Study on Synergistic Enhancement of Surface Properties of Ti-6Al-4V Alloy for Dental Applications by Magnetic Abrasive Finishing

Abstract

Titanium alloys are widely used in dental implants due to their superior biocompatibility and mechanical strength. However, these alloys are prone to corrosion and wear in the oral environment, thereby shortening their clinical lifespan. This study investigates the enhancement of titanium alloy surface properties using magnetic abrasive finishing (MAF) and examines the influence of magnetic needle diameters (0.2–1.5 mm) on surface modification. Titanium alloy samples were processed by MAF and systematically evaluated for surface morphology, grain size, surface hardness, residual stress, electrochemical corrosion behavior, and tribological performance. Results demonstrated that MAF improves surface morphology, significantly refines grain size, and enhances surface hardness and compressive residual stress, thereby optimizing surface properties. The 1.0 mm magnetic needle group demonstrated the best performance, achieving a Vickers hardness of 376.71 ± 12.48 HV and a compressive residual stress of −579.1 ± 8.49 MPa. In addition, this group showed a higher self-corrosion potential (−0.5661 V), a lower corrosion current density (0.0114 μA·cm⁻²), and the lowest wear rate ((4.49 ± 0.42) × 10⁻⁴ mm³/N·m) in artificial saliva, demonstrating superior corrosion and wear resistance. Overall, MAF technology markedly enhances the surface integrity of titanium alloys in artificial saliva through the synergistic effects of grain refinement and stress modulation. These findings provide valuable experimental evidence supporting future efforts to optimize the surface properties of titanium alloy dental implants.

1. Introduction

With the ongoing advancements in modern dentistry and dental implantology, there is a growing demand for improved surface properties of implant materials. Titanium alloys are among the most successful biomaterials, owing to their high strength-to-weight ratio and excellent biocompatibility. They not only promote tissue integration effectively but also exhibit unique physical properties, making them widely used in dental applications [1,2,3]. However, within the complex and dynamic oral environment, titanium alloys are still significantly affected by combined corrosion and wear, severely limiting their long-term clinical reliability [4]. Electrolytes, microbial metabolites, and pH fluctuations in saliva can induce electrochemical corrosion, resulting in the continuous release of metal ions. This process may trigger local inflammation, cytotoxicity, and reduced biocompatibility [5,6,7,8]. Repeated masticatory stress and occlusal forces cause mechanical wear, leading to implant loosening, reduced stability, and eventual failure [9,10]. The synergistic interaction between corrosion and wear has emerged as a critical factor limiting the long-term clinical performance of titanium implants. Consequently, developing efficient and controllable surface modification strategies to improve corrosion and wear resistance has become a central focus in dental implant research.

Traditional surface modification methods primarily include laser cladding, plasma spraying, and anodizing. Among these, laser cladding employs high-energy lasers to deposit materials onto the substrate surface. It provides high coating density and strong interfacial bonding, but its limitations include costly equipment and complex operation [11,12]. Plasma spraying offers high processing efficiency; however, maintaining consistent coating quality is challenging, and adhesion is influenced by multiple factors, leading to reliability issues [13]. Anodizing is simple to perform and produces an oxide film on metal surfaces; however, the film exhibits low mechanical strength, compromising its long-term stability [14]. Although these methods have achieved progress, they still encounter challenges such as high energy consumption, complicated procedures, high costs, and limited adaptability to complex structures in practical applications, which restrict their clinical application. Therefore, it is urgent to develop a low-energy, environmentally friendly surface modification technology suitable for complex curved structures.

