Functionalized 3D-printed scaffolds for enhanced osteogenesis and guided bone regeneration

Abstract

In this study, we introduced an innovative approach to guided bone regeneration (GBR) that effectively addresses the challenges of treating large bone defects. Our pioneering 3D-printed multifunctional scaffolds uniquely integrate polycaprolactone (PCL), chitosan (Cs), L-arginine (L-Arg), and β-tricalcium phosphate (β-TCP), leveraging the synergistic effects of these materials to enhance immunomodulation, bioactivity, and mechanical integrity. These PCL/Cs-L-Arg/βTCP scaffolds exhibit remarkable mechanical properties (Young's modulus ∼32.84 ± 4.11 MPa) and maintain structural integrity for 60 days under physiological conditions when fabricated through extrusion-based 3D printing. A key feature of this composite is the dual role of L-Arg, which not only supports osteogenesis but also acts as a potent immunomodulator. The scaffolds facilitate the sustained release of L-arginine over 21 days, fostering a pro-regenerative environment that promotes significant immunomodulatory effects, including a decrease in pro-inflammatory cytokines (IL-6, TNF-α) and an enhancement of anti-inflammatory and osteogenic growth factors (BMP-2, TGF-β) in macrophages. This cytokine profile shift suggests a transition from a pro-inflammatory M1 phenotype to an anti-inflammatory M2 phenotype. A progressive increase in alkaline phosphatase activity, nearly double that of PCL/Cs scaffolds by day 21, reflects enhanced osteogenic differentiation. Additionally, the scaffolds demonstrate exceptional bioactivity, with over 83% and 93% reductions in calcium and phosphorus ions, respectively, in simulated body fluid over 28 days, as evidenced by Alizarin red staining. This integrated approach signifies a major breakthrough in biomaterial design for GBR, presenting transformative potential for treating bone defects in dental and orthopedic applications, and marking a significant leap forward in the field of bone regeneration.

---

1. Introduction

Sufficient bone volume is a critical requirement for effective therapy in dental implantology. Insufficient horizontal or vertical bone volume can compromise the stability of implant placement, creating significant challenges in achieving successful outcomes.¹ Moreover, the process of bone regeneration and repair remains a major challenge in medical care, particularly in cases involving infection, trauma, or malignancies.² To address these challenges, various techniques have been developed to enhance bone volume and facilitate bone tissue regeneration.^3,4

Guided bone regeneration (GBR) is a widely employed strategy to address bone volume deficits by promoting bone repair and regeneration. This technique enhances the reliability of bone augmentation and ensures long-term stability in the regenerated area.⁵ A barrier membrane is often used as a biological separator between gingival and bone tissue to prevent fibroblasts and epithelial cells from infiltrating the bone defect site. This separation ensures uninterrupted bone regeneration, allowing for complete and gradual bone healing. An ideal GBR membrane should provide effective tissue integration, ease of clinical application, spatial control, and dimensional stability.^6,7

Various GBR barriers have been developed and clinically utilized, primarily combining bone grafts with collagen membranes. Collagen membranes, as natural biomaterials, offer excellent cell affinity, biodegradability, and bioresorbability, eliminating the need for a second surgery. However, their limited mechanical strength, inadequate space maintenance, and suboptimal osteogenic activity hinder their long-term effectiveness.^8,9 To overcome these challenges, researchers have integrated synthetic and natural biodegradable polymers with bioactive materials and bioceramics, enhancing both mechanical properties and biological functionality. This approach better mimics natural tissue regeneration, ensuring stability.¹⁰

Poly(ε-caprolactone) (PCL), a synthetic polymer approved by the food and drug administration (FDA), has emerged as a promising material for GBR applications due to its superior biocompatibility, controllable degradation, and non-toxic byproducts (water and carbon dioxide) during hydrolysis.¹¹ Additionally, its low melting point (50–60 °C) makes it highly versatile for various manufacturing processes. However, PCL's slow degradation rate often lags behind the rate of new bone formation, and its hydrophobic nature hinders cell attachment and osteogenic activity. To address these challenges, natural polymers such as alginate, chitosan, and silk fibroin, as well as bioactive additives, have been incorporated to enhance its physicochemical and biological properties.^12,13

Chitosan (CS), a natural polymer, has garnered significant attention as a biodegradable membrane material due to its low toxicity, strong biocompatibility, mucosal adherence, appropriate mechanical strength, and cost-effectiveness.¹⁴ Moreover, CSs structure closely resembles the polysaccharides found in the extracellular matrix, enabling optimal biodegradation and facilitating new bone tissue development.¹⁵

L-Arginine (L-Arg), a critical nutrient, plays a significant role in tissue development, particularly during bone repair. L-Arg promotes hydroxyapatite formation by facilitating interactions between PO₄³⁻ and Ca²⁺ ions and enhancing bone mineral density by activating growth hormone expression. Additionally, L-Arg metabolism produces nitrogen-containing compounds essential for tissue regeneration, such as nitric oxide (NO), creatine, and polyamines. NO, in particular, regulates bone tissue formation and inhibits osteoclast activity, making L-Arg a valuable component for enhancing bone regeneration.^16–19

The integration of bioactive ceramic additives, such as bioglass and calcium phosphates (e.g., hydroxyapatite and β-tricalcium phosphate [β-TCP]), into polymer-based GBR membranes is a well-established approach to improve bioactivity, mechanical strength, and mineralization capacity.²⁰ Previous study has documented the development of biomaterials employing a combination of PCL and β-tricalcium phosphate (βTCP) to enhance regeneration characteristics. It is important to mention that βTCP (Ca₃(PO₄)₂) is a synthetic bone graft substitute that is capable of promoting the growth of bone tissue and being absorbed by cells.¹² Nevertheless, there is still potential for improvement in the process of enhancing bone regeneration by effectively mixing different polymers and ceramics to achieve their preferred unique behaviors, including mechanical strength and compatibility with bone tissue.²¹

GBR membranes can possess a wide range of properties depending on the method that they are manufactured by and their structure. The porosity of the scaffold is a crucial factor that affects tissue regeneration.²² Recently, 3D printing has emerged as a convenient and reproducible method for fabricating flexible and porous membranes, offering precise control over scaffold properties. Various 3D printing techniques, including fused deposition modeling (FDM), selective laser sintering (SLS), and stereolithography (SLA), have been explored for scaffold fabrication, each with unique advantages in terms of resolution, material compatibility, and mechanical properties. In this study, the FDM technique was selected due to its suitability for processing thermoplastic polymers, cost-effectiveness, and ability to create highly porous structures with controlled architecture.^23,24

While clinical GBR membranes are often composed of soft materials to ensure flexibility and ease of adaptation, the choice of rigid materials such as PCL/Cs and β-TCP in this study is based on their superior mechanical stability, osteoconductive properties, and controlled degradation. Soft membranes, despite their biocompatibility, may lack the necessary structural integrity to maintain the defect space, leading to early collapse and insufficient bone regeneration. In contrast, rigid scaffolds provide enhanced dimensional stability, allowing for sustained osteogenic activity and better integration with surrounding bone tissue. Furthermore, the gradual degradation of PCL/Cs and β-TCP aligns with the bone healing timeline, supporting long-term regeneration without premature resorption. By leveraging these advantages, the proposed GBR membrane aims to address the limitations of traditional soft membranes while maintaining an optimal balance between mechanical support and bioactivity.^25,26

In this study, we developed a GBR membrane using a 3D-printed polycaprolactone–chitosan (PCL/CS) scaffold functionalized with L-arginine (L-Arg) and β-TCP. We evaluated its physical, chemical, and mechanical properties, as well as its biodegradability, biological compatibility, and immune response. Our study introduces a groundbreaking approach by developing a unique PCL/Cs scaffold functionalized with L-arginine (L-Arg) and β-TCP, where the strategic combination of these components creates an unprecedented synergy to optimize bone regeneration. The controllable degradation of PCL, the antibacterial properties of chitosan, and the osteogenic enhancement provided by L-Arg, together with the bone-regenerative support of β-TCP, distinguish this innovative composite scaffold as a highly promising candidate for advancing GBR applications.

