Impact of Adding Cerium Zirconium Oxide Nanofibers in 3D-Printed Denture Base Material

Abstract

Purpose: Pure three-dimensional (3D)-printed resin for denture base shows strength in comparison with the conventional heat-cured materials. The purpose of this study was to assess how physical and mechanical properties of 3D-printed denture base resins are affected by the addition of cerium zirconium oxide nanofibers (CeZrO₄ NFs), which have a unique combination of thermophysical and mechanical properties. Materials and Methods: The specimens were digitally created utilizing Microsoft Corporation’s 3D builder software through computer-aided design. To meet the test criteria for transverse strength, impact strength, hardness, radiopacity, and degree of conversion (DC), specimens were designed and printed with specific dimensions according to the relevant standards. The 3D-printed denture base resin was mixed with CeZrO₄ NFs (diameter: 300–800 nm, length: 2–10 µm) at weight percentages of 0.5, 1.0%, 1.5%, 2%, and 2.5%. The data were analyzed using Tukey’s post hoc test (α = 0.05) and ANOVA. Field emission scanning electron microscopy (FESEM) and energy dispersive X-ray spectroscopy (EDX) were used to evaluate surface morphologies of the composites and nanofibers, and the dispersion of the NFs within the resin matrix respectively. Results: The results demonstrated that compared with those of the control group, the average transverse strength, impact strength, and hardness values of the CeZrO₄ NF reinforcement groups significantly increased up to a nanofiller concentration of 1.5 wt.%., whereas those of the other reinforcement groups significantly decreased. For example, the impact strength significantly increased from 5.84 kJ/m² (0 wt.%) to the maximum value 8.76 kJ/m² at 1.0 wt.% CeZrO₄ NF. On the other hand, the Shore D hardness increased from 80.84 for the control group to the maximum value 83.27 at 1.5 wt.% CeZrO₄ NF. The radiopacity increased as the NF concentration increased. Although Fourier transform infrared (FTIR) spectroscopy analysis did not show any noticeable change in the chemical structure of the resin after incorporating the NFs, there was a notable improvement in the DC of the nanocomposites with NF concentrations of 0.5, 1.0 and 1.5 wt.%. Energy dispersive X-ray spectroscopy (EDX) and field emission scanning electron microscopy (FESEM) showed evidence of uniform distribution of the CeZrO₄ NFs in the 3D-printed specimens. Conclusions: The properties of the denture bases fabricated from 3D-printed resin were enhanced by the addition of 0.5%, 1 wt.% and 1.5 wt.% CeZrO₄-milled NFs, though the latter two concentrations produced the most significant results.

1. Introduction

A worldwide dental issue that mostly affects elderly individuals is complete or partial edentulism [1]. The rising trend of the aging population has led to an increasing demand for prosthetic tooth replacements that provide both functionality and esthetics. Since removable dentures are the most popular option clinically for replacing lost teeth, many studies have focused on developing new materials and manufacturing processes for next-generation dentures [2]. The primary goal is to create a material that is cost-effective, easily accessible, biologically compatible, and simple for the dentists to manipulate [3]. Despite its shortcomings, including poor mechanical and physical qualities and a lengthy production procedure, PMMA is still the most commonly used denture base material by the practitioners [4]. Digital technology has proven to have numerous advantages in dentistry in general and particularly in prosthodontics, in terms of manufacturing speed and precision and number of clinical visits. However, certain challenges such as inadequate mechanical qualities of the raw materials used to create dentures using 3D printing still need to be overcome [4,5]. In recent years, dentistry has made extensive use of digital manufacturing technology. Currently, removable dentures can be fabricated using computer-aided design and computer-aided manufacturing (CAD-CAM) technologies, which blend additive and subtractive methods. This technology is also known as rapid prototyping and 3D printing [6,7].

Despite being the current thrust for creating fully removable 3D-printed denture bases, they still have certain limitations with respect to their mechanical and physical characteristics [8,9]. While 3D-printed denture base materials possess flexural strength close to the ISO-accepted value of 65 MPa, they have the lowest flexural strength and surface hardness compared with the traditional and milled denture base materials. As a result, their clinical applications are restricted [5,10,11]. Furthermore, compared with the traditional acrylic resins, 3D printing resins result in inferior double-bond conversion [8].

The use of NPs has been proposed to enhance the characteristics of the polymethylmethacrylate (PMMA) denture base [12]. Following that trend, to overcome the limitations of the 3D-printed denture bases, many studies have explored the effects of adding reinforcing metal oxide nanoparticle (NP) fillers, such as TiO₂, Al₂O₃, and SiO₂, along with adjusting the post-polymerization time, layer thickness, and printing orientation, which appeared to improve some of the mechanical qualities of the 3D-printed resin used in dentures [13,14,15]. Furthermore, more recently, nanoscale molecules were considered as a functional filler for modifying dental restorative materials [16].