Magnetic abrasive finishing (MAF) is an advanced surface modification technique that employs a magnetic field to control magnetic abrasives, enabling precise material removal from workpiece surfaces. It provides advantages including uniform material removal, adaptability to complex geometries, and the absence of mechanical contact wear [15]. Moreover, the process is environmentally friendly, significantly mitigating ecological impacts by lowering the consumption and emission of pollutants during machining [16]. MAF enhances material performance in complex service environments by inducing residual compressive stress and increasing surface hardness [17,18,19]. In recent years, advances in MAF technology and its strong controllability of abrasive media structures have revealed distinct advantages in processing complex cavities and miniature workpieces, highlighting its substantial industrial potential and clinical relevance.

Although the potential of magnetic abrasive finishing (MAF) in precision machining is increasingly recognized, studies on its application to titanium alloys—particularly in dental surface modification—remain limited. Specifically, systematic experimental evaluations of the influence of magnetic needle diameter—a critical process parameter—on grinding performance are lacking. Therefore, this study investigates how magnetic needle diameter regulates the surface properties of Ti-6Al-4V alloy using magnetic needle grinding technology, with the goal of providing experimental evidence and optimization guidelines for its application in dental titanium alloy surface modification. The methodology adopted for this study is described in detail below, outlining the experimental setup and evaluation criteria.

2. Materials and Methods

2.1. Sample Preparation

Cylindrical specimens of Ti-6Al-4V alloy (Zapp Precision Metals, Schwerte, North Rhine-Westphalia, Germany) were fabricated using CNC lathe processing, with a diameter of 6 mm and a length of 2.5 mm. The Ti-6Al-4V alloy primarily consists of approximately 90% titanium, 6% aluminum, and 4% vanadium. The specimens were sequentially ultrasonically cleaned in acetone, ethanol, and distilled water for 10 min each, then dried and prepared for subsequent experiments. A schematic of the magnetic abrasive finishing process is presented in Figure 1. The experiment was conducted using the single-factor variable method, in which the magnetic needle diameter—identified as the primary factor influencing surface quality [20]—was varied while all other parameters were maintained at predetermined constant values to investigate its effect on surface finishing performance. The magnetic needles were made of SUS304 stainless steel, and the grinding fluid used was HD-233A (Holding Electric Co., Ltd., Taichung City, Taiwan Province, China). Magnetic abrasive finishing was performed using an HD-728 machine (Holding Electric Co., Ltd., Taichung City, Taiwan Province, China). During processing, the magnetic pole rotation direction was reversed halfway to improve finishing uniformity. After grinding, the specimens were ultrasonically cleaned again with acetone (≥99.5%), anhydrous ethanol, and distilled water, and then dried. The experiment was divided into four groups, each treated with MAF using different magnetic needle diameters. The detailed parameters are listed in

Table 1. Experimental design.

Group	Magnetic Needle Diameter/mm	Grinding Fluid Ratio	Magnetic Needle Quality/kg	Magnetic Needle Length/mm	Magnetic Pole Rotation Speed/Hz	Grinding Time/min
MAF0.2	0.2	1/50	1.0	5.0	40	60
MAF0.5	0.5	1/50	1.0	5.0	40	60
MAF1.0	1.0	1/50	1.0	5.0	40	60
MAF1.5	1.5	1/50	1.0	5.0	40	60

. Unprocessed samples served as the control group.

2.2. Surface Morphology

The surface morphology of the samples was examined using a FlexSEM 1000 II scanning electron microscope (Hitachi Industrial Products, Ltd., Marunouchi, Tokyo, Japan) operated at an acceleration voltage of 10 kV.

2.3. Phase Analysis

Phase analysis was conducted using a D/Max 2500PC X-ray diffractometer (Rigaku Holdings Co., Ltd., Akishima City, Tokyo, Japan) equipped with a Cu target and Kα radiation (λ = 0.154 nm). Each group consisted of one independent sample. The operating conditions were: tube voltage 40 kV, tube current 40 mA, and a scanning range of 2θ = 30–90°. X-ray diffraction (XRD) patterns were processed with Jade 6.5 software (Materials Data, Inc., Livermore, CA, USA) to determine phase composition, full width at half maximum (FWHM). Based on the obtained FWHM values, the grain size was calculated using the Scherrer formula (Equation (1)).

$$D={\displaystyle \frac{K\mathrm{\lambda }}{\mathrm{\beta }\cos \mathrm{\theta }}}$$

(1)

Here, D denotes the average crystallite size, K is the shape factor, λ represents the wavelength of the X-ray radiation, β corresponds to the FWHM of the diffraction peak, and θ is the Bragg diffraction angle.