2. Materials and methods

2.1. Materials and reagents

The research encompassed the following materials: polycaprolactone (PCL) with a molecular weight of 40 000 Da, low molecular weight chitosan (Cs), L-arginine (L-Arg), β-tricalcium phosphate (β-TCP), lipopolysaccharide (LPS), and the reagents 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT), 4′,6-diamidino-2-phenylindole (DAPI), bovine serum albumin (BSA), and the alkaline phosphatase activity (ALP) evaluation kit were procured from Sigma-Aldrich (USA). Trifluoroacetic acid (TFA: CF₃CO₂H) and dimethyl sulfoxide (DMSO) were procured from Merck (Germany). Supplementary components, including Dulbecco's Modified Eagle Medium (DMEM-high), streptomycin, penicillin, and fetal bovine serum (FBS), were procured from Bioidea (Iran). All tests utilized the Griess reagent kit (Natrix, Iran), the Bradford protein assay kit (Partocib, Iran), tissue growth factor-β (TGF-β), tumor necrosis factor-α (TNF-α), interleukin-6 (IL-6), and interleukin-10 (IL-10) ELISA kits (KPG Co., Iran), along with distilled water (DDW, Bioidea, Iran).

2.2. Fabrication 3D-printed scaffolds

All microporous composite scaffolds in the current study were engineered using the liquid deposition modeling (LDM) 3D printing machinery (Z-Morph, China). At the initial stage of scaffold construction, PCL was dissolved in TFA solvent at a concentration of 30% (w/v) at a temperature of 55 °C. In addition, Cs was dissolved in TFA solvent at a concentration of 15% (w/v) at a temperature of 50 °C to generate the composite scaffold solution. After obtaining a uniform mixture of PCL and Cs in a polymer solution, β-TCP and L-Arg powder were added to the solution with concentrations of 20% and 5% (w/v), respectively. The mixture was then manually stirred to confirm uniformity to develop a 3D printing ink. Scaffolds with a cubic form were created using a three-dimensional microstructure, specifically provided as PCL/Cs-L-Arg/β-TCP. A model in the shape of a square with a small thickness, measuring 40 × 40 × 0.5 mm³, was designed for this particular purpose. The distance between lines in the model is 500 μm. The design was created using SolidWorks, version 2020. The 3D printing process was executed utilizing an extrusion-based 3D printer with the ideal parameters listed in Table 1.

Table 1 Parameters for 3D printing of microporous scaffolds

Speed (mm s^−1)	Nuzzle diameter (mm)		Extrusion multiplayer	Retraction	Layer height	Bed temperature (°C)
	Inner	Outer
5	0.6	0.8	0.0088	1.2	0.1	40

---

2.3. Material characterization

2.3.1. FTIR. The scaffolds were analyzed for their chemical composition with Fourier transform infrared (FTIR: 6300, JASCO, Japan) spectroscopy in the range of 400–4000 cm⁻¹.

2.3.2. SEM. The morphology of 3D-printed scaffolds was investigated via scanning electron microscopy (SEM: TESCAN, MIRA model, Czech Republic) at an accelerating voltage of 5 kV. The filament's diameter was calculated using ImageJ software based on SEM images.²⁷

2.3.3. Contact angle and biodegradation. The water contact angle evolution at room temperature was used to assess the wettability of the PCL/Cs, PCL/Cs-L-Arg/β-TCP scaffolds (n = 3).

The degradation study was done according to the ISO 13781: 1997, MOD standard.²⁸ Several specimens of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/β-TCP composites were each placed in a 30 mL centrifuge tube. The tubes were filled with PBS with a pH ∼ 7.4 ± 0.2 and stored at a temperature of 37 °C. At specific times during 1, 3, 5, 7, 14, 21, 28, 35, 42, 49, 56, and 63 days, the samples were collected subsequently dried using blotting before being weighed. At each time point, the samples were placed again into vials containing a fresh PBS solution. The mean results were obtained from a minimum number of three repeats. The percentage of mass change was calculated using the following eqn (1):²⁹

In this equation, W₂ represents the sample weight after degradation in time points, and W₁ represents the sample weight before soaking in PBS.

It is important to mention that a distinct experimental group was utilized to measure the pH fluctuation, whereas the degradation solution remained same during the whole degradation experiment.

2.4. Mechanical behavior

The mechanical properties of 3D-printed membranes have been evaluated using a tensile test in accordance with the ASTM D638 standard. The test utilized a Hounsfield machine from the United Kingdom, which included a load cell capable of handling up to 50 N. The scaffolds, with dimensions of 30 mm × 10 mm (n = 3), were cut and exposed to tensile loading at a constant strain rate of 10 mm min⁻¹ using a load cell with a capacity of 10 N. The obtained stress–strain curves were used to determine the tensile strength and elastic modulus, which relates to the slope of the stress–strain curve in the linear area.³⁰

2.5. L-Arg release

The amount of L-Arg in PCL/Cs-β-TCP and its subsequent release from the membrane were quantified using the Bradford protein assay kit. To determine the concentration of L-Arg, the samples were initially divided into 1.5 × 1.5 cm² segments. The samples were thereafter immersed in a 10 mL solution of phosphate buffered saline (PBS) and subjected to rotation at 150 rpm at a temperature of 37 °C.

To assess the release of L-Arg, PCL/Cs-L-Arg/β-TCP membranes were submerged in a 10 mL PBS solution at pH 7.4 (n = 3) and maintained at 37 °C for 21 days. At designated intervals, 10 μL of PBS was introduced to a well and mixed with 80 μL of water to formulate a dilution. Subsequently, a 20 μL aliquot of reagent A, containing Coomassie brilliant blue G-250 dye, was introduced into the wells and permitted to incubate for 5 minutes to facilitate interaction with the L-Arg peptide. The absorbance of the solutions was quantified at 595 nm utilizing a microplate reader. The concentration of L-Arg was determined by constructing a standard curve for BSA in accordance with the manufacturer's guidelines.³¹

2.6. Bioactivity assay

The assessment of the scaffolds' bioactivity and the formation of hydroxyapatite was done by immersing the samples in a simulated body fluid (SBF) solution, according to the ASTM F2150-19 standard. Afterward, 1 × 1 cm² of the scaffolds were submerged in SBF and kept at a temperature of 37 °C for a duration of 28 days. Each week, starting from the first week and continuing until the fourth week, was assessed individually. Indeed, the bioactivity of the scaffolds was evaluated on a weekly basis for each individual scaffold. The release of the Ca and P ions was assessed using inductively coupled plasma mass spectrometry (ICP-MS), generally known as ICP. Inductively coupled plasma (ICP) is a widely used method in emission spectrometry for the identification and measurement of components in sample analysis. Inductively coupled plasma spectrometry use an inductively coupled plasma (ICP) to ionize the elements in the sample, and then measures these ions using a mass spectrometer (MS).³⁰ Also, the samples were taken out of SBF and rinsed with distilled water at 28-day intervals. Surface changes of the samples were then analyzed using SEM.³²

2.7. In vitro biological assay

2.7.1. MTT and DAPI staining. The viability of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/β-TCP composites was assessed utilizing the MG63 osteoblast-like cell line obtained from the Pasteur Institute Cell Bank in Iran. Membranes were produced by 3D printing and configured as quarter circles with a 10 mm radius. Sterilization was accomplished by exposing the samples to UV radiation for 20 and 30 minutes on each side.

