Polymethyl methacrylate–hydroxyapatite: A viable non-metallic alternative for dental implants? Comprehensive surface characterization

Abstract

Objectives

The surface characteristics of implants made from polymethyl methacrylate (PMMA) and hydroxyapatite (HA) were evaluated in this study.

Methods

In vitro experiments were carried out using four types of screw implants fabricated from different materials: PMMA–HA, PMMA coated with HA (PMMA–HAc), PMMA, and sandblasted, large-grit, acid-etched titanium (SLA-Ti). Several characterization procedures were performed in this study. X-ray diffraction was applied to verify the incorporation of HA into the PMMA matrix by analyzing the crystalline structure and phase composition. Atomic force microscopy was used to measure surface roughness. Scanning electron microscopy (SEM) visualized pore sizes in detail. Contact angles were measured to evaluate the wettability of the materials.

Results

SLA-Ti had the highest crystallinity value (90 %). By contrast, the polymer-based samples exhibited semicrystalline properties, with 22 % for PMMA–HAc, 16 % for PMMA–HA, and 12 % for PMMA. SLA-Ti also exhibited the greatest surface roughness (2.020 nm), followed by PMMA–HAc (220 nm), PMMA–HA (128 nm), and PMMA (108 nm), which had noticeably smoother surfaces. PMMA–HAc had the largest average pore diameter (141.07 ± 217.56 μm), with progressively smaller pores observed in PMMA–HA (30.86 ± 24.13 μm), PMMA (18.30 ± 9.58 μm), and SLA-Ti (3.97 ± 1.05 μm). Wettability analysis showed that PMMA–HA exhibited the most hydrophilic behavior, with the lowest contact angle (47.28°). The other materials exhibited lower hydrophilicity, with contact angles of 74.53° for PMMA–HAc, 75.70° for PMMA, and 84.36° for Ti.

Conclusions

The surface characteristics of dental implants composed of PMMA–HA, PMMA–HAc, PMMA, and SLA-Ti were investigated in this study to assess their potential for creating a favorable environment for osseointegration.

Introduction

Dental implants are the most effective option for replacing missing teeth to address a major global health issue that can greatly reduce a person's quality of life.1, 2, 3 However, the primary materials used, i.e., titanium and its alloys, still have several important limitations. Poor bonding with bone and gingival tissue has often been reported for titanium implants produced without appropriate surface modification. This inadequate attachment can weaken the stability of implants and increase the likelihood of infection and inflammation in the peri-implant area, ultimately leading to implant failure.⁴ Furthermore, metallic implants are associated with limitations such as hypersensitivity, reduced bone cell recruitment, and aesthetic compromise due to the potential visibility of metal through thin tissue.^5,6 Consequently, these limitations have motivated the search for alternative biomaterials by optimizing surface modifications to ensure clinical success.

Among the various types of biomaterials, polymeric biomaterials have gained popularity due to their many advantages.⁷ In particular, polymethyl methacrylate (PMMA, [-CH₂-C(CH₃) (COOCH₃)-]_n) is a lightweight synthetic polymer characterized by tunable mechanical properties, adjustable porosity, softness, the absence of electrolytic currents, and favorable aesthetic qualities. PMMA is used extensively in clinical dentistry for creating various dental prostheses, such as denture bases, liners, reliners, repair materials, artificial teeth, temporary crowns, bridges, implant screws, and obturators.⁸ In addition, PMMA has potential as an implant material due to its thermal and electrical insulation, resistance to biodegradation, and surface characteristics that facilitate cell adhesion.^5,9, 10, 11 However, PMMA is a bioinert material that lacks the biological activity required to establish efficient osseointegration, thereby leading to inferior interfacial strength between the implant and host bone tissue compared with bioactive materials. In addition to these biological limitations, PMMA has mechanical shortcomings, particularly its low mechanical strength, modulus of elasticity, and fracture toughness.^12,13 Thus, the incorporation of hydroxyapatite (HA) has been investigated to overcome these issues. This modification is expected to reinforce the mechanical integrity but the primary aim is to provide the appropriate surface characteristics required for biological interaction.