2.4. Surface Micro Vickers Hardness

The Vickers hardness of the samples was measured using a 200HV-5 Vickers hardness tester (Laizhou Huayin Test Instrument Co., Ltd., Laizhou City, Shandong Province, China). During testing, each sample was mounted on the stage and adjusted to the focal plane of the microscope. A 136° rhombic diamond indenter applied a load of 200 g, which was maintained for 10 s. After unloading, the lengths of the two diagonals of the indentation were measured under the microscope, and the hardness was calculated using the standard Vickers formula with the applied load. Each group consisted of five independent samples (n = 5). The Vickers hardness was calculated according to Equation (2):

$$HV=0.1891{\displaystyle \frac{F}{d^2}}$$

(2)

Here, HV denotes Vickers hardness, d refers to the average length of the two indentation diagonals, and F indicates the applied test force.

2.5. Surface Residual Stress

The surface residual stress of the titanium alloy was measured using the Laboratory X-ray Residual Stress Diffractometer MG2000 and analysis system (Proto Manufacturing Ltd., Kitchener, ON, Canada). Prior to testing, the Cu Kα X-ray source was set to a tube voltage of 25.00 kV and a tube current of 30.00 mA, and the sample surface was thoroughly cleaned. The specimen was then securely fixed to the measurement platform to ensure that its normal was approximately parallel to the incident X-ray beam. Residual stress was determined using the sin²ψ method by sequentially collecting diffraction peaks of the crystal planes at seven inclination angles (ψ = 0.00°, 10.83°, 18.84°, 19.00°, 19.16°, 27.17°, and 38.00°) [21]. Each group consisted of three independent samples (n = 3). Finally, the surface residual stress was calculated according to Equation (3):

$${\mathrm{\sigma }}_{\mathrm{\varphi }}=-{\displaystyle \frac{E}{2\left(1+\mathrm{\nu }\right)}}\cot {\mathrm{\theta }}_0\cdot {\displaystyle \frac{\mathrm{\pi }}{180}}\cdot {\displaystyle \frac{\partial \left(2{\mathrm{\theta }}_{\mathrm{\varphi }\mathrm{\psi }}\right)}{\partial \left({\sin }^2\mathrm{\psi }\right)}}$$

(3)

Among these, σ_ϕ denotes the residual stress on the material surface; E is the Young’s modulus of the crystal plane; v is the Poisson’s ratio of the crystal plane; θ_ϕψ denotes the X-ray diffraction angle of the stressed material; and ψ is the angle between the surface normal and the normal of the diffraction crystal plane. A negative σ_ϕ indicates compressive stress, whereas a positive value indicates tensile stress.

2.6. Electrochemical Corrosion

Electrochemical corrosion tests were conducted using a CHI660E electrochemical workstation (Shanghai Chenhua Instrument Co., Ltd., Shanghai, China). The testing system comprised a saturated calomel electrode (SCE) as the reference electrode, a platinum wire as the counter electrode, and the sample as the working electrode. All electrodes were immersed in artificial saliva (AS, pH 6.8) at 37 °C to simulate the oral environment. The composition of the artificial saliva is shown in Table 2. Before both polarization and impedance measurements, the samples were stabilized under open-circuit potential (OCP) conditions for 60 min to ensure that the system reached a relatively stable state. After OCP stabilization, electrochemical impedance spectroscopy (EIS) was performed over a frequency range of 10⁻² to 10⁵ Hz. Subsequently, dynamic potential polarization tests were carried out by scanning from –1.6 V to +1.6 V (vs. Ag/AgCl), and the polarization curve was recorded. The corrosion potential (E_corr) and corrosion current density (I_corr) were obtained from the polarization curve using the Tafel extrapolation method. Each group consisted of one independent sample. The polarization resistance (R_p) was calculated using the Stern–Geary Equation (4):