The cytotoxicity of the materials was evaluated using the direct contact method with the MG63 cell line. Ten 10⁴ cells were inoculated onto the samples and cultivated in Dulbecco's Modified Eagle's Medium (DMEM) enriched with 10% fetal bovine serum (FBS) and 1% streptomycin. The cells were incubated at 37 °C in a 5% CO₂ atmosphere.

Cell viability was evaluated at 1, 3, and 5 days by the MTT test, in accordance with the methodology described in prior work. Subsequent to the removal of the culture media, a 0.5 wt% MTT solution in PBS was introduced, and the samples were incubated for 4 h. Dimethyl sulfoxide (DMSO, Sigma) was subsequently employed to dissolve the formazan crystals, and the optical density (OD) was assessed at 450 nm utilizing a microplate reader (Hiperion, MPR4). The cell viability was determined by applying the following eqn (2):³³

where OD_S (Sample), OD_B (blank), and OD_C (control) indicate the optical density of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP samples, DMSO (blank), and TCP (control), respectively.

Five days after cell attachment, the number of viable cells was assessed using DAPI staining (4′,6-diamidino-2-phenylindole) to visualize the cell nuclei. After the removal of the medium at specified time intervals, the 3D-printed scaffolds were washed with PBS. The samples were subsequently subjected to ethanol solutions of increasing concentrations (50%, 70%, 90%, and 100%) for 20 minutes each. After ethanol removal, the scaffolds were stabilized by immersion in a 2.5% glutaraldehyde solution and subsequently rinsed with PBS. The samples were stained with DAPI for 30 seconds, thereafter, washed with PBS. Fluorescent images of live cells were obtained using a fluorescent microscope (Olympus BX51, Japan).³⁴

2.7.2. ALP activity. ALP activity was measured according to the manufacturer's guidelines, using an ALP test kit to assess para-nitrophenyl phosphate (pNPP) levels. To quantify the total protein content, the samples were treated with a radioimmunoprecipitation lysis buffer and incubated on ice for 2 h. The samples were then transferred to microtubes and centrifuged at 15 rpm for 15 minutes at 4 °C. The supernatant, containing the total protein, was collected and combined with the ALP kit reagents. The absorbance at 405 nm was measured using a microplate reader to determine both ALP activity and protein concentration. ALP activity, expressed in units per liter, was normalized to the total protein content of each sample.³⁵

2.7.3. Alizarin red staining. The Alizarin red staining method was employed to assess the osteoblasts' ability to deposit calcium and form a mineralized matrix.³⁶ On the 7th day of cell culture, extracellular calcium formation on the scaffolds was examined using this staining technique. At the designated time point, the culture medium was removed, and the cells were fixed with 10% formaldehyde for 15 minutes at room temperature. The samples were then rinsed three times with distilled water to remove any residual fixative. Following this, the samples were incubated with 1% Alizarin red solution (pH 4.1) for 30 minutes and subsequently washed six times with distilled water to eliminate excess stain. The stained cells were captured using a camera (SCC-101P, Samsung/Korea).³⁷

2.7.4. Inflammatory responses. The activity of RAW264.7 murine macrophages, sourced from the Pasteur Institute Cell Bank in Iran, was assessed by directly seeding the cells onto PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/β-TCP substrates. Following sterilization of the samples, macrophages were inoculated at a density of 10^5 cells per well and grown in high-glucose DMEM enriched with 10% FBS and 1% streptomycin. The cells were incubated at 37 °C in an atmosphere containing 5% CO₂.

In the enzyme-linked immunosorbent test (ELISA), 12-well plates were utilized, and 2 × 10⁵ cells were distributed to the samples. The release of proinflammatory and anti-inflammatory cytokines, including IL-6, TNF-α, BMP2, and TGF-β, was evaluated via ELISA. The cell culture media was treated with 100 ng mL⁻¹ lipopolysaccharide (LPS) from Escherichia coli for 24 h to produce inflammation. Following a 24-hour incubation of RAW264.7 cells with the materials, the culture medium was harvested and subjected to centrifugation. The levels of cytokines IL-6, TNF-α, BMP2, and TGF-β were measured using ELISA kits in accordance with the manufacturer's instructions.³⁸

2.7.5. Nitric oxide (NO) measurement. To evaluate nitric oxide (NO) production, the RAW-264.7 mouse macrophage cell line was used. Macrophages were cultured on 3D-printed membranes in a 48-well plate at a density of 10⁵ cells per mL, with TCP used as the control. After a 48-hour incubation, NO production was measured using the Griess reagent method. For this, 50 μL of the supernatant, collected after centrifugation, was mixed with 50 μL of a solution containing 0.1% N-1-naphthylethylenediamine dihydrochloride (NED) and 50 μL of a 1% sulfanilamide solution in 5% phosphoric acid. The reaction was incubated at room temperature for 10 minutes, and absorbance was measured at 540 nm using a microplate reader. NO levels were quantified by comparing the absorbance to a standard nitrite reference curve.³⁹

2.8. Statistical analysis

The ANOVA test was used for statistical analysis of the data. The Tukey-Kramer post hoc test was conducted using GraphPad Prism Software (Version 9) to ascertain the statistically significant differences across the groups. A p-value of less than 0.05 was deemed as statistically significant.

3. Results and discussion

3.1. FTIR

The primary functional groups of the PCL/CS, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP substances have been identified using the FTIR analysis (Fig. 1A).

Fig. 1 (A) Chemical characterization of 3D printed structures using FTIR analysis and (B) SEM images of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds (scale bar: 500 μm).

As seen in Fig. 1A, The PCL/CS, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP spectra have peaks at 2880 cm⁻¹ and about 3350 cm⁻¹, which correspond to the stretching vibrations of –CH and –OH of CS, respectively.⁴⁰ The peaks shown at 1650 cm⁻¹ and 1599 cm⁻¹ correspond to the amide I and II functional groups of CS in the spectra of PCL/CS, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP.⁴¹ The peaks observed in the spectral regions of 1580–1690 cm⁻¹ and 3155–3650 cm⁻¹ correspond to the H₂O molecules present in L-Arg.³³ The peak considered at 1728 cm⁻¹ corresponds to the symmetric stretching of the C=O bond, including the carbonyl group. Additionally, the peaks at 1157 cm⁻¹ and 1269 cm⁻¹ indicate the stretching of the C–O bond in the crystalline section of the PCL present in the PCL/CS, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP samples.⁴² The spectral lines seen at 2819–2996 cm⁻¹ indicate the stretching vibrations of C–H and C–C bonds in the amorphous phase. The presence of phosphate groups (P=O) has been shown by wavenumbers of 955 cm⁻¹ and 1082 cm⁻¹, which are associated with the contribution of βTCP.⁴³ In the PCL/Cs-L-Arg/βTCP composite, the bond at 1071 cm⁻¹ corresponds to the stretching vibration of the PO₄³⁻group.⁴⁴ The interaction between the carbonyl group (C=O) in PCL/Cs-L-Arg and the hydroxyl group (OH) in βTCP in the PCL/Cs-L-Arg/βTCP sample leads to a decrease in peak intensity. The peaks seen at 1443–1538 cm⁻¹ are indicative of the CH₂ bond, which provides evidence for the existence of PCL in all samples. The FTIR analysis of the PCL/Cs-L-Arg/βTCP sample revealed a peak at 1479 cm⁻¹, which corresponds to the asymmetric stretching of the CH₂ bond.⁴⁵ The presence of CS, βTCP, and PCL in the composite scaffolds was confirmed based on the exhibited peaks.

