrous tissue and orientation of regenerated bone. Int. J. Bioprint. 2024;10:2323. doi: 10.36922/ijb.2323. [DOI] [Google Scholar]
  34. Shibahara K., Hayashi K., Nakashima Y., Ishikawa K.. Controlling the pore size of carbonate apatite honeycomb scaffolds enhances orientation and strength of regenerated bone. Biomater. Adv. 2025;166:214026. doi: 10.1016/j.bioadv.2024.214026. [DOI] [PubMed] [Google Scholar]
  35. Hayashi K., Munar M. L., Ishikawa K.. Effects of macropore size in carbonate apatite honeycomb scaffolds on bone regeneration. Mater. Sci. Eng., C. 2020;111:110848. doi: 10.1016/j.msec.2020.110848. [DOI] [PubMed] [Google Scholar]
  36. Hayashi K., Shimabukuro M., Kishida R., Tsuchiya A., Ishikawa K.. Honeycomb scaffolds capable of achieving barrier membrane-free guided bone regeneration. Mater. Adv. 2021;2:7638–7649. doi: 10.1039/D1MA00698C. [DOI] [Google Scholar]
  37. Hayashi K., Shimabukuro M., Kishida R., Tsuchiya A., Ishikawa K.. Structurally optimized honeycomb scaffolds with outstanding ability for vertical bone augmentation. J. Adv. Res. 2022;41:101–112. doi: 10.1016/j.jare.2021.12.010. [DOI] [PMC free article] [PubMed] [Google Scholar]
  38. Shimabukuro M., Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Effects of carbonate ions in phosphate solution on the fabrication of carbonate apatite through a dissolution–precipitation reaction. Ceram. Int. 2022;48:1032–1037. doi: 10.1016/j.ceramint.2021.09.188. [DOI] [Google Scholar]
  39. Sugimoto K., Zhou Y., Galindo T. G. P., Kimura R., Tagaya M.. Investigation of surface layers on biological and synthetic hydroxyapatites based on bone mineralization process. Biomimetics. 2023;8:184. doi: 10.3390/biomimetics8020184. [DOI] [PMC free article] [PubMed] [Google Scholar]
  40. Shibahara K., Hayashi K., Nakashima Y., Ishikawa K.. Reconstruction of load-bearing segmental bone defects using carbonate apatite honeycomb blocks. ACS Mater. Au. 2023;3:321–336. doi: 10.1021/acsmaterialsau.3c00008. [DOI] [PMC free article] [PubMed] [Google Scholar]
  41. Shibahara K., Hayashi K., Nakashima Y., Ishikawa K.. Honeycomb scaffold-guided bone reconstruction of critical-sized defects in rabbit ulnar shafts. ACS Appl. Bio Mater. 2021;4:6821–6831. doi: 10.1021/acsabm.1c00533. [DOI] [PubMed] [Google Scholar]
  42. Shibahara K., Hayashi K., Nakashima Y., Ishikawa K.. Effects of channels and micropores in honeycomb scaffolds on the reconstruction of segmental bone defects. Front. Bioeng. Biotechnol. 2022;10:825831. doi: 10.3389/fbioe.2022.825831. [DOI] [PMC free article] [PubMed] [Google Scholar]
  43. Bodde E. W. H., Spauwen P. H. M., Mikos A. G., Jansen J. A.. Closing capacity of segmental radius defects in rabbits. J. Biomed. Mater. Res. 2008;85:206–217. doi: 10.1002/jbm.a.31549. [DOI] [PubMed] [Google Scholar]
  44. El Backly R. M., Chiapale D., Muraglia A., Tromba G., Ottonello C., Santolini F., Cancedda R., Mastrogiacomo M.. A modified rabbit ulna defect model for evaluating periosteal substitutes in bone engineering: A pilot study. Front. Bioeng. Biotechnol. 2014;2:80. doi: 10.3389/fbioe.2014.00080. [DOI] [PMC free article] [PubMed] [Google Scholar]
  45. Boudaoud A., Burian A., Borowska-Wykręt D., Uyttewaal M., Wrzalik R., Kwiatkowska D., Hamant O.. FibrilTool, an ImageJ plug-in to quantify fibrillar structures in raw microscopy images. Nat. Protoc. 2014;9:457–463. doi: 10.1038/nprot.2014.024. [DOI] [PubMed] [Google Scholar]
  46. Patel A., Zaky S. H., Li H., Schoedel K., Almarza A. J., Sfeir C., Sant V., Sant S.. Bottom-up self-assembled hydrogel-mineral composites regenerate rabbit ulna defect without added growth factors. ACS Appl. Bio Mater. 2020;3:5652–5663. doi: 10.1021/acsabm.0c00371. [DOI] [PubMed] [Google Scholar]
  47. Ślósarczyk A., Paszkiewicz Z., Paluszkiewicz C.. FTIR and XRD evaluation of carbonated hydroxyapatite powders synthesized by wet methods. J. Mol. Struct. 2005;744–747:657–661. doi: 10.1016/j.molstruc.2004.11.078. [DOI] [Google Scholar]