HA is an inorganic mineral defined by the chemical formula Ca₁₀(PO₄)₆(OH)₂.⁵ It is known for its excellent osteoconductive properties in bone tissue repair, as well as its biocompatibility, bioactivity, and non-inflammatory nature, allowing it to integrate seamlessly with surrounding bone tissue without eliciting adverse reactions.¹⁴ Moreover, HA is beneficial for enhancing the biological integration of implants through improved osteoblast attachment and increased surface hydrophilicity.¹⁵ Furthermore, the incorporation of HA into the PMMA matrix aligns with current advancements in additive manufacturing, where polymer–ceramic composites are utilized for enhancing mechanical strength and biological interaction. The composite approach also serves as a foundational step toward four-dimensional printing applications, offering the potential for smart implants that can dynamically adapt to the physiological environment post-implantation.^16,17

The PMMA–HA blending ratio critically affects the properties of composites.¹⁸ Previous studies have shown that HA loadings within the range of 10–20 wt% enhance surface roughness, wettability, and bioactivity, while also minimizing particle agglomeration and the structural discontinuities that are commonly reported with higher HA contents.19, 20, 21 It is hypothesized that the addition of HA at 16.2 wt%, which falls within this optimal range, may yield a PMMA composite with favorable surface properties for dental implants. To validate this hypothesis, in the present study, the crystallinity, roughness, pore size, and wettability were assessed as critical early indicators of an implant's biological potential. Consequently, PMMA–HA implants with this specific ratio were synthesized and the surface characteristics were evaluated to determine their potential uses as non-metallic options for dental implants.

Materials and Methods

Study design

This study was based on in vitro laboratory experiments with a comparative control group design. The sample groups consisted of a PMMA–HA composite (83.8:16.2 ratio), PMMA coated with HA, pure PMMA, and titanium control. Commercially available sandblasted, large-grit, acid-etched titanium (SLA-Ti) implants (NeoBiotech IS-II, NeoBiotech Co., South Korea) were employed as the control group because they are widely used and highly regarded in dental implantology due to their excellent biological, physical, and mechanical properties.⁵ Samples were prepared at the Research Center of the Faculty of Dental Medicine, Universitas Airlangga, Indonesia, and the coating process was performed at the Physics Laboratory, Faculty of Science and Technology, Universitas Airlangga. All characterization tests were conducted at the Faculty of Advanced Technology and Multidiscipline (FTMM), LIHTR UNAIR, and the Institut Teknologi Sepuluh Nopember (ITS).

Extraction and preparation of HA

Bovine teeth were selected as the source of HA due to their high crystallinity (82.5–86.3 %), structural similarity to human teeth, ready availability, and favorable particle morphology. Bovine-derived HA also has physicochemical characteristics, including trace ions such as Mg, Na, and Sr, that are biologically advantageous compared with standard synthetic HA, making it a more suitable and cost-effective material for evaluating the surface properties of PMMA–HA.^22,23

In the extraction process, bovine teeth were first extracted from the jaw and thoroughly cleaned, before boiling at 100 °C for 3 h, where the water was replaced every hour. The teeth were then dried and softened using pressure treatment at 1.2 psi for 1 h, before changing the water and washing. This pressure–washing cycle was repeated twice. The softened teeth were then dried in an oven at 400 °C for 24 h, before ultrasonic treatment in alcohol for 3 h, where the alcohol was replaced every hour to remove residual fats. Subsequently, the samples were re-dried at 400 °C for 24 h. Calcination was then performed at 1000 °C for 2 h in a muffle furnace. After calcination, the teeth were ground and sieved through an 80-mesh filter to obtain HA powder. The final powder appeared white to off-white and the particle sizes ranged from 400 to 650 nm.²⁴

Preparation of PMMA samples

PMMA implant samples (Cemex® System Genta, Tecres, Italy) were prepared by mixing 0.4 g of PMMA powder with 400 μL of monomer to form PMMA cement. The mixture was stirred carefully to reduce the formation of air bubbles and poured into a three-dimensional (3D) cylindrical silicone mold with a diameter of 5 mm and height of 10 mm. The samples then were left to cure at room temperature, before trimming and storing in a clean setting with an average relative humidity of about 40 % at 20 °C.