3.2. Morphology

The SEM images of 3D printed PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds are shown in Fig. 1(B). It is evident that the 3D-printed construction demonstrated a uniform structure with interconnected pores. The pore diameters of the PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/TCP samples were 405.1 ± 25.6 μm, 548.1 ± 41.5 μm, and 473.9 ± 16.1 μm, respectively. The viscosity of the solution was reduced by the addition of L-Arg to PCL/Cs, which in turn reduced the diameter of the 3D printed filaments. Consequently, the pore diameter increased.^46,47 However, the viscosity of the PCL/Cs-L-Arg solution was increased by the addition of TCP, resulting in a decrease in pore diameter in comparison to PCL/Cs-L-Arg. The presence of macro pores (400–600 μm) in the membrane structure being designed can facilitate nutrient supply, and waste removal and enhance cell-matrix interaction.⁴⁸

3.3. Water wettability

The wettability of membranes in various groups, including PCL, PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP, was investigated using the water contact angle measurement (Fig. 2A and B). The findings indicated that the contact angle of PCL was reduced because of the addition of Cs. The contact angle of the PCL/Cs composite scaffold is 74.6 ± 4°, while the contact angle of PCL alone is reported to be 120.6 ± 3°. The contact angle of the PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP scaffolds has significantly decreased, with values of 58.01 ± 2° and 45.05 ± 3°, respectively. This decrease may be attributed to the hydrophilic nature, the presence of amine and hydroxyl groups in CS and L-Arg, and the hydrophilic structure of βTCP.^33,49,50 The results indicate that the wettability of the scaffold is influenced by a variety of parameters, such as its chemical composition, surface texture, surface porosities, and crystallinity.⁵¹

Fig. 2 (A) and (B) Water contact angles of PCL, PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds. All values are presented as the averages (n = 3) ± standard deviation. (*: P < 0.05, **: P < 0.005, and ***: P < 0.0005).

3.4. In vitro degradation

The degradation profile of biomaterials is an essential aspect in the field of tissue engineering.⁵² An ideal scaffold shows degradation at a similar rate as that of new tissue would forms.⁵³ The degradation profiles of polymers in water-based environments typically consist of four distinct stages: hydration, reduced tensile strength resulting from the breaking of ester linkages, loss of mass, and ultimately, solubilization. Temperature, pH, molecular weight, and degree of crystallinity are key parameters that significantly affect the degradation process.⁵⁴ In this study, the degradation of scaffolds was observed by measuring weight loss and monitoring the trend of pH changes over a period of 63 days (Fig. 3A and B). The addition of Cs to the PCL polymer will accelerate the degradation process due to Cs's amorphous nature, which contrasts with the crystalline structure of PCL. Fig. 3A shows the different degradation processes that occurred in three sample groups throughout this time period. Alternatively, PCL/Cs membranes had an almost straight degradation pattern, whereas PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP saw the highest weight loss compared to PCL/Cs scaffolds. The rapid degradation of scaffolds containing L-Arg may be attributed to the hydrolysis of this hydrophilic agent. Despite all these changes in the weight of the 3D- printed scaffolds had maintained their integrity for 60 days. Similarly, López et al.⁵⁵ assessed the mechanical properties and degrading characteristics of the PCL/Cs scaffold. Their findings indicated that the inclusion of hydrophilic drugs resulted in an increased rate of hydrolytic the degradation compared to PCL/Cs scaffolds.

Fig. 3 The degradation process and mechanical behavior (A) degree of degradation according to weight loss of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds during 60 days of degradation, (B) PH change of control, PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds during 60 days of degradation, (C) stress/strain curve, (D) Young's modulus diagram of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP samples. All values are presented as the averages (n = 3) ± standard deviation. (*: P < 0.05, ***: P < 0.0005, and ****: P < 0.00005).

Typically, the pH decreases throughout the degradation process for PCL/Cs. The reason can be attributed to degradation products, such as the breakage of ester bonds by hydrolysis. The formation of monomers resulting in the separation of hydrogen ions from the carboxyl groups will concurrently lead to a reduction in pH. In this study, the PCL/Cs scaffold demonstrated a more significant change followed by an increase in the pH value, as shown in Fig. 3B. Following a duration of 63 days, the ultimate pH values were known 6.41 ± 0.3, 7.9 ± 0.5, and 8.3 ± 0.2 for scaffolds classified as PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP, respectively.

3.5. Mechanical behavior

It is necessary for GBR/GTR membranes to possess favorable mechanical properties, such as considerable strength and flexibility, is essential in order to endure the manipulation and placing into specific organs during implantation.⁵⁶Fig. 3C and D show the stress/strain curve and Young's modulus diagram for the PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP sample scaffolds. Among all evaluated samples, PCL/Cs has the lowest tensile strength, measuring at 1.36 ± 0.31 MPa. The presence of L-Arg in the PCL/Cs composite resulted in a considerable increase in both the strength (2.51 ± 0.41 MPa) and Young's modulus (14.62 ± 3.45 MPa). PCL/Cs-L-Arg/βTCP has the highest strength (6.71 ± 0.38 MPa) and Young's modulus (32.84 ± 4.11 MPa) compared to the other test samples. The enhanced mechanical features of the PCL/Cs scaffold can be attributed to the presence of L-Arg and the chemical interactions between PCL/Cs and L-Arg. In addition, the incorporation of βTCP in the PCL/Cs-L-Arg membrane resulted in an enhancement of the Young's modulus and stress strength, achieved by increasing the crystallite phases of the filament structure.⁵⁷ Similarly, Ye et al.⁵⁸ assessed the mechanical characteristics of polyhydroxyalkanoates (PHA) with different amounts of β-TCP that were developed using the fused deposition modeling (FDM) technology. The addition of β-TCP to PHA 3D filament resulted in a considerable improvement in compressive strength and Young's modulus.

Based on the findings of the current study and previous research, it can be concluded that the tensile strength and Young's modulus of PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP samples were within the appropriate range for their usage as GBR membranes and dental tissue engineering.⁵⁹

3.6. L-Arg release

The controlled release of bioactive drugs and chemicals is a crucial challenge in the field of biomaterials and tissue engineering, specifically in GBR membranes.⁷ Based on our understanding, L-Arg has a dual function with both positive and negative aspects in the field of tissue repair.³¹ Therefore, the level of L-Arg should be regulated by encapsulating it in a PCL/Cs composite in order to promote healing and minimize inflammatory reactions. Consequently, we investigated the release of L-Arg from the PCL/Cs-L-Arg/βTCP 3D scaffold in PBS with pH of 7.4 over a period of 21 days (as shown in Fig. 4A). A direct relationship may be shown between the cumulative drug release and time, suggesting that L-Arg exhibits sustained release characteristics. The rates that L-Arg is released are affected by factors such as the rate at which it degrades and its solubility in the medium. The reduced rate of L-Arg release from PCL/Cs-L-Arg/βTCP can be attributed to the existence of the PCL layer, which functions as a barrier to control the penetration and diffusion of the media and drug. In addition, the continuous degradation and decomposition of the PCL/Cs-L-Arg/TCP membrane may be a factor in the slower release.⁶⁰ The distinct release profile of L-Arg, characterized by an initial accumulation of release followed by a persistent release pattern, may be well-suited for promoting bone repair and regulating the immune response.¹⁶

Fig. 4 (A) In vitroL-Arg cumulative release of PCL/Cs-L-Arg/βTCP sample incubated in PBS (pH 7.4) for 21 days, ICP analysis of (B) Ca ions, and (C) P ions present in SBF solution during 28 days of immersion, and (D) SEM images of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds after 28 days immersion in SBF (scale bar: 20 μm). The results are reported as the means (n = 3) ± standard deviation.