  48. Wang M., Qian R., Bao M., Gu C., Zhu P. R.. FT-IR and XRD study of bovine bone mineral and carbonated apatites with different carbonate levels. Mater. Lett. 2018;210:203–206. doi: 10.1016/j.matlet.2017.09.023. [DOI] [Google Scholar]
  49. Li W., Dai F., Zhang S., Xu F., Xu Z., Liao S., Zeng L., Song L., Ai F.. Pore size of 3D-printed polycaprolactone/polyethylene glycol/hydroxyapatite scaffolds affects bone regeneration by modulating macrophage polarization and the foreign body response. ACS Appl. Mater. Interfaces. 2022;14:20693–20707. doi: 10.1021/acsami.2c02001. [DOI] [PubMed] [Google Scholar]
  50. Qin H., Wei Y., Han J., Jiang X., Yang X., Wu Y., Gou Z., Chen L.. 3D printed bioceramic scaffolds: adjusting pore dimension is beneficial for mandibular bone defects repair. J. Tissue Eng. Regen Med. 2022;16:409–421. doi: 10.1002/term.3287. [DOI] [PubMed] [Google Scholar]
  51. Petersen A., Princ A., Korus G., Ellinghaus A., Leemhuis H., Herrera A., Klaumünzer A., Schreivogel S., Woloszyk A., Schmidt-Bleek K., Geissler S., Heschel I., Duda G. N.. A biomaterial with a channel-like pore architecture induces endochondral healing of bone defects. Nat. Commun. 2018;9:4430. doi: 10.1038/s41467-018-06504-7. [DOI] [PMC free article] [PubMed] [Google Scholar]
  52. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Carbonate apatite micro-honeycombed blocks generate bone marrow-like tissues as well as bone. Adv. Biosyst. 2019;3(12):e1900140. doi: 10.1002/adbi.201900140. [DOI] [PubMed] [Google Scholar]
  53. Hayashi K., Ishikawa K.. Effects of nanopores on the mechanical strength, osteoclastogenesis, and osteogenesis in honeycomb scaffolds. J. Mater. Chem. B. 2020;8:8536–8545. doi: 10.1039/D0TB01498B. [DOI] [PubMed] [Google Scholar]
  54. Hayashi K., Tsuchiya A., Shimabukuro M., Ishikawa K.. Multiscale porous scaffolds constructed of carbonate apatite honeycomb granules for bone regeneration. Mater. Des. 2022;215:110468. doi: 10.1016/j.matdes.2022.110468. [DOI] [Google Scholar]
  55. Hayashi K., Ishikawa K.. Honeycomb scaffolds fabricated using extrusion molding and the sphere-packing theory for bone regeneration. ACS Appl. Bio Mater. 2021;4:721–730. doi: 10.1021/acsabm.0c01279. [DOI] [Google Scholar]
  56. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Channel aperture characteristics of carbonate apatite honeycomb scaffolds affect ingrowths of bone and fibrous tissues in vertical bone augmentation. Bioengineering. 2022;9:627. doi: 10.3390/bioengineering9110627. [DOI] [PMC free article] [PubMed] [Google Scholar]
  57. Hayashi K., Zhang C., Taleb Alashkar A. N., Ishikawa K.. Carbonate apatite honeycomb scaffold-based drug delivery system for repairing osteoporotic bone defects. ACS Appl. Mater. Interfaces. 2024;16:45956–45968. doi: 10.1021/acsami.4c08047. [DOI] [PMC free article] [PubMed] [Google Scholar]
  58. Hayashi K., Shimabukuro M., Zhang C., Taleb Alashkar A. N., Kishida R., Tsuchiya A., Ishikawa K.. Silver phosphate-modified carbonate apatite honeycomb scaffolds for anti-infective and pigmentation-free bone tissue engineering. Mater. Today Bio. 2024;27:101161. doi: 10.1016/j.mtbio.2024.101161. [DOI] [PMC free article] [PubMed] [Google Scholar]
  59. Hayashi K., Kishida R., Tsuchiya A., Ishikawa K.. Hematopoietic function restoration by transplanting bone marrow niches in vivo engineered using carbonate apatite honeycomb bioreactors. Small Struct. 2024;5(10):2400065. doi: 10.1002/sstr.202400065. [DOI] [Google Scholar]
  60. Hayashi K., Shimabukuro M., Ishikawa K.. Antibacterial honeycomb scaffolds for achieving infection prevention and bone regeneration. ACS Appl. Mater. Interfaces. 2022;14:3762–3772. doi: 10.1021/acsami.1c20204. [DOI] [PubMed] [Google Scholar]

Associated Data

This section collects any data citations, data availability statements, or supplementary materials included in this article.

Supplementary Materials

ao4c11634_si_001.pdf (1.2MB, pdf)

Articles from ACS Omega are provided here courtesy of American Chemical Society

RESOURCES