Antistatic Melt-Electrowritten Biodegradable Mesh Implants for Enhanced Pelvic Organ Prolapse Repair

Abstract

Pelvic organ prolapse (POP) is a health condition that can significantly impact patients’ quality of life. Unfortunately, most available treatments present drawbacks such as high recurrence rates, risk of complications, poor tissue integration, and the need for reintervention. One promising alternative is the use of biodegradable implantable meshes, which can support the organs, guide tissue regeneration, and be fully absorbed without damaging the surrounding tissues. In this study, biodegradable polycaprolactone (PCL) meshes were fabricated using melt electrowritten (MEW), incorporating the antistatic agent Hostastat^® FA 38 (HT) to address these limitations. The goal was to produce microscaffolds with suitable biophysical properties, particularly more stable fiber deposition and reduced fiber diameter. Different HT concentrations (0.03, 0.06, and 0.1 wt%) were investigated to assess their influence on the fiber diameter and mechanical properties of the PCL meshes. Increasing HT concentration significantly reduced fiber diameter by 14–17%, 39–45%, and 65–66%, depending on mesh geometry (square or sinusoidal). At 0.06 wt%, PCL/HT meshes showed a 24.10% increase in tensile strength and a 55.59% increase in Young’s Modulus compared to pure PCL meshes of similar diameter. All formulations demonstrated cell viability >90%. Differential scanning calorimetry (DSC) revealed preserved thermal stability and changes in crystallinity with HT addition. These findings indicate that the antistatic agent yields promising results, enabling the production of thinner, more stable fibers with higher tensile strength and Young’s Modulus than PCL meshes, without adding cellular toxicity. Developing a thinner and more stable mesh that mimics vaginal tissue mechanics could offer an innovative solution for POP repair.

1. Introduction

Pelvic organ prolapse (POP) is a condition in which pelvic organs descend due to weakened support structures such as muscles, ligaments, and fascia [1,2,3]. It affects nearly half of women over 50 years who have had vaginal childbirth, and its prevalence is expected to increase with population aging [4,5,6,7]. While conservative treatments exist, many advanced cases require surgical intervention, with up to 15% of women undergoing POP surgery during their lifetime [8].

Synthetic meshes, initially designed for hernia repair [9], were introduced for POP treatment in the early 2000s [10]. However, their poor mechanical compatibility with vaginal tissue, along with complications such as inflammation, erosion, and excessive stiffness, led to safety concerns [8,11]. In 2019, the FDA banned transvaginal mesh products for POP repair in the U.S., highlighting the need for alternative, tissue-compatible solutions [12].

These limitations have driven the development of new, more compatible solutions for POP repair. Additive manufacturing, particularly melt electrowritten (MEW), has emerged as a promising strategy to fabricate customized biodegradable meshes with high resolution and precise control over fiber architecture. MEW enables the production of scaffolds with tunable geometry, porosity, and fiber alignment, mimicking the microstructure of native tissue [13,14].

Polycaprolactone (PCL) is widely used in MEW [15,16] due to its biocompatibility, mechanical stability, and slow degradation rate [17,18,19,20]. Unlike solvent-based techniques, MEW allows solvent-free processing by melting and extruding PCL directly, which is advantageous for biomedical applications [21]. Although medical-grade PCL is more expensive than technical-grade versions, it ensures greater purity, reproducibility, and printing precision [19], which are critical factors for clinical translation. Prior studies have demonstrated its use in FDA-approved devices and its ability to enhance tissue regeneration without increasing local stiffness [22].

Antistatic agents are amphiphilic molecules commonly used to reduce surface charge accumulation, which can destabilize the electrohydrodynamic jet during MEW. Their inclusion enhances surface conductivity by attracting moisture and increasing ion concentration at the polymer–air interface [23], resulting in improved jet stability and finer fiber formation. Hostastat^® FA 38 (HT), a commercial antistatic additive [24], has previously demonstrated potential to reduce fiber diameter and improve ionic conductivity in melt electrospun systems [25]. However, such additives may also affect biocompatibility, making cytotoxicity testing essential when targeting biomedical applications [26]. Mixing HT with PCL via melt blending enables homogeneous distribution, potentially influencing both mechanical performance and cell compatibility [27].