3.7. Bioactivity assay

The evaluation of bioactivity is an essential stage in establishing the capacity of a chemical to promote bone repair.⁶¹ The main role of polymer characteristics is to significantly influence bioactivity behavior, especially in the mineralization process, by facilitating the deposition of minerals. Every polymer has the ability to influence the creation of hydroxyapatite crystals by chemically reacting with specific organic groups that possess a negative charge, such as carboxylates.⁶² This interaction occurs when the polymer is immersed in simulated bodily fluid (SBF), resulting in the deposition of hydroxyapatite on these organic groups. Therefore, in this work, PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP were submerged in this solution for a duration of 28 days. Afterwards, the production of hydroxyapatite on the materials was evaluated by ICP analysis. SBF refers to simulated body fluid, which is a solution with a high concentration of calcium and phosphate salts. The presence of an active component in this solution induces the formation of apatite-like crystals on the surface. The inclusion of βTCP in the scaffolds improves the development of crystal deposits by providing locations for mineral nucleation in the surrounding area.

Fig. 4B and C display the findings of the study on the levels of calcium and phosphorus ions in the SBF solution for 28 days. The analysis occurred on the scaffolds using inductively coupled plasma (ICP) analysis. Typically, it is anticipated that the levels of these ions in the surroundings will diminish gradually as a calcium phosphate layer develops on the surface.³⁰ The amount of calcium in the PCL/Cs-L-Arg/βTCP sample exhibited a substantial reduction in comparison to the other scaffolds after a duration of 28 days. The reduction can be ascribed to the surface conditions that facilitate the creation of a calcium phosphate layer, impeding the assimilation of many calcium ions on the surface. The calcium decrease in the PCL/Cs-L-Arg/βTCP samples is around 83%. The inclusion of βTCP in the PCL/Cs-L-Arg/βTCP samples results in a notable decrease in calcium levels in the SBF. This is attributed to the interaction between calcium ions and the surfaces of PCL/Cs-L-Arg/βTCP. The phosphorus content in the samples exhibited a substantial decline over a period of 28 days, with reductions of over 93% seen in the PCL/Cs-L-Arg/βTCP sample, as shown in Fig. 4C. The presence of βTCP in the samples' composition leads to a reduction in phosphorus concentration by interacting with phosphorus ions and PCL/Cs-L-Arg/βTCP surfaces. Previous research has documented a comparable phenomenon occurring when a hydroxyapatite layer is created by integrating βTCP into the composition.⁶³

As shown Fig. 4D, the SEM images show the presence of calcium phosphate particles in the PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP scaffolds. The addition of L-Arg and βTCP significantly promotes the formation of spherical apatite-like crystals, particularly at higher concentrations. Incorporating βTCP into the scaffold increases surface roughness, which helps facilitate crystal deposition by creating nucleation sites for minerals in the surrounding environment. βTCP is known for its bioactivity and biocompatibility, primarily due to its large surface area relative to its volume and its ability to promote bone growth (osteoconductivity).⁶⁴ These properties make it suitable for bone tissue engineering, as it can form a chemical bond with natural bone tissue by generating a physiologically active apatite layer. The SEM images suggest that βTCP plays a role in both the bone formation process in vivo and the formation of apatite-like crystals in vitro. The results confirm that PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP scaffolds have a bioactive surface capable of supporting hydroxyapatite formation.

3.8. Relative MG-63 cell viability

The vitality of MG-63 cells following interaction with the specimen scaffolds was assessed using the MTT assay. Fig. 5A illustrates the cell proliferation statistics when scaffolds are present. The relative cell viability (% control) showed a considerable increase after 3 days when treated with PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP. This increase in cell viability was preserved until day 5. L-Arg peptides are biocompatible substances that enhance cell proliferation and have no negative effects on cells.³¹ Similarly, βTCP has a beneficial effect on cell growth. The combination of these distinct components resulted in a synergistic enhancement of cell growth and viability.⁶⁵

Fig. 5 (A) Cell viability of MG-63 cells seeded on PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds at days 1, 3, and 5 and the achieved results were normalized against the control group (TCP), (B) Relative ALP activity of MG- 63 cells on the control group (TCP), PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds during 21 days, (C) CLSM images of MG-63 cells cultured on the PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP samples during 5 days (scale bar: 50 μm), and (D) Alizarin red staining of MG-63 cells seeded on the PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds on the 7th day of the cultivation. All values are presented as the averages (n = 3) ± standard deviation. (*P < 0.05, **P < 0.005, ***P < 0.0005, and ****P < 0.00005).

To confirm the MTT experiment, MG-63 cells were subjected to DAPI staining. The fluorescence microscopy images of cells stained with DAPI, as seen in Fig. 5B, provide crucial data on cellular viability and proliferation. The findings indicated that MG-63 cells exhibited strong development on the surfaces of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP membranes for a duration of 5 days. Fig. 5C clearly demonstrates that cell proliferation on PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP was much greater than on the PCL/Cs scaffold. This indicates that L-Arg plays an effective function in enhancing cell proliferation. The possible cause might be attributed to the production of L-Arg and NO, which may stimulate cell growth and improve attachment.¹⁸

3.9. Alkaline phosphatase assay

Alkaline phosphatase (ALP) activity, a crucial marker of osteoblast development, was assessed on days 7, 14, and 21. This enzyme is vital for bone formation as it hydrolyzes inorganic pyrophosphates, releasing calcium and phosphate ions necessary for the mineralization of the extracellular matrix.⁶⁶ The incorporation of roughness and surface chemical alteration as efficient strategies in the formation of bioscaffolds enhances the process of osteogenic differentiation.⁶⁷ Overall, MG-63 cells cultured with PCL/Cs-L-Arg/βTCP scaffold exhibited more ALP activity compared to cells cultured with PCL/Cs and PCL/Cs-L-Arg scaffolds throughout the incubation period (Fig. 5B). The ALP activity of PCL/Cs-L-Arg/βTCP nanocomposite scaffolds on days 7, 14, and 21 was about 1.3, 1.5, and 2 times more than that of PCL/Cs scaffolds. The observed differences were found to be statistically significant (p < 0.005). The hydrophilicity, cell proliferation of PCL/Cs-L-Arg/βTCP nanocomposite scaffolds were enhanced by modifications in surface chemistry, including βTCP as a bioactive agent. Collectively, these factors have the potential to greatly improve the process of cell osteogenic differentiation. Increased ALP activity is evidence of accelerated bone formation.⁶⁸

Previous study showed the beneficial impact of βTCP on ALP activity. The biological characteristics results clearly shown that the inclusion of βTCP into the polymer scaffolds could accelerate the differentiation of MG-63 cells and the generation of new bone.⁶⁹ Furthermore, scaffolds containing βTCP exhibited a greater amount of hydroxyapatite formation, which corresponded to the observed enhancement in the osteogenesis process.