Preparation of PMMA–HA composite samples

PMMA was mixed with bovine-derived HA using a planetary mixer for 6 h. The monomer was added at a volume of 0.016 mL for every 0.1 g of PMMA–HA powder to obtain a ratio of 83.8:16.2. This mixture was stirred for 1–1.5 min until it reached a dough-like consistency and subsequently poured into a silicone mold (5 mm diameter × 10 mm height). The material was allowed to set for 5–10 min, before removing from the mold. After removal, the samples were sterilized by following a series of steps, which involved rinsing with phosphate-buffered saline, immersing in 70 % ethanol for 2 h, and exposing to ultraviolet light for 2 h.²⁵ Ultraviolet irradiation was applied as a non-thermal sterilization approach to decontaminate the surfaces of PMMA-based materials, which are known to be susceptible to heat-related damage.²⁶

Coating PMMA implant surfaces with HA

PMMA samples were sterilized by soaking in 70 % ethanol, rinsing thoroughly with distilled water, and air drying at room temperature. The coating mixture was prepared by dissolving 1.4 g of HA and 0.6 g of gelatin (ratio of 70:30) in 6 mL of distilled water and 1 mL of 96 % ethanol, before continuous stirring for 24 h using a magnetic stirrer. Gelatin was incorporated as an organic binder to enhance the stability and homogeneity of the slurry, thereby facilitating uniform HA dispersion and contributing to the formation of surface microporosity. The ratio of 70:30 was selected based on previous research, which demonstrated that this formulation achieved an optimal coating thickness of 40 μm. Furthermore, the polar functional groups of gelatin (-NH₂, –COOH, and –OH) are known to enhance interfacial adhesion and modify the surface chemistry of composite coatings.27, 28, 29

The samples were then immersed for 10 min in the coating solution using a dip-coating apparatus, which operated at a constant voltage of 9 V. After immersion, the samples were lifted vertically at a steady speed of 0.198 mm/s while keeping the setting constant at 9 V to ensure that a uniform and firmly adhered coating was obtained. Finally, the coated samples were dried at room temperature in a well-ventilated area to prevent contamination and surface cracking.²⁹

X-ray diffraction (XRD)

XRD analysis was conducted using a diffractometer (Rigaku Miniflex 600-C, Rigaku, Japan) equipped with Cu radiation and a nickel filter, which operated at a scan speed of 2°/min across a 2θ range of 10°–60°.^30,31 Crystal phases were identified using JCPDS reference code 00-009-0432. The degree of crystallinity was determined by dividing the crystalline area by the total diffraction area and multiplying the value by 100 %.³² Crystallinity was evaluated using Origin software and the percentage crystallinity was calculated with the following formula:

$$\frac{TotalCrystalArea}{TotalArea\left(TotalCrystalArea+AmorphusArea\right)}\times 100\%\text{.}$$

Surface roughness measurement with atomic force microscopy (AFM)

Surface roughness was assessed for the PMMA samples using an atomic force microscope (Bruker N8-NEOS, Bruker, Germany), which operated in non-contact mode to avoid any surface damage. Before scanning, the samples were cleaned with isopropyl alcohol. An area of 10 μm × 10 μm was scanned at a rate of 1 s per line using an etched silicon cantilever tip while maintaining an approximate tip–sample distance of 10 nm. AFM analysis generated 3D surface morphology images, which were subsequently processed using dedicated software to calculate the arithmetic average roughness (Ra). This parameter represents the mean deviation of the surface height from the central plane, and is reported in nanometers.³³

Pore size measurement with scanning electron microscopy (SEM)