According to a study by Aati et al., provisional restorations enhanced with ZrO₂ created with 3D-printed resin performed better over an extended period of time [17]. Mubarak et al. added less than 1 weight percent of silver–titanium dioxide nanofiller to 3D-printed material, which increased the tensile strength, tensile modulus, and flexural strength [18]. A recent study examined the mechanical and surface properties of 3D-printed resins that included silicon dioxide nanoparticles and reported that the hardness, flexural strength, and impact strength of the resins improved [19].

Nano-ZrO₂ has recently drawn much interest as a filler which can enhance the mechanical properties of acrylic resins due to its advantages, including superior toughness, mechanical strength, abrasion, corrosion resistance, and biocompatibility. High mechanical properties of nano-ZrO₂ enable resistance to crack propagation, and its crystalline structure increases the opacity value [15].

Since cerium oxide (CeO₂) nanoparticles have special surface characteristics, high stability, and biocompatibility, they are utilized in many different fields, including biomedicine and technology, and provide surface protection from UV rays and oxidation [20]. In dentistry, CeO₂ nanoparticles are used to reduce microbial development and inflammatory responses [21]. Adding varying amounts of CeO₂ nanoparticles to 3D-printed resins can enhance the mechanical characteristics of the denture bases, including their bending resistance and impact strength [22,23]. CeZrO₄ is a mixed metal oxide composed of cerium oxide (CeO₂) and zirconium oxide (ZrO₂). The CeO₂ component offers an oxygen storage capacity (OSC) and redox behavior (Ce⁴⁺ ↔ Ce³⁺), whereas zirconia contributes to the thermal stability and mechanical strength of the material.

To the best of the authors’ knowledge, no prior research has investigated the impact of CeO₂ or CeZrO₄ nanofibers (NFs) on the mechanical characteristics of 3D-printed denture base materials. Therefore, in this study, hardness, flexural strength, impact strength, and radiopacity of the 3D-printed denture base resin were assessed after the addition of CeZrO₄ NFs. High surface area-to-volume ratio of the nanoparticles increases nanofiber-matrix interfacial bonding (interlocking), which improves the mechanical properties, prevents crack propagation, and reduces polymerization shrinkage [24]. The null hypothesis predicted that the addition of CeZrO₄ NFs would not affect the properties of the 3D-printed denture base resin.

2. Materials and Methods

2.1. Materials and Specimen Grouping

The 3D-printed specimens were fabricated using DentaBase (Asiga, Sydney, Australia, under ISO 13485:2016 and EN ISO 13485:2016), a denture base resin that is known for its great strength and natural pink shade. The manufacturer’s data sheet states that the resin has 7,7,9(or 7,9,9)-trimethyl-4,13-dioxo-3,14-dioxa-5,12-diazahexadecane-1,16-diyl bismethacrylate, tetrahydrofurfuryl methacrylate, and diphenyl(2,4,6-trimethylbenzoyl) phosphine oxide. CeZrO₄ NFs with a diameter ranging between 300 and 800 nm and length varying between 2 and 10 µm (Sigma–Aldrich, Darmstadt, Germany) were added to the resin at various concentrations (0.5, 1.0, 1.5, 2.0, and 2.5, weight percent) to create nanocomposites. Three hundred specimens were prepared in conduct with various tests (60 specimens each for flexural strength, impact strength, hardness, radiopacity and degree of conversion with n = 10).

2.2. Specimen Preparation

2.2.1. Specimen Design

The flexural strength hardness test specimens had dimensions of 65 × 10 × 2.5 ± 0.2 mm in accordance with ISO 20795-1:2013 guidelines [25], whereas impact test specimens were fabricated with dimensions of 80 × 10 × 4 ± 0.2 mm in accordance with ISO 179 (2003) [26]. The specimen dimensions for the radiopacity test were 10 × 20 × 3 mm [27], and dusty powder was used for evaluating chemical bonding characteristics and polymerization. To create the specimens, CAD software was used, and a stereolithography (STL) file—a popular file format for 3D printing—was exported.

2.2.2. Mixing of CeZrO₄ NFs with Resin

Initially, the NF suspension was made by mixing CeZrO₄ NFs with 3 mL of 99.9% ethyl alcohol in an ultrasonic mixer (MSE Soniprep 150, Amsterdam, The Netherlands) for 3 min. The number of NFs added to the groups was 0.3 for 0.5 wt.%, 0.7 for 1.0 wt.%, 1.05 gm for 1.5 wt.%, 1.4 for 2 wt.% and 1.75 for 2.5 wt.% weight. A special shaker lc-3d mixer (NextDent, Rock Hill, SC, USA) was used to create a homogeneous mixture of basic resin for denture printing. Then, the NF suspension was mixed in a dark amber glass container with 70 g of denture base resin. An Alfa HS-860 magnetic stirrer (A-LEA, China) was used to mix the denture base resin and CeZrO₄ NFs suspension for 30 min at 60 °C, and the glass container was partially covered to allow the evaporation of alcohol. The mixture was then kept in the dark amber glass container and mixed for 8 h at room temperature (25 ± 5 °C).