Alizarin red staining was performed on the 7th day of MG-63 cell cultivation on the PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds, and the results are shown in Fig. 5D. This staining method highlights extracellular matrix (ECM) mineralization, indicating successful in vitro bone formation.⁷⁰ The deep red color observed in the images is due to the reaction between calcium ions and Alizarin red dye.⁷¹ As shown in the images, the intensity and number of calcium deposition nodules were significantly higher on the PCL/Cs-L-Arg/βTCP scaffolds compared to the PCL/Cs scaffold. This observation aligns with the cell attachment and alkaline phosphatase (ALP) secretion results, suggesting that the incorporation of L-Arg and βTCP enhances ECM mineralization and promotes osteoblast differentiation.

3.10. Immune responses and NO production

To investigate the inflammatory and anti-inflammatory effects of PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffolds, we measured the levels of IL-6, TNF-α, BMP-2, and TGF-β in macrophages using an ELISA kit. The cells were first stimulated with LPS (100 ng mL⁻¹) to induce the production of pro-inflammatory cytokines (IL-6 and TNF-α) before assessment. As shown in Fig. 6, there were no significant differences in the release of pro-inflammatory and anti-inflammatory cytokines from the PCL/Cs scaffolds compared to the control group, indicating no immunomodulatory effects. However, both PCL/Cs-L-Arg and PCL/Cs-L-Arg/βTCP scaffolds showed a decrease in pro-inflammatory cytokines (IL-6 and TNF-α), which may be due to the sustained release of L-Arg and subsequent nitric oxide (NO) production. In contrast, BMP-2 and TGF-β, osteogenic growth factors, showed a significant increase in activity, suggesting a shift toward anti-inflammatory responses. This pattern is likely linked to a higher proportion of M₂ macrophages relative to M₁ macrophages.³⁵ Notably, the PCL/Cs-L-Arg/βTCP scaffold exhibited a pronounced increase in the expression of anti-inflammatory cytokines (BMP-2 and TGF-β) compared to the other materials. He et al.⁷² conducted a study on the immunomodulation properties of a copolymer made of L-Arg-based poly(ester amine). They found that increasing the L-Arg content and adjusting the production of nitric oxide (NO) to around 9 μM resulted in higher expression of TGF-β1 and lower expression of TNF-α. This modulation of cytokine expression could potentially influence the behavior of different types of cells and facilitate the transition from an inflammatory state to an anti-inflammatory state at the site of damage. Nevertheless, our investigation demonstrated a large increase in the BMP-2 and TGF-β markers, whereas the IL-6 and TNF-α markers showed a notable decrease when PCL/Cs-L-Arg/βTCP was utilized (Fig. 6A–D). The results consistently showed that PCL/Cs-L-Arg/βTCP successfully stimulated the transformation of macrophages into the M₂ phenotype, regulated inflammatory responses, and enhanced bone regeneration.

Fig. 6 Macrophage cell responses to control group (TCP), PCL/Cs, PCL/Cs-L-Arg, and PCL/Cs-L-Arg/βTCP scaffold: In vitro immune response, ELISA analysis of macrophage cytokines (A) BMP-2, (B) TGF-β, (C) IL-6, (D) TNF-α during 48 h, (E) In vitro NO production by RAW-264.7 macrophage during 48 h, and (F) the schematic of the effect of NO and L-Arg release on immune cells response. The results are reported as the means (n = 3) ± standard deviation. (*P < 0.05, **P < 0.005, ***P < 0.0005, and ****P < 0.00005).

Several 3D-printed scaffolds have been studied to determine their ability to stimulate the generation of nitric oxide (NO). According to Fig. 6E, macrophages promptly reacted to the samples and produced NO. Following a 48 h incubation period, the release of NO increased from 1.26 ± 0.3 μM to 9.11 ± 1.1 μM when L-Arg was added to the PCL/Cs composite scaffold (Fig. 6E). The data exhibited a substantial increase compared to the control group (0.73 ± 0.1 μM). Macrophages are capable of metabolizing L-Arg through two distinct pathways, which include the generation of NO and an increase in cellular metabolism [52]. L-Arg additionally acts as the only substrate for the synthesis of NO by inducible nitric oxide synthase (iNOS). NO is well acknowledged as a powerful immunomodulatory agent produced by macrophages.³¹ It has a vital function in enhancing the development of bones and tissues by employing many processes, including the prevention of infections, the promotion of blood vessel formation, and the facilitation of the proliferation of cells.^18,73Fig. 6F shows that every result showed that NO and L-Arg of PCL/Cs-L-Arg/βTCP sample could successfully trigger macrophage polarization to M2 phenotype, moderate inflammatory responses. In conclusion, NO is recognized as a powerful immunomodulatory agent, facilitating the transition of macrophage polarization from the pro-inflammatory M1 phenotype to the anti-inflammatory M2 phenotype, thus creating a regenerative milieu. Specifically, NO generation has been demonstrated to increase the expression of critical osteogenic growth factors, including BMP-2 and TGF-β, which are vital for osteogenesis.⁷⁴ TGF-β regulates osteoblast development and bone matrix deposition via its Smad-dependent signaling pathways. Likewise, BMP-2 operates via the Smad signaling pathway to promote osteogenic differentiation and bone production. The elevation of TGF-β and BMP-2 activity noted in our work corroborates this approach, suggesting that the PCL/Cs-L-Arg/βTCP scaffold facilitates bone regeneration by diminishing inflammation and concurrently activating the signaling pathways essential for osteogenic differentiation.^75,76 This dual action emphasizes the significance of both TGF-β and BMP-2 in facilitating osteogenic differentiation, underscoring the scaffold's promise in clinical applications for massive bone defect repair and regeneration.

For future in vivo validation, studies should focus on assessing the degradation profiles of the PCL/Cs-L-Arg/βTCP scaffold, including its breakdown and interactions with surrounding tissues over time. Monitoring the rate of degradation and by-products is crucial, as these can impact tissue regeneration. Additionally, the foreign body response needs to be evaluated to ensure no significant inflammatory reactions or adverse effects. Finally, investigating the scaffold's integration with host tissue, particularly its interaction with bone and soft tissues, is vital for successful regeneration.

To expand the applications of this scaffold beyond guided bone regeneration (GBR), its potential in areas such as craniofacial reconstruction, long bone defects, and cartilage repair should be explored. The unique combination of PCL, chitosan, L-arginine, and β-tricalcium phosphate provides mechanical strength, biocompatibility, immune regulation, and osteogenic promotion, making it suitable for complex applications like craniofacial abnormalities. Additionally, the scaffold's ability to deliver bioactive compounds for localized drug release offers potential for improved tissue repair with minimal systemic side effects. This adaptability also opens possibilities for cartilage repair, broadening its potential in regenerative medicine. Thus, this scaffold shows great promise for advancing tissue engineering and regenerative therapies.

4. Conclusion

This study introduces a novel 3D-printed scaffold that combines polycaprolactone, chitosan, L-arginine, and β-tricalcium phosphate for guided bone regeneration (GBR). The unique synergy between these components, coupled with extrusion-based 3D printing, enhances the scaffold's mechanical strength, bioactivity, and sustained drug release, promoting an optimal environment for bone repair. The scaffold's ability to modulate macrophage cytokine release, to foster an anti-inflammatory response, and to enhance osteogenic properties, underscore its potential for bone regeneration. The integration of L-arginine and β-TCP represents a significant advancement in scaffold design, making this composite an innovative tool in regenerative dentistry and orthopedic treatments. Our findings demonstrate the capability of 3D printing to create highly customized biomaterials, paving the way for future advancements in bone regeneration therapies.