A scanning electron microscope (Phenom Pro X, Thermo Fisher Scientific, Netherlands) was used to evaluate the pore size and surface morphology of the samples. Before imaging, the samples were placed on carbon tape and sputter-coated with gold (Au) for 3 s using a sputter coater (Quorum Technologies, United Kingdom) after undergoing a vacuum process for 30 min. SEM images were obtained at magnifications of 100 × , 700 × , and 1000 × . Pore diameters were measured using ImageJ software by selecting the largest pore within each field of view to ensure that the analysis was consistent.³⁴

Wettability analysis

Wettability was analyzed based on contact angle measurements obtained under room temperature and ambient laboratory conditions. The test solution was prepared as a simulated body fluid (SBF) by dissolving 58.43 g of NaCl, 2.77 g of CaCl₂, and 1.39 g of NaHPO₄·H₂O (Sigma–Aldrich) in 1 L of deionized water. To acquire each measurement, the sample was positioned with its upper surface parallel to a flat base, and an approximately 3 μL droplet of the SBF solution was dispensed onto the surface using a syringe. A side-view image was captured for each droplet using a camera positioned horizontally and aligned with the sample plane. Images were recorded 10 s after depositing the droplet and analyzed using ImageJ software. The software measured the contact angle formed between the droplet and the material's surface, which was then used to classify wettability according to the following criteria: perfect wetting (0°), high wettability (0° < contact angle <90°), and low wettability (90° < contact angle <180°).³⁵ These measurements allowed the determination of each sample's affinity for water, indicating whether the surface was hydrophilic or hydrophobic.

Data analysiss

Statistical analyses were performed using the Statistical Package for the Social Sciences (SPSS) version 25 (IBM Corporation, USA). Data normality was examined with the Shapiro–Wilk test, followed by Levene's test to evaluate the homogeneity of variances. When the data satisfied the assumptions of normality and homogeneity, one-way analysis of variance was conducted to compare the differences in means between groups. Tukey's post hoc test was subsequently applied to identify pairwise differences, and statistically significant differences were accepted at p < 0.05.

Results

Degree of crystallinity

The differences in the XRD peaks obtained for the PMMA, PMMA–HA, and HA coating samples are shown in Figure 1. For PMMA, the characteristic XRD peaks were obtained at angles (2θ) of 19°, 20°, 30°, and 45°, where these peaks were predominant, moderately diffused, and broad. The PMMA–HA sample produced characteristic Bragg diffraction peaks at approximately 2θ = 25.9°, with strong peak regions at 31.8°–33°, and at around 39°, 46.7°, and 49.5°. These diffraction patterns indicated the crystalline arrangement of HA integrated into the material. SLA-Ti, which served as the control, produced distinct sharp diffraction peaks at 2θ angles of 35° and 40°, confirming its strong crystalline characteristics and high degree of crystallinity. The findings showed that PMMA was characterized by the lowest crystallinity (12 %), followed by PMMA–HA (16 %) and PMMA–HAc (22 %), whereas SLA-Ti exhibited the highest crystallinity (90 %) (Table 1).

Figure 1.

XRD patterns obtained for PMMA, PMMA–HAc, PMMA–HA, and SLA-Ti. (∗) Peak labels indicate crystalline phase origins.

Table 1.

Degrees of crystallinity for dental implant materials.

Dental Implant Material	PMMA	PMMA–HAc	PMMA–HA	Ti
Degree of crystallinity (%)	12 %	22 %	16 %	90 %

Surface roughness

The Ra measurements obtained for the dental implant materials are presented in Table 2. The findings showed that PMMA–HAc had the highest Ra value, followed by PMMA–HA, and PMMA exhibited the lowest surface roughness. These results suggest that the mixing and coating procedures involving HA led to significant topographical changes in the PMMA-based implants. Furthermore, the surface roughness of the SLA-Ti implant material was reported as 2.02 μm (2020 nm) by the manufacturer.

Table 2.

Surface roughness parameter (Ra) mean value ± standard deviation for test materials.