$$R_p={\displaystyle \frac{{\mathrm{\beta }}_a\left|{\mathrm{\beta }}_c\right|}{2.303I_{\mathrm{corr}}\left({\mathrm{\beta }}_a+\left|{\mathrm{\beta }}_c\right|\right)}}$$

(4)

In this context, βa denotes the anodic Tafel slope, while βc denotes the cathodic Tafel slope.

Table 2. Artificial saliva components.

NaCl	KCl	NaH2PO4, 2H2O	CaCl2, 2H2O	Na2S, 2H2O	Urea	Distilled-Water
0.4 g	0.4 g	0.78 g	0.795 g	0.005 g	1 g	1000 mL

Table 2. Artificial saliva components.

NaCl	KCl	NaH2PO4, 2H2O	CaCl2, 2H2O	Na2S, 2H2O	Urea	Distilled-Water
0.4 g	0.4 g	0.78 g	0.795 g	0.005 g	1 g	1000 mL

2.7. Friction and Wear Performance

The friction and wear performance was evaluated using a CSM-TRB3 tester (Anton Paar, Graz, Austria). Prior to testing, the friction pair surfaces were polished and subsequently cleaned with acetone in an ultrasonic bath. The samples were then mounted in the drive unit and fixture of the testing machine and preloaded to ensure precise system alignment. To simulate the oral environment, tests were performed under artificial saliva (AS) conditions at pH 6.8 and 37 °C, using 3-mm-diameter Si₃N₄ balls as counter bodies to evaluate the wear resistance of the substrate. Test parameters included a constant load of 2 N, a reciprocating frequency of 2 Hz, and a stroke amplitude of 3 mm. Wear volume was determined by white-light scanning and analyzed with the accompanying software, Image Plus (Anton Paar, Graz, Austria). Each group consisted of three independent samples (n = 3). The wear rate was calculated according to Equation (5):

$$W={\displaystyle \frac{V}{L\times F}}$$

(5)

Here, W denotes the wear rate, V represents the wear volume, L is the sliding distance, and F is the applied normal load.

2.8. Statistical Analysis

All data related to micro-Vickers hardness, residual stress, and friction/wear performance are expressed as mean ± standard deviation (mean ± SD), with at least three independent specimens per group (n ≥ 3). Prior to statistical analysis, data normality was assessed using the Shapiro–Wilk test, and homogeneity of variances was verified using the Levene test. When both assumptions were satisfied, one-way analysis of variance (ANOVA) followed by Tukey’s multiple comparison test was performed to evaluate intergroup differences. If either assumption was violated, the Kruskal–Wallis non-parametric test was applied. Statistical significance was defined as p < 0.05. For phase analysis and electrochemical corrosion tests, since only one specimen was examined per group, results are presented descriptively without statistical evaluation.

3. Result

3.1. Surface Morphology

Figure 2 shows the surface morphology of samples processed under different MAF parameters. The control group exhibited continuous deep grooves and small pits, with grooves oriented in a consistent direction and overall high roughness. In the MAF_0.2 group, macroscopic grooves were partially smoothed, while parallel micro-striations and local irregular micro-elevations remained, indicating a slight improvement in surface topography. In the MAF_0.5 group, most grooves were smoothed, forming shallow and broad wavy traces; the residual grooves became shallower and narrower, further improving surface flatness. The MAF_1.0 group showed an even flatter surface with only a few shallow scratches. At MAF_1.5, machining marks were almost entirely eliminated, although localized particle deposits were present.