2.2.3. 3D Printing of Specimens and Post-Processing

The specimens were printed via a digital light processing (DLP) 3D printer (Asiga MAX UV, Sydney, Australia) with the following settings: layer thickness = 0.05 mm, separation velocity = 4.3 mm/s, separation pressure limit = 300 g/cm², heater temperature = 30 °C, and high-power ultraviolet solid-state 385 nm Light Emitting Diode (LED). To prevent the resin with the CeZrO₄ NF mixture from being exposed to outside light, the printer vat lid was sealed after the mixture was poured. On the build platform, the liquid resin started to solidify as a result of polymerization when the printer’s light was projected onto it. This process was repeated until the 3D models of the specimens were completed. All 30 specimens were printed within 40 min, and the layer development process was horizontal.

Using a sharp knife, the specimens were carefully removed from the 3D printer platform once the printing process was completed. Subsequently, the specimens were subjected to 3 min of ultrasonic cleaning (Dewin, Zhuhai, China); the layers were exposed to UV light at a wavelength of 405 nm to continue the polymerization process, which continued until the 3D-printed model was produced. The specimens were cured in a post-curing device (Otoflash G171, China) for 20 min in accordance with the manufacturer’s guidelines. The procedure was conducted with two cycles of 2000 flashes on each side of each specimen under an atmosphere of inert nitrogen gas.

The specimens were subsequently cleaned with isopropyl alcohol (99.9%) to remove any uncured resin and polished using an acrylic bur (Syndent Tools Co., Jiangsu, China) and a lathe polishing tool (Renfert, Hilzingen, Germany). The dimensions of the specimens were measured using a digital Vernier scale (Kirti NDT/Maharashtra, India) with a precision of ±0.01 mm. Before testing, all specimens were submerged in distilled water for 48 h in accordance with ADA specification No. 12, 1999. Different steps of fabrication and post-processing are shown in Figure 1.

2.3. Characterization Procedures

2.3.1. Flexural Strength Test

A universal Instron testing apparatus (Jianqiao Testing Equipment, Dongguan, China) was used to perform three-point bending test. Each specimen was positioned using a fixture, which had two parallel supports spaced 50 mm apart. A plunger positioned in the middle of the supports applied a steady force at a cross-head speed of 1.0 mm/min until it broke. The test was carried out in compliance with ADA specifications No. 12, 1999, and with a 500 N load cell. The following formula was used for calculating the flexural strength [28].

$$\mathrm{Flexural} \mathrm{strength}={\displaystyle \frac{3PI}{2{bd}^2}}\left({\displaystyle \frac{\mathrm{N}}{{\mathrm{m}\mathrm{m}}^2}}\right)$$

(1)

where I is the span length, P is the peak load, b is the specimen width (mm), and d is the specimen thickness (mm).

2.3.2. Impact Strength Test

In accordance with the ISO 179 guidelines, the amount of impact energy absorbed by the specimen during fracture was evaluated. An impact testing equipment (Testing Machines Inc. (TMI), New Castle, DE, USA) was used to perform the Charpy unnotched impact tests. The specimen was struck by a free-swinging pendulum with a two-joule capacity while being supported horizontally at both ends. The impact strength was calculated using the following formula [29].

$$\mathrm{Impact} \mathrm{strength}={\displaystyle \frac{E}{B·D}}\times {10}^3\left({\displaystyle \frac{\mathrm{K}\mathrm{J}}{{\mathrm{m}}^2}}\right)$$

(2)

where E is the impact energy absorbed, D is the specimen thickness (mm), and B is the specimen width (mm).

2.3.3. Surface Hardness Test

The hardness of the specimens was evaluated via a Shore D device (Time Group Inc., Capannori, Italy). It was composed of an indenter that was blunt-pointed and had a diameter of 0.8 mm, tapering to 1.6 mm. The indenter was fitted with a digital scale that ranged from 0 to 100 units. The specimens were subjected to strong and quick indenter acceleration and hardness value was directly read from the digital scale. Shore D hardness was calculated using the mean of the three measurements on each specimen. The test was carried out in compliance with ADA standard, No. 12 (1999).