Ensuring implant sterility is a critical step in biomedical device development [28]. Sterilization and disinfection techniques such as ultraviolet (UV) light and ethanol (EtOH) immersion are commonly used for polymeric materials due to their accessibility and low cost. However, these methods can potentially alter polymer structure, especially when used in combination [29]. Therefore, assessing the cytocompatibility of sterilization and disinfection protocols with PCL/HT meshes is essential to ensure safety and performance.

Additionally, incorporating additives such as HT may influence the thermal behavior of the polymer. Differential scanning calorimetry (DSC) is a valuable tool for characterizing melting behavior and crystallinity, helping determine whether the additive alters the material’s thermal stability—an important factor for processing and degradation in vivo [30].

Beyond material selection, the structural design of the mesh, particularly fiber geometry, orientation, and pore architecture, plays a critical role in its mechanical behavior and biological integration. Studies have shown that scaffolds mimicking the anisotropic mechanical properties of vaginal tissue can reduce stress shielding and improve functional outcomes [31]. Meshes with tailored fiber patterns, such as square or sinusoidal geometries, offer different tensile responses [9], potentially influencing implant performance in vivo. Therefore, understanding how geometry interacts with material properties is essential for designing effective biodegradable meshes for POP repair.

The main goal of this study was to evaluate the influence of different concentrations of an antistatic agent, HT, on both the fiber diameter and mechanical properties of biodegradable meshes made of PCL. This was achieved by mixing PCL with the antistatic agent HT to mitigate electrostatic forces during the printing process. Additionally, the influence of the predefined fiber geometry on microscaffold properties was evaluated. The cytotoxicity of the meshes, based on the wt% of HT used, was also studied, along with the optimal sterilization method for these meshes.

2. Methodology

Figure 1 illustrates the methodology used in this study. First, the MEW prototype was calibrated for both technical- and medical-grade PCL to ensure that the resulting meshes had the desired fiber diameter. The polymer was then melt-blended with the antistatic additive at different weight percentages. The resulting pellets were used to fabricate the meshes via the MEW technique. Subsequently, cytotoxicity assays were conducted to assess the biological compatibility of the meshes, and uniaxial tensile tests were performed to evaluate their mechanical properties. Finally, a DSC analysis was carried out to investigate the thermal behavior of the material.

2.1. Polymer

PCL presents a variety of features that are appealing in biomedical applications, such as processability, mechanical properties, and high biocompatibility [32]. It has a low melting temperature (58–60 °C) [18,33], fast solidification, great malleability, 3D printing capacity, heat molding, and shape memory [33].

When considering a tissue engineering application, the degradation rate is an essential material characteristic. In the case of POP repair, the material’s degradation rate is expected to match the tissue’s regeneration rate. If degradation is slower than tissue regeneration, it will delay tissue growth. Conversely, if it is faster, a loss of connection between the tissue and the scaffold will occur, leading to a delay in the healing process [34]. Typically, PCL degrades after 2 to 3 years [17], but its degradation rate can be manipulated by changing its molecular weight and crystallinity or even by modifying its structure [20]. However, this depends on the properties of the PCL scaffolds, such as porosity.

In this study, both technical- and medical-grade PCL were used. The technical-grade PCL was commercially available as filament and obtained from 3D4Makers (Haarlem, The Netherlands) under the trade name Facilan™ PCL 100. It has a diameter of 1.75 mm, a density of 1.1 g/cm³ (ISO 1183 [35]), and a melting point of 58–60 °C. The medical-grade PCL was obtained from Corbion (Amsterdam, The Netherlands) in pellet form, under the trade name PURASORB^® PC 12. This polymer must be stored at –15 °C to preserve its properties for up to five years. After opening, it retains its characteristics for up to one year if kept at room temperature in its original packaging.

2.2. Antistatic Agent

The antistatic agent used in this study was HT, a commercially available additive from Clariant (Muttenz, Switzerland). This antistatic agent was designed to be properly mixed with thermoplastics, avoiding any chemical incompatibility between the melted polymer and this additive [25].