Implant Biomaterial	PMMA	PMMA–HAc	PMMA–HA	SLA-Ti
Ra	108 nm ± 11 nm	220 nm ± 65 nm	128 nm ± 19 nm	2,020 nm

Pore size

The pore diameter was relatively small in PMMA and pores were uniformly distributed across the material's surface. By contrast, PMMA–HAc had larger pores with substantial variability, including several extreme outliers. PMMA–HA had a more moderate pore distribution, characterized by a combination of both small and large pores. Among all materials, SLA-Ti had the smallest and most uniform pore structure. The largest and most variable pore sizes were found in PMMA–HAc, followed by PMMA–HA, PMMA, and SLA-Ti (Figure 4, Table 3).

Figure 4.

SEM images of dental implant materials: (a) PMMA (350 × ), (b) PMMA–HAc (700 × ), (c) PMMA–HA (500 × ), and (d) SLA-Ti (1000 ×).

Table 3.

Pore size (mean ± standard deviation (SD)) and Tukey's honestly significant difference test results (p-value) between groups.

Group	Pore Size (μm) (Mean ± SD)	PMMA	PMMA–HAc	PMMA–HA	SLA-Ti
PMMA	18.3±9.58
PMMA–HAc	141.07±217.56	< 0.001a
PMMA–HA	30.86±24.13	0.974	0.002a
SLA-Ti	3.97±1.05	0.964	< 0.001a	0.804

^ap < 0.05 shows statistically significant difference.

Wettability contact angle measurements

Figure 5 shows the wettability profiles obtained for the PMMA, PMMA–HAc, PMMA–HA, and SLA-Ti dental implants. Among the materials, PMMA–HA exhibited the highest hydrophilicity, as indicated by the lowest mean contact angle (47.28°), followed by PMMA–HAc (74.53°) and PMMA (75.70°), whereas SLA-Ti exhibited the greatest hydrophobicity with an average contact angle of 84.36° (Table 4). The data satisfied the assumptions of normality and homogeneity of variance (p > 0.05). Post hoc analysis (Table 4) revealed that the contact angle obtained for PMMA–HA differed significantly from those for the other materials (p < 0.05), except there was no significant difference compared with that for PMMA–HAc (p = 0.998).

Figure 5.

Wettability assessments for dental implant materials using SBF: (a) PMMA, (b) PMMA–HAc, (c) PMMA–HA, and (d) SLA-Ti.

Table 4.

Contact angles (mean ± standard deviation (SD)) and post hoc test comparisons (p-value) between groups.

Group	Contact Angle (°) (Mean ± SD)	PMMA	PMMA–HAc	PMMA–HA	SLA-Ti
PMMA	75.70 ± 5.64
PMMA–Hac	74.53 ± 5.65	0.998
PMMA–HA	47.28 ± 8.14	0.011a	0.014a
SLA-Ti	84.36 ± 11.36	0.577	0.482	0.002a

^ap < 0.05 denotes a statistically significant difference.

Discussion

The surface attributes of dental implants play vital roles in ensuring appropriate integration with the surrounding tissues because they directly affect the osseointegration process.³⁶ In the present study, XRD analysis was applied to confirm the incorporation of HA into the PMMA matrix by examining the crystalline structure of the implant materials. XRD is a fundamental analytical method in materials science for providing critical insights into factors such as the crystallite size and phase composition within a material.³⁷

XRD analysis showed that PMMA produced broad and diffuse peaks at 2θ angles of 19°, 20°, 30°, and 45°, indicating its predominantly amorphous nature with very low crystallinity (12 %). These findings are consistent with earlier literature, which describes PMMA as a polymer with minimal crystalline regions due to its molecular chain structure.³⁸ By contrast, the addition of HA induced a shift toward a semicrystalline structure. The PMMA–HA sample produced characteristic Bragg diffraction peaks at approximately 2θ = 25.9°, 31.8°–33°, 39°, 46.7°, and 49.5°. Similarly, the PMMA–HAc sample produced HA-associated peaks at around 25.9°, within the region of 31.8°–33°, and at 44°, 46.7°, and 49.5°. Consequently, the calculated crystallinities increased to 16 % for PMMA–HA and 22 % for PMMA–HAc, confirming the successful incorporation of HA in the polymer matrix.³⁹ SLA-Ti was analyzed as a positive control to serve as a benchmark, and it produced distinct and sharp diffraction peaks at 2θ angles of 35° and 40°, confirming its highly crystalline nature, with a calculated crystallinity of 90 %. This high level of crystallinity is consistent with the characteristics of standard titanium dental implants.