Authors’ contributions

Mohammad Hosseini Hooshiar: investigation, writing – original draft, visualization. Negin Ostadsharifmemar: investigation, writing – original draft, methodology. Tohid Javaheri: investigation, writing – original draft, methodology. Negin Salehinia: investigation, writing – original draft, software. Melika Golozar: conceptualization, editing draft, software. Ensieh Sagheb Sadeghi: writing – original draft, methodology. Atefeh Zamani: conceptualization, formal analysis, editing draft. Parisa Heydari: writing – original draft, project administration, investigation, visualization, formal analysis, supervision. Ali zarabbi: project administration, reviewing and editing original draft, investigation, supervision. Mohammad Mahdevar: project administration, investigation, visualization, funding acquisition.

Ethics statement

All methods were executed in compliance with pertinent guidelines and regulations, and all experiments adhered to established standards.

Data availability

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

Conflicts of interest

The authors declare no conflict of interest.

References

1. C. Xu, F. Wu, J. Yang, H. Wang, J. Jiang, Z. Bao, X. Yang, G. Yang, Z. Gou and F. He, Chem. Eng. J., 2022, 450, 138003.
2. F. Cestari, M. Petretta, Y. Yang, A. Motta, B. Grigolo and V. M. Sglavo, Sustainable Mater. Technol., 2021, 29, e00318.
3. D. Bahati, M. Bricha, A. Semlali and K. El Mabrouk, Mater. Chem. Phys., 2024, 129637.
4. Z. Shu, C. Zhang, L. Yan, H. Lei, C. Peng, S. Liu, L. Fan and Y. Chu, Int. J. Biol. Macromol., 2023, 224, 1040–1051.
5. J. Liu, Q. Zou, C. Wang, M. Lin, Y. Li, R. Zhang and Y. Li, Mater. Des., 2021, 210, 110047.
6. G. L. Abe, J.-I. Sasaki, C. Katata, T. Kohno, R. Tsuboi, H. Kitagawa and S. Imazato, Dent. Mater., 2020, 36, 626–634.
7. H. Shakeri, M. H. Nazarpak, R. Imani and L. Tayebi, Int. J. Biol. Macromol., 2023, 231, 123201.
8. B. S. Heidari, J. M. Dodda, L. K. El-Khordagui, M. L. Focarete, P. Maroti, L. Toth, S. Pacilio, S. E. El-Habashy, J. Boateng and O. Catanzano, Acta Biomater., 2024, 184, 1–22.
9. B. Allan, R. Ruan, E. Landao-Bassonga, N. Gillman, T. Wang, J. Gao, Y. Ruan, Y. Xu, C. Lee and M. Goonewardene, Tissue Eng., Part A, 2021, 27, 372–381.
10. S.-M. Solomon, I.-G. Sufaru, S. Teslaru, C. M. Ghiciuc and C. S. Stafie, Appl. Sci., 2022, 12, 1042.
11. X. Wang, P. Shen, N. Gu, Y. Shao, M. Lu, C. Tang, C. Wang, C. Chu, F. Xue and J. Bai, ACS Biomater. Sci. Eng., 2023, 10, 537–549.
12. A. H. Mahmoud, Y. Han, R. Dal-Fabbro, A. Daghrery, J. Xu, D. Kaigler, S. B. Bhaduri, J. Malda and M. C. Bottino, ACS Appl. Mater. Interfaces, 2023, 15, 32121–32135.
13. X. Zhao, S. Wang, F. Wang, Y. Zhu, R. Gu, F. Yang, Y. Xu, D. Xia and Y. Liu, J. Magnesium Alloys, 2024, 12, 966–979.
14. Z. Liu, J. Shi, L. Chen, X. He, Y. Weng, X. Zhang, D.-P. Yang and H. Yu, Int. J. Biol. Macromol., 2024, 133172.
15. X. Niu, L. Wang, M. Xu, M. Qin, L. Zhao, Y. Wei, Y. Hu, X. Lian, Z. Liang and S. Chen, Carbohydr. Polym., 2021, 260, 117769.
16. L. Huang, J. Zhang, X. Liu, T. Zhao, Z. Gu and Y. Li, Chin. Chem. Lett., 2021, 32, 234–238.
17. Y. Zhou, Z. Gu, J. Liu, K. Huang, G. Liu and J. Wu, Carbohydr. Polym., 2020, 230, 115640.
18. H.-P. Bei, X. Ji, T. Xu, Z. Chen, C.-H. Lam, X. Zhou, Y. Yang, Y. Zhang, C. Wen and Y. Liu, Composites, Part B, 2024, 278, 111410.
19. G. A. Khoman, M. H. A. Kalijaga, N. Aisah, R. Fidyaningsih, J. Raharjo, O. P. Arjasa and E. Prajatelistia, R. Soc. Open Sci., 2024, 11, 231694.
20. M. Ezati, H. Safavipour, B. Houshmand and S. Faghihi, Prog. Biomater., 2018, 7, 225–237.
21. T. Tanaka, H. Komaki, M. Chazono, S. Kitasato, A. Kakuta, S. Akiyama and K. Marumo, Morphologie, 2017, 101, 164–172.
22. S. Liao, W. Wang, M. Uo, S. Ohkawa, T. Akasaka, K. Tamura, F. Cui and F. Watari, Biomaterials, 2005, 26, 7564–7571.
23. J.-H. Shim, J.-Y. Won, J.-H. Park, J.-H. Bae, G. Ahn, C.-H. Kim, D.-H. Lim, D.-W. Cho, W.-S. Yun and E.-B. Bae, Int. J. Mol. Sci., 2017, 18, 899.
24. W. L. Ng, J. An and C. K. Chua, Engineering, 2024, 36, 146–166.
25. K. Kim, Y. Su, A. J. Kucine, K. Cheng and D. Zhu, ACS Biomater. Sci. Eng., 2023, 9, 5457–5478.
26. Z. Ma, X. Hu, X. Li, Q. An, Y. Zhang, C. Guo, Y. Zhao and Y. Zhang, ACS Biomater. Sci. Eng., 2024, 10, 3984–3993.
27. J. Uddin, R. Dubey, V. S. Balasubramaniam, J. Kabel, V. Khare, Z. Salimi, S. Sharma, D. Zhang and Y. K. Yap, Micromachines, 2024, 15, 349.
28. L. Biglari, M. Naghdi, S. A. Poursamar, M. R. Nilforoushan, A. Bigham and M. Rafienia, Int. J. Biol. Macromol., 2024, 258, 128716.
29. S. M. R. Hosseini, P. Heydari, R. N. Azadani, S. Iravani and A. Zarrabi, Emergent Mater., 2024, 1–12.
30. R. N. Azadani, S. Karbasi and A. Poursamar, Int. J. Biol. Macromol., 2024, 129407.
31. P. Heydari, M. Kharaziha, J. Varshosaz, A. Z. Kharazi and S. H. Javanmard, Biomater. Adv., 2024, 213762.
32. K. Shobana and S. Swamiappan, Inorg. Chem. Commun., 2023, 148, 110347.
33. P. Heydari, J. Varshosaz, M. Kharaziha and S. H. Javanmard, J. Mater. Sci.: Mater. Med., 2023, 34, 16.