According to the technical sheet [24], HT is an ethoxylated alkylamine derived from renewable fatty acids, with a dripping point of 97 °C. The seller states that this additive is a highly effective internal antistatic agent for polymers, primarily used to equalize or reduce static charges on the polymer surface.

2.3. MEW Device

The MEW prototype used during this study was built with a modular architecture based on the different functions of each component, including structure, control, movement, heating, material supply, material collection, and a high-voltage generator [10].

The device was built using an XY moving collector (whose movements are guaranteed by linear actuators driven by the stepper motors [10]) and a Z-moving printing head (maximum height of 70 mm) [15]. The collector is a square-shaped aluminum plate with a 3 mm thickness and an area of 270 × 270 mm² [10]. The positive voltage is applied to the collector and the negative to the nozzle, leading to a uniform electrical field that allows MEW printing [15]. This device is equipped with both fiber and pellet extruders, meaning that a wide range of materials can be used.

2.4. MEW Device Calibration

The printing process involved two main steps: calibration of the MEW device, followed by the fabrication of the desired meshes. For calibration purposes, 90 × 90 µm square meshes were printed with fiber diameters of 160, 200, and 260 µm. Fiber diameter was measured using a ZEISS AxioPhot microscope (ZEISS, Oberkochen, Germany) in combination with OLYMPUS Stream Basic software.

Calibration began with technical-grade PCL due to its lower cost compared to medical-grade PCL. Only square meshes were printed during this step to minimize material waste. As technical-grade PCL is supplied in filament form, it was cut into small segments—referred to as pellets—for compatibility with the pellet extruder. The distance between the nozzle and collector (Z-height) was fixed at 0.3 mm to eliminate its influence as a variable parameter.

Table 1. Technical-grade PCL calibration printing parameters.

Temperature (°C)	160–180–200–220
Collector’s Speed (mm/min)	500–1000–1200–1400–1800
Voltage (kV)	2.63–3.03–3.23–3.43–3.63–4.03–4.63

summarizes the printing parameters tested during the calibration process.

To initiate the calibration process, square meshes with a fiber diameter of 260 µm were first printed to minimize material waste. Subsequently, meshes with fiber diameters of 160 and 200 µm were fabricated. A single-parameter variation approach was adopted to ensure precise calibration, in which only one printing parameter—such as temperature, collector speed, or applied voltage—was adjusted at a time.

The initial printing parameters were selected based on a previous study [9], which reported that an ideal voltage for printing PCL meshes was 6 kV. However, in the present setup, stable printing was achieved with voltage values in the range of 3–4 kV.

2.5. Meshes Design

To obtain the desired fiber’s diameter, it is necessary to adjust the extrusion value, E (mm), in the G-code files, for which the FullControl software (v3) was used.

To perform this calculation, Equation (1) was used, where $D_{extrusion}$ is the desired fiber diameter (µm), $L_{extrusion}$ is the length of the printed fiber (mm), and $D_{marketfilament}$ is the diameter of the PCL filament, which is 1.75 mm.

$$E={\displaystyle \frac{{(D_{extrusion}\times 0.001\times 1.5)}^2\times L_{extrusion}}{D_{marketfilament}^2}}$$

(1)

The influence of mesh geometry on tensile behavior was investigated by comparing two designs, a square pattern and a sinusoidal pattern, as illustrated in Figure 2.

For the square-shaped meshes, a pore size of 1.5 mm was selected. The sinusoidal mesh design was based on Equation (2), where a represents the amplitude of the wave, b its wavelength, and c and d are constants that shift the wave in the y- and x-directions, respectively.

$$y=a·\mathit{sin}\left({\displaystyle \frac{2\pi }{b}}·\left(x-c\right)\right)+d$$

(2)

Previous studies [16] on different sinusoidal meshes showed that wavy fibers with an amplitude of 1.25 mm and wavelength of 3.5 mm behaved like vaginal tissue when submitted to tension. Since the mechanical tests are uniaxial, a horizontal wave with an amplitude of 3 mm and a wavelength of 35 mm was designed to support the vertical waves.