AFM was applied to measure the surface roughness of the implant samples and obtain nanometer-scale details, where the roughness was quantified using the Ra parameter. Ra was selected rather than 3D parameters in order to ensure direct comparability with the positive control because the manufacturer-specified data for the SLA-Ti implant material was provided only in Ra. However, it is recognized that using Ra alone may not fully capture the complexity of surface topography features. More conventional techniques such as mechanical or contact profilometry are often employed to incorporate additional parameters, including the maximum height difference (Rz), and thus obtain deeper insights into peak–valley distributions. To address this limitation and provide clear visualizations of the surface profiles, characteristic AFM images are presented in Figure 2.⁴⁰

Figure 2.

Three-dimensional topography and two-dimensional views of: (a) PMMA, (b) PMMA–HAc, and (c) PMMA–HA.

The PMMA sample had the lowest Ra value of 108 nm, indicating a smooth surface characteristic of polymeric materials. The synthesized variants PMMA–HAc and PMMA–HA had contrasting surface roughness profiles, where PMMA–HAc exhibited the highest Ra value of 220 nm and PMMA–HA had an intermediate Ra value of 128 nm. The differences in surface roughness between PMMA–HA and PMMA–HAc were attributed to the presence of HA particles producing a more textured surface. The surface roughness of the coated PMMA–HAc samples increased even further because the HA particles remained exposed on the surface rather than being fully embedded within the matrix.^41,42 The surface roughness of these coated biomaterials can be influenced by the coating method and processing parameters, including the applied voltage, drying conditions, and solvent selection. In addition, the chemical composition of the composite and any thermal oxidation treatments can further modify the resulting surface characteristics.⁴³

The SLA-Ti implants had a significantly rougher surface (2.02 μm), which was produced by the sandblasting and acid-etching procedures. These mechanical treatments create deep micro-pits that are known to improve bone–implant contacts and promote osseointegration.⁴⁴ Theoretically, this surface roughness may influence the behavior of bacteria, but friction and wear properties are also critical for determining load transfer and the long-term durability of a prosthesis.^45,46 However, these specific biological responses and tribological behaviors require further confirmation.

Analysis of the pore size and surface morphology showed that the SLA-Ti implants had the smallest and most uniform pores, with an average diameter of 3.97 μm, due to their precision manufacturing and industrial surface treatment processes (See Figure 3). The PMMA samples also contained relatively small and uniform pores, with an average diameter of 18.3 μm.

Figure 3.

SLA titanium (Neobiotech) surface images.

By contrast, the PMMA–HA samples had larger pores (30.86 μm), which could be attributed to disruption of the PMMA matrix caused by the incorporation of HA particles. The incorporation of HA mainly generated micro-voids due to the polarity mismatch between hydrophilic HA and hydrophobic PMMA. This chemical incompatibility promoted the agglomeration of particles and weak interfacial adhesion, resulting in the formation of physical gaps between the filler and polymer matrix.⁴⁷ This increased porosity may have affected the structural integrity, but the difference in the pore size between PMMA–HA and PMMA was not statistically significant (p = 0.974).⁴⁸ PMMA–HA also exhibited microporosity within the range of 10–100 μm, and this pore size interval was previously reported to favor cell attachment and facilitate protein adsorption.⁴⁹ Finally, The PMMA–HAc samples had the largest and most irregular pores, with an average diameter of 141 μm, and this substantial increase was probably due to the gelatin component of the HA slurry coating swelling and trapping air during the drying process, thereby increasing the overall porosity.⁵⁰