2.3.4. Radiopacity Test

To assess radiopacity, the specimens were spread out across a wax plate of 3 mm thick. Wax was added to replicate soft tissue absorbing and dispersing media. To standardize the film density, a 10 mm aluminum step wedge with a sequential 1 mm step height was affixed next to the specimens. The specimens, wax plates, and aluminum step wedges were positioned on the exposure side of a 35 × 43 cm cassette. An X-ray machine (Siemens, Berlin, Germany) was used, with a distance of one meter between the specimens and the X-ray source. The equipment was operated at 53 kV and 5 mA and its exposure period was 0.35 s, much like that of standard chest radiography [27]. A light transmission densitometer (Pehamed Densonorm 21i, France) was used to measure the variation in image density (optical density, OD) of each specimen in comparison with standard acrylic resin and aluminum step wedges. Each specimen was subjected to five measurements at various locations, and a mean value was calculated. The procedure for measuring OD of the X-ray films by the densitometer is shown in Figure 2. The radiopacity of the tested samples was quantified in terms of aluminum equivalent thickness (mmAl) by applying a calibration curve derived from the aluminum step wedge thickness ranging from 1 to 10 mm according to the guidelines specified in ISO 4049. This calibration served as a reference for interpreting the radiopacity of the tested materials. By applying quadratic regression, the approximate conversion equation is: mmAl = 10.352 × (OD)² + 26.870 × (OD) + 18.268, as shown in Figure 2D. The optical density decreased with increasing Al step wedge thickness, from 1.15 at a 1 mm Al thickness to 0.34 at a 10 mm Al thickness. This demonstrated the calibration and validity of the radiographic measurement system where thicker aluminum absorbed more X-rays resulting in lower optical density values and provided a reference scale for comparing the radiopacity of the tested materials with known aluminum thickness.

2.3.5. Chemical Characterization and Degree of Conversion Calculations

Fourier transform infrared (FTIR) spectroscopy (SHIMADZU, Kyoto, Japan) was employed to obtain FTIR spectra at wavelengths ranging from 4000 to 400 cm⁻¹. Ten specimens from each group and CeZrO₄ NF powder were tested. Furthermore, degree of conversion (DC) during polymerization was determined using the FTIR spectra according to Equation (3) [30]. Following polymerization, the nanocomposite specimens were compared to the liquid resin, which was scanned as the baseline [31,32].

$$\begin{array}{c}\mathrm{DC}(\%)=\\ (1-{\displaystyle \frac{\left(\frac{{\mathrm{T}}_{1637}}{{\mathrm{T}}_{1608}}\right)\mathrm{p}\mathrm{e}\mathrm{a}\mathrm{k}\mathrm{h}\mathrm{i}\mathrm{g}\mathrm{h}\mathrm{t}\mathrm{s}\mathrm{a}\mathrm{f}\mathrm{t}\mathrm{e}\mathrm{r}\mathrm{p}\mathrm{o}\mathrm{l}\mathrm{y}\mathrm{m}\mathrm{e}\mathrm{r}\mathrm{i}\mathrm{z}\mathrm{a}\mathrm{t}\mathrm{i}\mathrm{o}\mathrm{n}}{\left(\frac{{\mathrm{T}}_{1637}}{{\mathrm{T}}_{1608}}\right)\mathrm{p}\mathrm{e}\mathrm{a}\mathrm{k}\mathrm{h}\mathrm{i}\mathrm{g}\mathrm{h}\mathrm{t}\mathrm{s}\mathrm{b}\mathrm{e}\mathrm{f}\mathrm{o}\mathrm{r}\mathrm{p}\mathrm{o}\mathrm{l}\mathrm{y}\mathrm{m}\mathrm{e}\mathrm{r}\mathrm{i}\mathrm{z}\mathrm{a}\mathrm{t}\mathrm{i}\mathrm{o}\mathrm{n}}})\\ \times 100\end{array}$$

(3)

2.4. Characterization of Surface Morphology and Composition

CeZrO₄ NFs and broken surface were investigated using a Field Emission Scanning Electron Microscope (FESEM—INSPECT F50, FEI business, Hillsboro, OR, USA) at various magnifications. Since the resin specimens are not conductive, charges might result in optical distortions, blurring, or even destruction. To enhance image quality and disperse any charge build-up on the non-conductive resin surface, the specimens were sputter-coated with gold to a thickness of approximately 1 nm. Energy dispersive X-ray (EDX) spectroscopy (MIRA3TESCAN, Denver, CO, USA) was used to analyze chemical composition of the specimen.

2.5. Statistical Analysis

Data analysis was conducted using Prism 9 (GraphPad Software, San Diego, CA, USA) and SPSS (Statistical Package for Social Science, version 21). One-way analysis of variance (ANOVA) was used in order to compare each group’s mean value and the post hoc Tukey’s HSD test to identify any significant difference between the individual groups. Data homogeneity was verified using the Levene test, and the normality of the data distribution was assessed via the Shapiro–Wilk test. Significant differences were indicated by p values less than 0.05.

3. Results

3.1. Transverse Strength

Figure 3 shows that the flexural strength significantly increased as the NF concentration increased. The values ranged from 102 MPa (95% CI: 100.96–103.03) in the control group to 108.88 MPa (95% CI: 108.08–109.69) at 0.5 wt.%; CeZrO₄ NFs peaked at the 1 wt.% (116.85 MPa, 95% CI: 116.35–117.35) and then started to decrease at 1.5 wt.% (107.66 MPa, 95% CI: 106.92–108.40) until it reached the lowest value (62 MPa, 95% CI: 59.92–64.07) at the 2.5 wt.% NF concentration.