3.2. Phase Analysis

According to the XRD patterns in Figure 3, the titanium alloy samples primarily consisted of α-titanium, with a minor fraction of β-phase detected. No additional diffraction peaks were observed, indicating that MAF treatment did not induce significant phase transformation or formation of new phases. Compared with the control group, the MAF-treated samples exhibited slightly low-angle shifts in the main diffraction peaks.

Table 3. Characterization parameters of the main diffraction peak of the α-Ti (101) plane.

Group	FWHW (°)	Grain Size (nm)
Control	0.338	18.5
MAF0.2	0.362	17.7
MAF0.5	0.385	17.1
MAF1.0	0.400	15.4
MAF1.5	0.390	16.8

presents the full width at half maximum (FWHM) of the α-Ti (101) diffraction peak under different magnetic needle conditions, along with the corresponding grain sizes. The results indicate that with increasing magnetic needle diameter, the FWHM of the diffraction peak exhibits an overall upward trend. Specifically, the FWHM in the control group, the FWHM was 0.338°, increasing to 0.362°, 0.385°, and 0.400° after MAF_0.2, MAF_0.5, and MAF_1.0 treatments, respectively, before slightly decreasing to 0.390° in the MAF_1.5 group. Consistent with the FWHM changes, the grain size first decreased and then slightly increased. The grain size in the control group was 18.5 nm, which decreased to 17.7 nm, 17.1 nm, and 15.4 nm in the MAF_0.2, MAF_0.5, and MAF_1.0 groups, respectively, before increasing to 16.8 nm under MAF_1.5 conditions.

3.3. Surface Micro-Vickers Hardness

Figure 4 presents the surface micro-Vickers hardness of titanium alloy samples under various treatment conditions. The control group exhibited a Vickers hardness of 324.24 ± 12.88 HV. While MAF-treated samples showed a substantial increase in hardness. Specifically, the MAF_0.2 group reached 328.14 ± 13.52 HV, the MAF_0.5 group reached 331.83 ± 15.44 HV, and the MAF_1.0 group achieved the highest value of 376.71 ± 12.48 HV. The MAF_1.5 group exhibited a hardness of 369.97 ± 11.38 HV, slightly lower than that of the MAF_1.0 group. Statistical analysis revealed that the Vickers hardness values of the MAF_1.0 and MAF_1.5 groups were significantly higher than that of the control group (p < 0.05). Overall, the Vickers hardness increased with larger MAF magnetic needle diameters, peaking in the MAF_1.0 group and slightly decreasing thereafter.

3.4. Surface Residual Stress

Figure 5 presents the surface residual stress test results of titanium alloy samples from the control group and various MAF treatment groups. The control group exhibited an average residual stress of approximately −460.03 ± 17.57 MPa. With the progression of MAF treatment, the residual compressive stress gradually increased. The residual compressive stresses for the MAF_0.2, MAF_0.5, and MAF_1.0 groups were −504.57 ± 2.89 MPa, −539.6 ±7.79 MPa, and −579.1 ± 8.49 MPa, respectively, with MAF_1.0 exhibiting the highest value. Statistical analysis revealed that the residual compressive stresses of the MAF1.0 and MAF1.5 groups were significantly higher than that of the control group (p < 0.05), while no significant difference was observed between these two groups (p > 0.05). When the magnetic needle diameter was further increased to 1.5 mm, the residual stress of the MAF_1.5 group slightly decreased to −555.63 ± 12.79 MPa. Figure 6 presents the linear fitting results between 2θ and sin²ψ obtained from X-ray diffraction tests of samples across different groups and ψ angles. All samples exhibited strong linear correlations, with coefficients of determination (R²) exceeding 0.93, confirming the high stability and accuracy of the residual stress test. The linear fitting parameters corresponding to the table section in Figure 6 are provided in the Supplementary Materials (Table S1). According to the X-ray diffraction sin²ψ method, the slope of the fitted curve is directly proportional to the residual compressive stress. A larger slope corresponds to a higher residual compressive stress. The fitting results demonstrate that the slope variation trend aligns with the residual stress variation shown in Figure 5. The MAF_1.0 group exhibited the largest slope, corresponding to the maximum residual compressive stress. The control group showed the smallest slope, corresponding to the lowest residual compressive stress. Data points in each group were evenly distributed, and fitting errors were minimal, further confirming the reliability of the residual stress variation trend.