The wettability was evaluated for the dental implant materials by using SBF to obtain contact angle measurements. SBF was selected because it closely mimics the ionic composition and pH of human blood plasma, thereby providing a more realistic medium for predicting the in vivo behavior of materials.⁵¹ The results showed that the pure PMMA had a high contact angle of 75.70° and PMMA–HAc had a nearly identical value of 75.53°. Therefore, both surfaces were classified as moderately hydrophobic. The unexpectedly high value for PMMA–HAc was attributed to the presence of micro-voids or larger surface pores. These micro-textures can trap air pockets at the liquid–solid interface, generating a composite contact known as the Cassie–Baxter state.⁵² Consequently, these trapped air pockets could have hindered full water penetration and obscured the material's intrinsic hydrophilicity.

However, PMMA–HA had a substantially lower contact angle of 47.28°, indicating a markedly more hydrophilic surface. By contrast, the SLA-Ti implant had the highest contact angle of 84.36°. The incorporation of HA mainly enhanced the wettability due to the inherently hydrophilic nature of the ceramic phase increasing the surface's affinity for water molecules. In addition to this chemical contribution, HA particles increased the surface roughness to further amplify the hydrophilicity by promoting liquid spreading, which is consistent with the Wenzel principle. The combined effect of hydrophilic functional groups and a rougher surface texture reduced the contact angle and strengthened interfacial adhesion.^53.54

A key strength of this study was the successful synthesis of a PMMA–HA composite with a specific HA loading ratio of 16.2 %, which falls within the recommended range of 10–20 % that was previously suggested to balance bioactivity and material stability. However, this study had several limitations. First, the evaluations were primarily restricted to surface characterization, and the impacts of the observed porosity and roughness on the mechanical strength and structural integrity of the implant were not assessed. Second, the findings provide a theoretical basis for potential osseointegration based only on physicochemical properties, and actual biological performance requires confirmation through future in vivo investigations. Finally, the comparisons in this study were limited to a standard SLA-Ti surface. Modern SLA-Ti implants have diverse surface modifications with enhanced hydrophilicity, so future studies should include a broader range of control surfaces to allow more comprehensive comparisons.

Conclusion

In the present study, the PMMA–HA composite exhibited more favorable in vitro surface properties compared with the PMMA–HAc coated sample. In particular, the PMMA–HA composite exhibited a semi-crystalline structure, improved hydrophilicity, and a microporosity range that may be theoretically suitable for initial cell attachment, whereas the PMMA–HAc sample was limited by irregular macroporosity and hydrophobic behavior. Therefore, future studies should include comprehensive mechanical testing, energy dispersive spectroscopy analysis, and cytotoxicity assays to further validate the performance, long-term durability, and biocompatibility of PMMA–HA dental implant materials.

Ethical approval

All procedures in this study strictly followed the Guide for the Care and Use of Laboratory Animals and the Indonesian Ministry of Health's National Health Research and Development Ethics Standard and Guidelines (2017). Ethical clearance was obtained from the Research Ethics Committee, Faculty of Dental Medicine, Universitas Airlangga, Surabaya, Indonesia (Approval No. 1144/HRECC.FODM/X/2023).

Authors contributions

HRQ & MAA: Conceptualization, Methodology, Software, Data Curation. RDR: Supervision, Visualization, Validation, Funding Acquisition. HSB & ID: Formal Analysis and Investigation. TAW, XJN, & OBC: Writing - Review and Editing.

Declaration of generative AI and AI-assisted technologies in the writing process

During the preparation of this work, the authors used Gemini 3.5 Pro to enhance the language quality and readability. After using this tool, the authors thoroughly reviewed and edited the content as necessary, and take full responsibility for the final version of this publication.

Source of funding

This study was supported by the Institute for Research and Community Service, Universitas Airlangga, through the Airlangga Research Fund 2024 (Grant No. 673/UN3/2024).

Conflict of interest

The authors have no conflict of interest to declare.