3.2. Impact Strength

Figure 4 shows that impact strength (kJ/m²) significantly increased as the NF concentration increased from 5.84 kJ/m² (95% CI: 5.50–6.19) at 0 wt.% to 6.67 kJ/m² (95% CI: 6.32–7.01) at 0.5 wt.% with the maximum value at 1 wt.% (8.76 kJ/m², 95% CI: 8.16–9.35), and then began to decrease at 1.5 wt.% (7.06 kJ/m², 95% CI: 6.72–7.41) until reaching the lowest mean (5.1 kJ/m², 95% CI: 5.02–5.17) at 2.5 wt.%.

3.3. Surface Hardness

Figure 5 shows that hardness increased significantly as the NF concentration increased, with a peak value (83.27, 95% CI: 83.16–83.38) at 1.5 wt.%, after which it began to decrease gradually at 2.0 wt.% until reaching (79.86, 95% CI: 79.74–79.97) at 2.5 wt.%.

3.4. Radiopacity

As shown in Figure 6A, the optical density was the highest (lowest radiopacity) in the control group (1.22, 95% CI: 1.22–1.23) and decreased gradually as the NF concentration increased until it reached (1.03, 95% CI: 1.02–1.03) at 2.5% (highest radiopacity). In addition, Figure 6B showed a gradual increase in radiopacity as the filler content increased, rising from 0.89 mmAl in the control specimen to 1.57 mmAl in the material containing 2.5 wt.% CeZrO₄ NFs. This trend demonstrated that higher filler loading enhanced the material’s ability to attenuate X-rays. A filler concentration of 1.5 wt.% NPs is the minimum concentration that reached clinically acceptable radiopacity (≥1.0 mmAl) according to ISO 4049. However, other lower levels of concentrations such as 0.5 wt.% and 1.0 wt.% were also very close to the acceptable limit.

3.5. Microstructure and Composition

The FESEM images (Figure 7) demonstrated that CeZrO₄ NFs were uniformly distributed within the 3D-printed denture base resin. Furthermore, the good bonding between the NFs and the resin matrix was also observed.

Figure 8 shows an overview of the CeZrO₄ NFs, indicating a large amount of fiber material aggregation. The specimen consists of randomly oriented NFs mixed with smaller particles. The fibers and particles have rough surfaces, indicating that a high surface area was beneficial for application. Analysis of the NFs revealed that the effective particle size was 36.2 nm. The NFs appeared to have nonuniform diameters and lengths, with some being needle-like or rod-shaped.

Figure 9 displays EDX diagrams for 3D-printed denture base resin that does not contain CeZrO₄ NFs. EDX examination of the elemental makeup of the 3D-printed resin revealed carbon (C) and oxygen (O) as shown in Figure 10.

The elemental composition of the 3D-printed resin containing the CeZrO₄ NFs was as follows: cerium (Ce), zirconium (Zr), oxygen (O), and carbon (C). The EDX analysis results are shown in Figure 11 whereas EDX mapping of the nanocomposite is presented in Figure 12.

3.6. FTIR and Degree of Conversion

The FTIR spectrum for polymethylmethacrylate (PMMA) liquid in Figure 13 shows the presence of CH2 rocking mode peaks at 777.31, 813.96, and 943.19 cm⁻¹. The asymmetric stretching bond peak appears at 1105.21 cm⁻¹, which can be attributed to the C-O-C group. The methyl carbonyl group C-C-O stretching vibrations observed at about 1246.02 cm⁻¹. The presence peak at 1454.33 cm⁻¹ could be due to the CH2 deformation, while the peak at 1533.41 cm⁻¹ was attributed to C-H stretching bond. The high peak observed at 1728.22 cm⁻¹ could be due to the stretching vibration of the carbonyl group C=O. Moreover, a transmittance peak appears at about 2956.87 cm⁻¹ resulting from the C-H bonds. Also, the peak at 3379.29 cm⁻¹ was attributed to (O-H) stretching bond.

The FTIR spectrum for polymethylmethacrylate (PMMA) liquid after polymerization shows similar peaks as before polymerization. The presence of CH2 rocking mode peaks at 777.31 cm⁻¹ and 813.96, cm⁻¹. The asymmetric stretching bond peak appears at 1118.71 cm⁻¹, which can be attributed to the C-O-C group. The methyl carbonyl group C-C-O stretching vibrations were observed at about 1244.09 cm⁻¹. The peak at 1463.97 cm⁻¹ is due to the CH2 deformation, while the peak at 1533.41 cm⁻¹ was attributed to C-H stretching bond. The peak observed at 1734.01 cm⁻¹ was due to the stretching vibration of the carbonyl group C = O but this time it was smaller than before polymerization. Also, the peak appears at approximately 2958.80 cm⁻¹, which could be attributed to the C-H bonds. The peak at 3410.15 cm⁻¹ was attributed to (O-H) stretching bond.