3.5. Electrochemical Corrosion

Figure 7 presents the dynamic polarization curves of titanium alloy samples subjected to different MAF treatments in an artificial saliva environment. All groups exhibit similar polarization profiles, indicating comparable redox reactions during electrochemical corrosion. Distinct passivation plateaus with broad passivation zones were observed in all treated samples, confirming excellent passivation performance regardless of magnetic needle size. Electrochemical parameters derived from curve fitting are summarized in

Table 4. Electrochemical corrosion parameters.

Group	Ecorr (V vs. SCE)	Icorr (μA·cm−2)	βa (mv/dec)	|βc| (mv/dec)	Rp ( $\mathrm{\Omega }$·cm2)
Control	−1.0647	0.2869	362	127	1.43 × 105
MAF0.2	−0.9553	0.1349	161	86	1.80 × 105
MAF0.5	−0.8329	0.1164	345	211	4.88 × 105
MAF1.0	−0.5661	0.0114	489	454	8.97 × 106
MAF1.5	−0.6798	0.1069	294	204	4.89 × 105

. A more positive self-corrosion potential (E_corr) indicates higher corrosion resistance, whereas a lower corrosion current density (I_corr) indicates a slower corrosion rate. As the magnetic needle diameter increased from 0.2 mm to 1.0 mm, E_corr shifted positively and I_corr decreased markedly, demonstrating progressive enhancement of corrosion resistance. The MAF_1.0 group exhibited the best performance, with E_corr reaching −0.5661 V and I_corr reduced to 0.0114 μA·cm⁻², while the untreated control group showed the poorest resistance (E_corr = −1.0647 V, I_corr = 0.2869 μA·cm⁻²). Figure 8 shows the Nyquist impedance plots of each sample group. All curves display the characteristic semicircular shape typical of charge-transfer-controlled capacitance. As the magnetic needle diameter increased, the curve radius first expanded and then slightly decreased, indicating that polarization resistance (R_p) followed an overall upward trend. The maximum Rp of 8.97 × 10⁶ Ω·cm² was observed in the MAF_1.0 group, significantly higher than the 1.43 × 10⁵ Ω·cm² measured in the control group, confirming the superior corrosion resistance achieved with this treatment.

3.6. Friction Wear

Figure 9 depicts the temporal evolution of friction coefficients. The control group maintained a coefficient of ~0.40 up to 400 s, followed by a sharp increase above 0.55, indicating deteriorating tribological performance. In contrast, MAF-treated samples exhibited lower and more stable friction coefficients. Specifically, MAF_1.0 maintained the lowest values (0.35–0.40), while MAF_0.2 and MAF_0.5 showed slight increases after 400 s. The MAF_1.5 group also rose at later stages but remained below the control group. As shown in Figure 10, the average coefficient in the control group was 0.422 ± 0.016, whereas all MAF-treated groups showed reductions. MAF_1.0 achieved the lowest value (0.387 ± 0.023), followed by MAF_0.5 (0.392 ± 0.012), MAF_0.2 (0.403 ± 0.008), and MAF_1.5 (0.416 ± 0.022). Figure 11 shows the wear profiles. The control group exhibited the deepest and sharpest grooves, while all MAF treatments reduced both depth and width. MAF_1.0 produced the shallowest, U-shaped profile, indicating superior surface protection. Wear volume and rate results (Figure 12) further confirmed these trends. The control group had the highest wear ((9.25 ± 1.24) × 10⁻³ mm³, (6.43 ± 0.86) × 10⁻⁴ mm³/N·m). All MAF-treated groups performed better, with MAF_1.0 showing the lowest wear ((6.47 ± 0.60) × 10⁻³ mm³, (4.49 ± 0.42) × 10⁻⁴ mm³/N·m), while MAF_1.5 ((6.57 ± 0.56) × 10⁻³ mm³, (4.56 ± 0.39) × 10⁻⁴ mm³/N·m) ranked second, followed by MAF_0.5 ((8.36 ± 0.73) × 10⁻³ mm³, (5.83 ± 0.52) × 10⁻⁴ mm³/N·m) and MAF_0.2 ((8.74 ± 1.11) × 10⁻³ mm³, (6.05 ± 0.78) × 10⁻⁴ mm³/N·m). Statistical analysis indicated that the reductions in wear volume and wear rate for MAF_1.0 and MAF_1.5 were significant compared with the control group (p < 0.05).