34. S. M. R. Hosseini, P. Heydari, M. Namnabat, R. N. Azadani, F. A. Gharibdousti, E. M. Rizi, A. Khosravi, A. Zarepour and A. Zarrabi, Eur. J. Pharmacol., 2024, 176671.
35. S. Farshid, M. Kharaziha, H. Salehi and M. Ganjalikhani Hakemi, ACS Appl. Mater. Interfaces, 2023, 15, 48996–49011.
36. P. Jain, R. Y.-T. Lin, K. Mishra, H. Handral and N. Dubey, Dent. Mater., 2024, 40, 151–157.
37. M. Arastouei, M. Khodaei, S. M. Atyabi and M. J. Nodoushan, Mater. Today Commun., 2021, 27, 102176.
38. P. Heydari, A. Zargar Kharazi and L. Shariati, Sci. Rep., 2024, 14, 12019.
39. Z. Ling, Z. Chen, J. Deng, Y. Wang, B. Yuan, X. Yang, H. Lin, J. Cao, X. Zhu and X. Zhang, Chem. Eng. J., 2021, 420, 130302.
40. S. Ghadirian and S. Karbasi, Int. J. Biol. Macromol., 2023, 233, 123651.
41. H. Zarharan, M. Bagherian, A. S. Rokhi, R. R. Bajgiran, E. Yousefi, P. Heravian, M. N. Khazrabig, A. Es-haghi and M. E. T. Yazdi, Arab. J. Chem., 2023, 16, 104806.
42. E. Archer, M. Torretti and S. Madbouly, Phys. Sci. Rev., 2023, 8, 4391–4414.
43. C. S. Choy, W. F. Lee, P. Y. Lin, Y.-F. Wu, H.-M. Huang, N.-C. Teng, Y.-H. Pan, E. Salamanca and W.-J. Chang, Sci. Rep., 2021, 11, 9234.
44. A. Massit, A. El Yacoubi, A. Kholtei and B. C. El Idrissi, Biointerface Res. Appl. Chem., 2021, 11, 8034–8042.
45. T. S. Putri, A. Ratnasari, N. Sofiyaningsih, M. S. Nizar, A. Yuliati and K. A. Shariff, Ceram. Int., 2022, 48, 11428–11434.
46. X. Cui, J. Zhang, Y. Qian, S. Chang, B. J. Allardyce, R. Rajkhowa, H. Wang and K.-Q. Zhang, Engineering, 2024, 46, 92–108.
47. I. Koumentakou, A. Michopoulou, M. J. Noordam, Z. Terzopoulou and D. N. Bikiaris, J. Mater. Sci., 2024, 1–21.
48. F. Mukasheva, M. Moazzam, B. Yernaimanova, A. Shehzad, A. Zhanbassynova, D. Berillo and D. Akilbekova, Bioprinting, 2024, 39, e00341.
49. S. Tanpichai, Y. Srimarut, W. Woraprayote and Y. Malila, Int. J. Biol. Macromol., 2022, 213, 534–545.
50. S. W. Lee, O. Park, H. T. Rho and S. Kim, J. Korean Ceram. Soc., 2023, 60, 811–816.
51. M. A. Wsoo, S. I. Abd Razak, S. P. M. Bohari, S. Shahir, R. Salihu, M. R. A. Kadir and N. H. M. Nayan, Int. J. Biol. Macromol., 2021, 181, 82–98.
52. U. Meyer, T. Meyer, J. Handschel and H. P. Wiesmann, Fundamentals of tissue engineering and regenerative medicine, Springer, 2009.
53. P. Heydari, S. Parham, A. Z. Kharazi, S. H. Javanmard and S. Asgary, Fibers Polym., 2022, 1–10.
54. V. Mkhabela and S. S. Ray, Int. J. Biol. Macromol., 2015, 79, 186–192.
55. I. López-González, A. B. Hernández-Heredia, M. I. Rodríguez-López, D. Auñón-Calles, M. Boudifa, J. A. Gabaldón and L. Meseguer-Olmo, Pharmaceutics, 2023, 15, 1763.
56. J.-Y. Kim and J.-B. Park, Coatings, 2022, 12, 1059.
57. W. A. R. Neto, I. H. L. Pereira, E. Ayres, A. C. C. de Paula, L. Averous, A. M. Góes, R. L. Oréfice and R. E. S. Bretas, Polym. Degrad. Stab., 2012, 97, 2037–2051.
58. X. Ye, Y. Zhang, T. Liu, Z. Chen, W. Chen, Z. Wu, Y. Wang, J. Li, C. Li and T. Jiang, Int. J. Biol. Macromol., 2022, 209, 1553–1561.
59. S. Cao, J. Han, N. Sharma, B. Msallem, W. Jeong, J. Son, C. Kunz, H.-W. Kang and F. M. Thieringer, Materials, 2020, 13, 3057.
60. A. Behl, V. S. Parmar, S. Malhotra and A. K. Chhillar, Polymer, 2020, 207, 122901.
61. S. Skibiński, J. P. Czechowska, M. Guzik, V. Vivcharenko, A. Przekora, P. Szymczak and A. Zima, Sustainable Mater. Technol., 2023, 38, e00722.
62. J. A. Lett, S. Sagadevan, I. Fatimah, M. E. Hoque, Y. Lokanathan, E. Léonard, S. F. Alshahateet, R. Schirhagl and W. C. Oh, Eur. Polym. J., 2021, 148, 110360.
63. M. Duan, S. Ma, C. Song, J. Li and M. Qian, Ceram. Int., 2021, 47, 4775–4782.
64. M. C. B. Melo, B. R. Spirandeli, L. Barbosa, V. R. Dos Santos, T. M. B. de Campos, G. P. Thim and E. de Sousa Trichês, J. Mech. Behav. Biomed. Mater., 2025, 163, 106850.
65. Z. Javkhlan, S.-H. Hsu, R.-S. Chen and M.-H. Chen, J. Formosan Med. Assoc., 2024, 123, 71–77.
66. P. Jin, L. Liu, L. Cheng, X. Chen, S. Xi and T. Jiang, Biomed. Eng. Online, 2023, 22, 12.
67. P. M. Sivakumar, A. A. Yetisgin, S. B. Sahin, E. Demir and S. Cetinel, Environ. Res., 2023, 238, 117131.
68. S. Kumari, P. Srivastava and A. Mishra, J. Porous Mater., 2023, 30, 1085–1099.
69. Y.-F. Wu, Y.-T. Wen, E. Salamanca, L. M. Aung, Y.-Q. Chao, C.-Y. Chen, Y.-S. Sun and W.-J. Chang, J. Dent. Sci., 2024, 19, 1116–1125.
70. P. Zhou, Y. Xia, X. Cheng, P. Wang, Y. Xie and S. Xu, Biomaterials, 2014, 35, 10033–10045.
71. S. Shrestha, B. K. Shrestha, S. W. Ko, R. Kandel, C. H. Park and C. S. Kim, Carbohydr. Polym., 2021, 251, 117035.
72. M. He, L. Sun, X. Fu, S. P. McDonough and C.-C. Chu, Acta Biomater., 2019, 84, 114–132.
73. Z. Zhou, Y. Liu, W. Li, Z. Zhao, X. Xia, J. Liu, Y. Deng, Y. Wu, X. Pan and F. He, Adv. Healthcare Mater., 2024, 13, 2302153.
74. C. Bogdan, Nat. Immunol., 2001, 2, 907–916.
75. M. S. Rahman, N. Akhtar, H. M. Jamil, R. S. Banik and S. M. Asaduzzaman, Bone Res., 2015, 3, 1–20.
76. F. Guerrero, C. Herencia, Y. Almaden, J. M. Martínez-Moreno, A. Montes de Oca, M. E. Rodriguez-Ortiz, J. M. Diaz-Tocados, A. Canalejo, M. Florio and I. López, PLoS One, 2014, 9, e89179.

---