The FTIR spectrum of polymerized polymethylmethacrylate (PMMA) after the addition of CeZrO₄ NFs at varying quantities of 0.5, 1.0, 1.5, 2.0 and 2.5 wt.% shows similar peaks as those before the addition of the NFs. The bands and their wavenumbers, which represent the FTIR spectrum comparison for all the specimens, clearly revealed that no new bonds appeared after the polymerization of the PMMA after the addition of the CeZrO₄, i.e., there was no chemical reaction in the process, and it was only a physical blend. The degree of conversion (DC) values of the nanocomposites at lower concentrations (0.5 to 1.5 wt.%) showed statistically significant increase compared to the control group (Figure 14). However, at higher NF concentrations the DC values started to decrease.

4. Discussion

The incorporation of CeZrO₄ NFs into 3D-printed denture base resin produced significant changes in flexural strength, impact strength, surface hardness, radiopacity and degree of conversion (DC) while 1.0 wt.% NFs addition produced the most optimal mechanical and physical properties. Increasing above this NFs concentration leads to a decrease in these properties, except a continuous decrease in radiopacity with increasing filler concentration. Therefore, the null hypothesis was rejected.

The DLP printer was utilized in this research since it is a faster technique that creates items with a higher resolution [33]. By producing patterned laser light, DLP technology offers an advantage over stereolithography where all layers can be cured in a single laser exposure rather than requiring the laser to scan each area one at a time. Owing to this advantage, the building time is unaffected by the shape of the associated layer or the quantity of items [6].

The mechanical and physical properties of 3D-printed denture base could be affected by the orientation, curing technique, building parameter, CAD-CAM software, layer shrinkage, layer thickness, and layer numbers [34]. Air trapped in the resin liquid during specimen printing might have led to the formation of voids between printed layers, which would have impacted the mechanical and physical characteristics of the printed specimen [24]. A layer thickness of 50 µm and 0° orientation were used while printing the specimens for this research. By improving interlayer diffusion, air voids can be reduced as the layer thickness decreases [35]. According to Borella et al., specimens printed at a 50 µm layer thickness had superior qualities to those printed at 100 µm [36]. Owing to thinner individual printing layers, the flexural strength improved, and this increase could be most likely attributable to the fact that smaller layers make the polymerization process more effective. They allow light to flow through them more easily than the thicker layers do [37]. The angle of printing orientation influences the mechanical properties of 3D-printed materials. Alqutaibi et al. [38] found that the 0° orientation resulted in better flexural than the 45° and 90° orientations did while printing denture base specimens. Another study by Shim et al. [39] found that 0° orientation showed the highest flexural strength in 3D-printed resins compared to 45° and 90° printing orientations. This could be due to changing the layer building printing direction from parallel to perpendicular to the load direction, which can produce strong adhesion within the same layers as opposed to 45° and 90° orientations [40]. However, Jafarpour et al. [41] found that flexural strength was significantly greater (p < 0.001) at an orientation of 90° compared with 0° and 45° during 3D printing of denture base resin.

The flexural strength improved progressively with increasing CeZrO₄ NFs concentration, peaking at intermediate levels (0.5–1.5 wt.%), before decreasing at 2 wt.%. This could be due to the fact nanoparticles work as reinforcing agents, enhancing the resistance of the material to stress-induced fractures by filling the spaces between the polymer chains and improving the strength, stiffness, and overall structural integrity of the printed component. Previous research suggested that increasing the flexural strength of a nanocomposite might also depend on achieving a uniform dispersion of nanoparticles inside the resin matrix [42], confirming the findings of Majeed and Hamad, who reported similar improvements after the addition of 1.0 wt.% yttria-stabilized zirconia (YSZ) nanofillers to a 3D-printed denture base. However, the flexural strength of the 3D-printed resin decreased as YSZ nanoparticle concentrations increased beyond 1.0 wt.% [43]. Similarly, Kadhum and Hamad reported a significant increase in the flexural strength of the 3D-printed denture base after the addition of 0.5 wt.% barium titanate nanoparticles whereas an increase in BaTiO₃ concentration exceeding 1.0 wt.% resulted in a reduction in the flexural strength of the 3D-printed resin [44]. Additionally, Mhaibes et al. reported that the flexural strength of 3D-printed denture base increased significantly after the addition of 1.0 wt.% titanium oxide nanotubes and decreased at 2.0 wt.% nanotube concentrations [24]. The statistically significant decrease in flexural strength at higher nanofiber concentrations could be explained by a change in the internal structure of the polymerized resin where the nanofiller works as an impurity or internal defects caused by particle clustering. In this study, higher concentrations of CeZrO₄ (2.0 and 2.5 wt.%) in the resin matrix showed a negative impact on the DC and flexural strength possibly owing to an increase in the amount of unreacted monomer, which might act as a plasticizer. Therefore, in order to maximize the benefits of nano additives in terms of enhancing clinical performance of the denture base, it is crucial to determine their ideal concentration.