2.6. Uniaxial Tensile Testing

Tensile tests allow one to visualize how the mesh behaves when subjected to a high load in one direction. In this regard, two main evaluations were performed: how the meshes behave under load as a function of the HT concentration, and how the HT concentration affects the mechanical properties of the meshes.

To ensure consistency across all tensile tests, a “model test” was created in VectorPro software (v 2.1), with a minimum extension of 0 mm, a maximum extension of 200 mm, and an elongation speed of 10 mm/min. After the test is properly performed, the software VectorPro exports the values of load applied and displacement that the folds suffered.

Figure 3 presents examples of a square and sinusoidal mesh sample being tested in Mecmesin MultiTest 2.5-dV equipment (Mecmesin, Slinfold, UK) with a load cell of 100 N. All samples exhibited consistent deformation behavior during uniaxial tensile testing. No tearing or detachment occurred near the grips, and the deformation patterns were highly reproducible across specimens.

2.7. Differential Scanning Calorimetry

DSC analyzed the thermal behavior of the polymer samples to evaluate their thermal stability and identify phase transitions. Three consecutive temperature scans were recorded to ensure the reproducibility of the results, using the STA 449 F5 Jupiter from Netzsch, Selb, Germany. During each scan, the samples were heated from 20 °C to 100 °C at a controlled heating rate of 10 °C min⁻¹ in standard mode.

2.8. Melt Blending

To evaluate the effect of the additive on fiber diameter while ensuring biocompatibility, three concentrations of HT (Clariant Plastics & Coatings, Frankfurt am Main, Germany) were selected: 0.03, 0.06, and 0.1 wt%. Medical-grade PCL (12.0424 g) was placed into Griffin beakers lined with aluminum foil. Using Equation (3) and the known mass of PCL, the required HT amounts were calculated to be 0.0036, 0.0072, and 0.0121 g for 0.03, 0.06, and 0.1 wt%, respectively. Three samples were prepared for each concentration.

$$wt={\displaystyle \frac{m_{HT}}{m_{HT}+m_{PCL}}}\times 100⟺m_{HT}={\displaystyle \frac{wt\times m_{PCL}}{100-wt}}$$

(3)

After all the measurements, the samples were incubated at 100 °C for 40 min in a Drying Oven VWR^® VENTI-LINE^® (VWR, Radnor, PA, USA) with forced convection. Then, the beakers were removed from the incubator and left at room temperature until PCL solidification.

Since the main goal was to evaluate the additive’s influence on the diameter and mechanical properties of PCL, it was decided to proceed with this study by only printing meshes with fiber diameters of 260 µm to avoid material waste.

2.9. Cytotoxicity Evaluation of MEW Meshes

Cytotoxicity assays were conducted through the indirect contact assay according to ISO 10993-5:2009 [36]. To do so, the meshes were cut (8 mm diameter) and sterilized using two methods. One set of samples was sterilized by exposure to UV light (254 nm) for 30 min (termed UV samples). The other set was disinfected by immersion in a 70% EtOH bath for 20 min, washed with Phosphate Buffered Saline (PBS) for 5 min (3 times), followed by sterilization by UV light for 30 min (termed as ET + UV samples).

For cell culture, human neonatal dermal fibroblasts (hNDFs) isolated from human neonatal foreskin samples (Coriell Institute for Medical Research, Camden, NJ, USA) were cultured in Dulbecco’s Modified Eagle Medium (DMEM) supplemented with 10% fetal bovine serum (FBS, Gibco, Frederick, MD, USA) and antibiotics (1% v/v of both penicillin/streptomycin and amphotericin B (Sigma-Aldrich, Saint Louis, MO, USA)). Cells were cultured in 5% CO₂ at 37 °C in tissue culture polystyrene flasks. For transferring the cells to the 24-well plate, cells were trypsinized when reaching 80% confluence using 0.05 wt% trypsin/ethylenediamine tetraacetic acid (EDTA) solution (Sigma-Aldrich) and centrifuged at 1200 rpm for 5 min.