4. Discussion

This study investigates the enhancement of titanium alloy surface properties for dental applications using the MAF process with varying magnetic needle diameters. Experimental results show that MAF markedly improves corrosion and wear resistance, with the 1.0 mm magnetic needle group exhibiting the best performance. These improvements result from the synergistic effects of grain refinement and residual stress regulation induced by the MAF process.

Grain size is a critical factor influencing the mechanical and physical properties of materials. For nanomaterials with grain sizes below 100 nm, XRD is a widely used and effective method to determine grain size. Grain sizes before and after MAF were calculated from XRD data using the Scherrer formula. Results indicate that MAF markedly refines grain size, confirming its positive effect on the material’s microstructure. Grain refinement improves structural uniformity and enhances corrosion resistance. Wang et al. reported that grain refinement enhances corrosion resistance by increasing grain boundary density, accelerating passivation film nucleation and growth, and improving film semiconductor properties [22]. Electrochemical tests in this study further confirmed this conclusion. MAF-treated samples showed a positive shift in corrosion potential and a decrease in corrosion current density, indicating improved corrosion resistance. Electrochemical impedance spectroscopy revealed a larger Nyquist semicircle for MAF-treated samples, indicating a denser oxide film and higher interfacial impedance. This hinders corrosive ion migration, thereby enhancing material stability in corrosive environments [23,24,25].

According to the Hall–Petch relationship [26], surface hardness is inversely proportional to the square root of grain size. Thus, grain refinement directly increases surface hardness. Micro-Vickers hardness tests showed that surface hardness after MAF treatment increased significantly, with the MAF_1.0 group exhibiting the greatest improvement, consistent with grain refinement trends. This observation was also reported by Fujian et al. After MAF treatment, a refined microstructural layer formed on the titanium alloy surface, accompanied by a marked reduction in grain size. This impeded dislocation movement, thereby enhancing surface hardness [27]. Material wear is closely related to surface hardness and is often described by the Holm–Archard equation. According to this relationship [28], wear volume is inversely proportional to surface hardness. Thus, MAF treatment reduces wear volume by increasing surface hardness. Experimental results confirmed this: under identical conditions, the MAF_1.0 group showed the lowest wear volume, consistent with Archard equation predictions. This demonstrates that increased surface hardness mitigates wear. Optimizing MAF parameters therefore refines grain size and enhances surface hardness, markedly improving wear resistance.

Besides grain refinement, residual compressive stress induced during MAF is another key factor enhancing overall performance. Previous studies show that compressive stresses induced in the surface layer suppress crack initiation and propagation, thereby extending fatigue life [29,30,31]. In this study, the MAF_1.0 group exhibited the highest residual compressive stress, likely due to stronger grinding impact. Other studies indicate that MAF generates a stable microstress field on surfaces, enhancing fatigue performance and structural stability under complex conditions [32]. This stress state helps withstand alternating oral loads and mitigates interface stress concentration under corrosion–fatigue coupling. It delays crack propagation, lowers corrosion rates, and improves implant stability and biocompatibility [33,34,35,36]. Higher residual compressive stress also enhances surface hardness and lowers the friction coefficient, further improving wear resistance [37,38].