Lowery et al. [37] reported that a 50 µm layer thickness resulted in greater strength than a 100 µm layer thickness did and documented a general tendency toward better strength with a lower layer thickness. The smaller layer thickness used in this study could be the cause of the observed improvements in flexural strength. Furthermore, the post-curing unit plays a critical role in enhancing flexural strength. The inert post-curing process in a nitrogen gas atmosphere counteracts the detrimental effects of the oxygen inhibition layer [37]. The use of 3D-printed denture in clinical settings could be possible as both the control and modified groups at lower NF concentration exceeded the ISO-20795-1:2013 flexural strength criteria, which call for a minimum value of 65 MPa.

To avoid fractures in the case of a sudden drop, the denture base material must have sufficient impact strength. In this study, the incorporation of CeZrO₄ NFs into 3D printing resin caused a significant increase in the impact strength reaching peak with the 1.0 wt.% and 1.5 wt.% NF concentrations compared with that of the control group, and these results were statistically significant for both experimental groups. This improvement could be attributed to the consequence of the development of cross-links or supramolecular bonds within the nanocomposites, which could stop fracture propagation. Additionally, a strong bond between the resin matrix and the nanofiller could increase interfacial shear strength and arrest crack progression [45]. Also, this increase in impact strength could be related to effective dispersion of NFs at lower concentration, combined with the thin layer thickness and controlled post-curing in nitrogen environment, which enhanced cross-link density by reducing oxygen inhibition. This finding was consistent with the results of Fatalla et al., who reported that improved impact strength could be attributed to strong interfacial adhesion between the nanoparticles and the polymer matrix. The nanoparticles occupy the interchain regions, restrict chain displacement, and promote effective energy dissipation, which suppresses crack growth and fracture propagation [46]. Gad et al. reported that the addition of nano-ZrO₂ (with or without glass fibers) significantly improved the impact resistance [47]. These results agree with those of previous studies by Mhaibes et al. [24] Chen et al. [48] and Alshaikh et al. [13] which indicated that the impact strength was increased by the addition of titanium dioxide nanotubes (TiO₂), polyetheretherketone (PEEK), titanium dioxide (TiO₂), and zirconium dioxide nanoparticles (ZrO₂) to 3D-printed resins. A reduction in impact strength at higher concentration might be associated with fiber clustering and reduced DC.

The enhanced hardness suggests better wear resistance and maintenance of the occlusal form, which are key for clinical durability. The surface hardness increased with increasing NFs loading owing to the inherently high hardness of the zirconia and cerium oxide. Furthermore, the stacking and overlapping of polymer molecules, which limit their mobility, are also responsible for the hardness increase. This enhances the ability of a material to withstand scratches, cuts, and plastic deformation. The types of forces holding atoms together determine a material’s hardness. Here, the coherence of the mixture was enhanced by fiber–fiber mesh formation in the resin matrix and uniformly dispersed stiff and rigid NFs, which increases the hardness of the NF-reinforced resin. Similar reinforcement effects were found in the study by Pavlíková et al. [49] on hybrid nanocomposites, where surface hardness improved with modified zirconia nanoparticle content. Kim et al. [50] reported that the hardness values of the 2.5wt.% Ce group were significantly greater than those of the control group (p  <  0.05), and the study of Aati et al. [17] demonstrated that adding ZrO₂ NPs significantly increased the hardness of temporary restorations. Similarly, Gad et al. [18] reported that the hardness of 3D-printed resin increased with the addition of SiO₂ NPs in comparison to the pure resin. It should be noted that with the increase in nanofiber content the numerical values of hardness changed only by a small margin, although statistical analysis confirmed that the differences were statistically significant (p < 0.05). Therefore, this small change might not translate to a real-life clinical significance.

Broken denture fragments could be swallowed by the wearers in some cases. In order to make it easier to locate the denture fragments, radiopacity in denture base is desirable. Since most denture base materials are radiolucent, researchers are now focusing heavily on developing radio-opaque denture base materials. Since extraoral radiographs are more frequently employed as diagnostic tools in emergency situations, they were used in this investigation [51]. Numerous investigations have been carried out on the radiopacity of denture base resins; an increase in radiopacity is correlated with a decrease in radiographic optical density (OD). Aluminum is used to equalize the density of exposed films in radiopacity testing because tooth structures such as enamel have a similar coefficient of linear absorption. In accordance with ISO 4049 standard, aluminum allows for radiopacity up to 2 mm thick that reaches clinically acceptable radiopacity (>1 mmAl) [52]. With an increase in radiographic densities, radiopacity increases. When the results of the reinforced groups were compared with those of the unaltered group, there was a statistically significant increase in radiopacity as the concentration of CeZrO₄ NFs increased.