Cells were seeded into a treated 24-well plate ($4\times {10}^4$ cells/400 µL DMEM) and cultured for 24 h prior to the replacement of culture medium with the supernatants from the meshes. To obtain the supernatants, samples were incubated in culture medium for 24 h. After this incubation period, the medium in contact with the cells was replaced with the mesh-conditioned medium, and the cells were incubated for additional 24 h. Two control testing groups were used: (1) cells cultured in standard DMEM and (2) cells cultured in culture medium previously incubated with untreated PCL meshes (HT free).

To evaluate the metabolic activity, meshes were incubated in DMEM containing 20% (v/v) resazurin sodium salt (Sigma-Aldrich) for 2 h at 37 °C, followed by the analysis of fluorescence using a microplate reader (Synergy MX, BioTek, Winooski, VT, USA) at 530 nm (excitation) and 590 nm (emission) [37].

2.10. Statistical Analysis

The data obtained from the resazurin assay were expressed as metabolic activity. Statistical analysis was performed using one-way analysis of variance (ANOVA), and comparisons between different HT concentrations and sterilization methods were assessed using the nonparametric Mann–Whitney U test. A p-value of less than 0.05 was considered statistically significant. For each condition, a minimum of three specimens was analyzed. All statistical analyses were carried out using IBM SPSS Statistics, version 26.0.

3. Results

3.1. Technical-Grade PCL Calibration

3.1.1. Temperature

Firstly, the printing temperature was optimized while keeping all other parameters constant. The temperature varied between 160 °C and 220 °C, as shown in Table 1, while the collector speed was fixed at 1000 mm/min and the applied voltage at 3.63 kV.

It was observed that the temperature significantly impacts the fiber quality and the material viscosity. At a lower temperature (160 °C), there is high viscosity, leading to fiber discontinuities and material accumulation (Figure 4). When increasing the temperature to 180 °C, discontinuities and accumulation were reduced but were still present. At 200 °C, most fibers were well printed with minimal errors. Similarly, at 220 °C, the printed meshes resembled those at 200 °C, with only a few minor errors. Based on the obtained results and to avoid excessively low viscosity, a printing temperature of 200 °C was selected.

3.1.2. Collector’s Speed

Next, the collector speed was optimized while keeping the applied voltage fixed at 3.63 kV and the temperature at 200 °C, based on the best results obtained in the previous step.

After printing the meshes at 500 mm/min, most fibers were curved, and the polymer jet was unstable. Increasing the speed to 1000 mm/min reduced fiber deviations, while speeds of 1200 mm/min and 1400 mm/min produced straight, stable fibers. However, at 1800 mm/min, discontinuities appeared, with skipped lines and material deposition issues.

A non-straight fiber deposition and an unstable polymer jet indicate low collector speeds (500 and 1000 mm/min). On the other hand, discontinuities suggest that the material deposition rate is lower than the collector’s speed, causing skipped sections in the final product, meaning that 1800 mm/min is too high. Among the remaining viable speeds, both 1200 and 1400 mm/min produced straight and continuous fibers; however, 1200 mm/min was selected to minimize the risk of discontinuities.

3.1.3. Voltage

Finally, the applied voltage was optimized by varying it between 2.63 kV and 4.63 kV, as shown in Table 1. The temperature and collector speed were kept constant at 200 °C and 1200 mm/min, respectively.

At lower voltages, 2.63 kV and 3.03 kV, material accumulation at the nozzle caused fiber discontinuities, likely due to insufficient force to maintain a stable polymer jet. Increasing the voltage to 3.23 kV, 3.43 kV, and 3.63 kV stabilized the polymer jet, resulting in mostly continuous fibers with minor errors. Voltages above 4 kV (4.03 kV and 4.63 kV) led to fiber instabilities and discontinuities.

The voltage applied determines the material extrusion. If the voltage is too low, the electric field is weak and will not be able to extrude all the material, leading to material accumulation (2.63 and 3.03 kV). If the voltage is too high, the extrusion speed is too high, leading to points of discontinuity and instability (4.03 and 4.63 kV). These findings suggest that the optimal voltage range for achieving stable and continuous fiber formation lies between 3.23 kV and 3.63 kV. To make a safe choice and avoid ending up with an electrical field that is too strong or too weak, a voltage of 3.43 kV was selected.