MAF treatment markedly enhances corrosion and wear resistance of titanium alloys, a critical improvement for dental implant applications. The oral cavity, a complex and dynamic environment, readily induces frictional corrosion on implant surfaces during long-term use. This may cause adverse effects, including metal ion release, interface damage, and implant micromotion. In severe cases, it compromises implant stability and biocompatibility [39,40,41]. The surface modifications in this study enhance long-term stability and biocompatibility, reducing clinical complication risks. Specifically, improved corrosion resistance reduces metal ion release, lowers cytotoxicity and immune activation, and mitigates local inflammation by stabilizing passivation films. This ultimately improves biocompatibility and long-term implant safety [42]. Furthermore, improving wear resistance is crucial. It strengthens implant mechanical stability, reduces friction-induced damage in body fluids, and lowers risks of early loosening or failure [43]. This synergistic enhancement is expected to improve implant long-term stability and service life, while optimizing clinical outcomes and patient satisfaction.

This study demonstrates that MAF markedly enhances the surface integrity, microhardness, residual compressive stress, and corrosion resistance of the Ti-6Al-4V alloy. Nevertheless, several limitations should be acknowledged. The sample size was limited, and certain analyses (e.g., grain size and electrochemical measurements) were performed on a single specimen, lacking statistical validation. All experiments were conducted under static in vitro conditions that do not fully replicate the complex, dynamic oral environment. Surface roughness was mainly characterized qualitatively using SEM without quantitative assessment. This limitation partially restricts a detailed understanding of surface morphology evolution but does not compromise the overall conclusions. Unlike conventional roughening techniques, MAF produces smooth surfaces appropriate for abutment and connection interfaces, thereby reducing bacterial adhesion and wear and enhancing long-term stability. The process is relatively environmentally friendly; however, proper abrasive removal and recycling are necessary to avoid titanium residue contamination. Ongoing studies are focusing on the biological performance of MAF-treated Ti-6Al-4V, including its compatibility with gingival fibroblasts and epithelial cells and the adhesion behavior of Streptococcus mutans and Porphyromonas gingivalis, to further optimize its biological response and clinical applicability.

5. Conclusions

This study systematically investigated the influence of magnetic needle MAF processing on the surface properties of titanium alloys for dental applications, focusing on magnetic needle diameters ranging from 0.2 to 1.5 mm. Comprehensive characterization—including Surface morphology, grain refinement, surface hardness, residual stress, electrochemical corrosion resistance, and tribological behavior—led to the following findings.

MAF treatment significantly refined the grain structure and improved the surface morphology, with the smallest grain size and highest surface hardness (376.71 HV, >15% improvement) observed in the 1.0 mm group. Substantial compressive residual stress (−579.1 MPa) was also generated, which retarded fatigue crack propagation and improved structural stability. Electrochemical performance was notably enhanced in simulated oral environments: the 1.0 mm group exhibited a positive shift in self-corrosion potential (−0.5661 V), a reduced corrosion current density (0.0114 μA·cm⁻²), and an increased polarization resistance (8.97 × 106 Ω·cm²), indicating improved passivation stability. Furthermore, tribological properties were optimized, with the 1.0 mm group achieving a low friction coefficient (0.387) and wear rate (4.49 × 10⁻⁴ mm³/N·m), confirming superior anti-wear performance and surface integrity retention.

In summary, magnetic needle MAF processing enables comprehensive optimization of titanium alloy surfaces by tuning magnetic needle parameters. In particular, a 1.0 mm magnetic needle diameter synergistically enhances both electrochemical stability and mechanical durability, suggesting its potential as an effective surface modification approach for titanium alloys used in dental-related environments.