As a result, the transmission densitometer results indicated that the radiographic density decreases as greater amounts of CeZrO₄ NFs are added, whereas the control group has the highest mean radiographic density (low radio-opacity). The increase in radio-opacity was statistically significant, and the relative radio-opacity increased as the CeZrO₄ NFs concentration increased. This is evidently because the polymer matrix contains radio-opaque CeZrO₄ NFs, which absorb more radiation than the polymer matrix and are more radio-opaque. The high atomic number of Ce (Ce, atomic number: 58) and the high atomic number of Zr (Zr, atomic number: 40) in comparison with the low atomic number of the chemical constituent of acrylic (C, atomic number: 6, O, atomic number: 8) brought radio-opacity characteristics by the presence of CeZrO₄ NFs. An element’s ability to absorb X-rays is mostly determined by the cube of its atomic number [27].

FTIR spectra confirmed that the addition of CeZrO₄ NFs did not introduce new chemical bonds when the 3D-printed resin and CeZrO₄ NFs were combined, and all the bonds were the same in all the spectra except for those of the CeZrO₄ NFs. This made it very evident that the mixing was only physical in nature and not chemical. This also suggests that the long-term stability of denture depends primarily on dispersion uniformity and mechanical entrapment of the NFs within the resin rather than their chemical bonding.

Mechanical characteristics like hardness are significantly impacted by the degree of monomer conversion, and this characteristic would be improved in proportion to the rate of converted monomer into a polymer network [17]. The degree of polymerization of composite resins is closely related to the curing process, which may be impacted by the energy of the radiation, the range of wavelengths, the duration of exposure, translucency, and light scattering by the filled system [53]. In this study, the addition of a small amount of CeZrO₄ NFs (0.5 wt.%, 1.0 wt.% and 1.5wt.%) increased the DC (35%, 36% and 36%, respectively) compared with that of the control group (30%). This could be attributed to the photocatalytic properties of the CeZrO₄ NFs [54,55]. NFs helped the curing light energy diffuse throughout the 3D-printed material and accelerated photopolymerization and DC, which would decrease the amount of unreacted monomer. This could explain why the nanocomposites had better mechanical properties and surface hardness than the pure resin [56]. A greater DC reflects increased cross-linking and denser polymeric matrix with fewer unreacted monomers and gaps [57]. This is consistent with the results of Altarazi et al., who reported that the DC considerably increased when 0.1% TiO₂ was added to 3D-printed liquid resin [58]. A greater DC reduces the quantity of remaining monomer, which lowers the negative reactions in the oral tissues [59]. When the amount of NFs increased (2% and 2.5%wt NFs), the DC decreased (30% and 21%, respectively) and negatively affected the mechanical properties due to a reduction in light transmission and potential filler agglomeration which hinder polymerization [57]. Additionally, increased filler loading elevates the resin viscosity and restricts the mobility of unreacted monomers, further impacting polymerization kinetics [17].

Although these findings demonstrate promising improvement, several limitations must be considered. The present investigation focused on a single variety of CeZrO₄ NFs, a particular printing orientation, a single post-printing polymerization and a specific 3D-printed denture resin. Therefore, the present outcomes might be different from those of alternative resins, printing techniques, and printing settings. Future research should build on these findings by examining the effects of CeZrO₄ NFs on physical and mechanical properties of a variety of denture base resins in order to have a more complete understanding. The ideal NF percentages found in this study could not be directly compared due to a lack of prior studies on the utilization of CeZrO₄ NFs in 3D-printed dental resins. Nevertheless, the ideal proportion of CeZrO₄ NFs found in this investigation was comparable to that found in other studies that employed nanoparticles in 3D-printed resins. Furthermore, the current study was carried out in dry settings, which do not accurately reflect how they would function in an actual oral environment. Therefore, accelerated aging under different liquid media like water, artificial saliva, or coffee will be conducted to assess the performance of the nanocomposites.

The current results demonstrated that 3D-printed denture base materials represent a promising alternative to conventional heat-cured polymers in clinical settings. The integration of nanotechnology may further mitigate the limitations associated with current materials and improve their structural integrity and functional characteristics, leading to enhanced clinical durability and superior patient outcomes.

5. Conclusions

The foundation material for 3D-printed dentures was successfully reinforced with cerium zirconium oxide-milled nanofibers (CeZrO₄ NFs) with different concentrations (0.5, 1.0, 1.5, 2.0, and 2.5 wt.%). The results showed that 1.0 wt.% and 1.5 wt.% CeZrO₄ NFs nanocomposites had the highest flexural strength, impact strength, hardness and degree of conversion (DC) values. However, a general trend of decrease in mechanical properties at the highest concentration could be due to nanofiber clustering and the resultant reduction in DC. This new nanocomposite could help in improving the performance of dentures in clinical set-up.