3.1.4. Diameter and Pore Size Evaluation

As a final step of calibration, it was necessary to measure all the obtained fiber diameters, as well as the pore sizes. The pore sizes will indicate whether the geometry is correctly printed or not, and the fiber diameter values will indicate if the printer is extruding the right amount of material (Figure 4).

For each sample, six different measurements were taken, and the mean value of these measurements is present in

Table 2. Pore size and fiber diameter evaluation for technical-grade PCL calibration.

	Pore			Fiber Diameter
	260 µm	200 µm	160 µm	260 µm	200 µm	160 µm
Sample 1 (µm) Sample 2 (µm) Sample 3 (µm) Sample 4 (µm)	1483.17 1505.40 1527.60 1494.38	1485.11 1562.84 1547.89 1509.51	1520.66 1576.94 1583.67 1624.31	160.00 144.35 140.27 140.18	120.92 113.74 117.39 116.42	91.47 94.11 96.67 100.55
Mean (µm) Error (%)	1502.64 0.18	1526.34 1.76	1576.40 5.09	146.20 43.77	117.12 41.44	95.70 40.19
Mean error (%)		2.34		41.80

for the pore and fiber diameter size. With these results, it is possible to calculate the mean value for each pore/diameter, calculate the mean error, and analyze. For the diameter values, it is also possible to adjust Equation (1) and calibrate the device.

The predefined pore size was 1.5 mm (1500 µm), and based on the results in Table 2, the mean error is not significant (<10%), indicating that the geometry is well defined. However, regarding the fiber diameter, all the obtained values differ from the desired diameters of 260 µm, 200 µm, and 160 µm, with a mean error of 41.8%, indicating that the device was not extruding the correct amount of material.

3.1.5. Final Parameter Adjustments

Since the device was not extruding the correct amount of material, Equation (1) was adjusted by multiplying it by a correction factor: ${\displaystyle \frac{100}{100-41.8}}={\displaystyle \frac{100}{58.2}}=1.718$.

After updating all the necessary G-code files, new meshes were printed and their fiber diameters were evaluated using the same procedures as previously described. The results presented in

Table 3. Diameter evaluation for technical-grade PCL calibration, after adjustments.

	260 µm	200 µm	160 µm
Sample 1 (µm) Sample 2 (µm) Sample 3 (µm) Sample 4 (µm)	244.05 242.59 242.01 255.72	190.55 189.51 189.33 192.37	156.47 157.59 156.53 152.80
Mean (µm) Error (%)	246.09 5.35	190.44 4.78	155.85 2.60
Mean error (%)		4.24

show that the meshes have diameters close to the desired ones, with a mean error of 4.24%, which is considered not significant (<10%).

3.2. Medical-Grade PCL Calibration

In this section, the device was calibrated for medical-grade PCL (in pellet form), following the same procedure described in Section 3.1. Calibration was performed using square meshes; however, after identifying the final set of printing parameters, sinusoidal meshes were also printed to confirm their adequacy.

3.2.1. Initial Printing Parameters

The initial printing parameters were selected based on the results obtained with technical-grade PCL: a temperature of 200 °C, a collector speed of 1200 mm/min, and an applied voltage of 3.43 kV. Visual inspection of the resulting meshes indicated that they were well printed and comparable to those obtained during the technical-grade PCL calibration. Therefore, it was deemed unnecessary to perform single-parameter variation, and the process proceeded directly to the diameter evaluation step.

3.2.2. Diameter Evaluation

As in Section 3.1, all fiber diameters were measured (

Table 4. Diameter evaluation for medical-grade PCL calibration.

	260 µm	200 µm	160 µm
Sample 1 (µm) Sample 2 (µm) Sample 3 (µm) Sample 4 (µm)	256.05 239.04 244.30 240.20	192.78 190.89 190.08 190.50	157.22 159.31 158.31 158.85
Mean (µm) Error (%)	244.90 5.81	191.06 4.47	158.42 0.99
Mean error (%)		3.75

) to verify whether the printer was extruding the correct amount of material. Pore size was not measured, as the mesh geometry had already been calibrated and was well defined.

Analysis of Table 4 showed that no adjustment to Equation (1) was necessary, as the observed error in fiber diameter was not significant (<10%). Sinusoidal meshes were also measured, and similarly, the error was considered negligible. Therefore, the initial printing parameters—200°C, 1200 mm/min, and 3.43 kV—were confirmed to be suitable for printing both square and sinusoidal mesh geometries.

It was also observed that the mesh geometry was optimal in the central region, with well-defined and uniform squares. However, deviations were noted at the corners, where the squares were not equilateral (Figure 5A). To address this issue, the mesh design was modified (Figure 5B), resulting in uniform and fully usable samples measuring 35 × 35 mm, thereby reducing material waste.

3.3. PCL/HT Meshes: Fabrication and Testing

This section describes the process of printing the PCL/HT meshes, as well as measuring their diameters. Then, mechanical tests were performed to evaluate the meshes’ behavior under load, as well as cytotoxicity tests to evaluate the toxicity of these meshes.

3.3.1. Diameter Evaluation of the PCL/HT Meshes

The control meshes had an average fiber diameter of approximately 280 µm, slightly higher than the intended 260 µm. To avoid re-adjustment, the study proceeded under the condition that all control meshes remained consistent and exhibited similar diameters.

Table 5. Diameter evaluation of all wt% samples and calculation of the reduction of the fiber.

	Square Meshes		Sinusoidal Meshes
	Mean Diameter (µm)	Reduction (%)	Mean Diameter (µm)	Reduction (%)
Control (0 wt%)	281.30	-	273.62	-
0.03 wt% 0.06 wt% 0.1 wt%	234.04 153.92 98.41	16.80 45.28 65.02	234.05 165.35 95.07	14.46 39.57 65.26

presents the mean diameter values for each sample, along with the corresponding reductions in fiber diameter relative to the control group. Analysis of the data confirmed that the incorporation of the antistatic agent produced the intended effect, enhancing polymer stretchability and resulting in the formation of thinner fibers. The improved printing stability observed after HT incorporation can be attributed to the antistatic properties of the additive. HT, an amphiphilic molecule, reduces charge accumulation on the polymer surface by enhancing surface conductivity and facilitating uniform charge dissipation. This minimizes electrostatic disturbances during the MEW process, resulting in a more stable polymer jet and reducing fiber defects such as whipping, misalignment, and discontinuities. Consequently, this enables the production of thinner, more uniform fibers with consistent morphology and geometry across the printed mesh.

3.3.2. Uniaxial Tensile Testing

• Influence of HT concentration on the meshes’ behavior under load

For each concentration and mesh geometry, seven samples were tested. From the resulting stress–strain data, mean values were calculated and used to generate the stress–strain curves shown in Figure 6 for the square (A) and sinusoidal (B) meshes, respectively. The vaginal tissue curve was obtained from Vaz et al. [38].

To correctly analyze these results, it is necessary to introduce two terms, comfort zone and safety zone, which are associated with the expected strains experienced by the vaginal tissue. The comfort zone pertains to stresses encountered during daily activities and has a maximum value of 20% strain. The safety zone relates to high-stress peaks and is set with a threshold of up to 40% strain [16]. The following results will focus only on the comfort zone.

By analyzing Figure 6, it is evident that the pore geometry of the meshes affects both their behavior and the maximum load they can support. From Figure 6B, it can be inferred that the sinusoidal mesh best mimics the mechanical behavior of vaginal tissue, particularly the control mesh, which withstands comparable stress levels up to a strain of 0.15. However, when considering higher strain values, native tissue can support greater stress values. Thus, the slope of the vaginal tissue curve is steeper than that of the control mesh, indicating that the tissue is stiffer than the mesh.

Another conclusion from the graphics is that the load-bearing capacity is also influenced by fiber diameter, meaning that thicker fibers can support higher stress. An individual analysis of each graph confirms that this trend holds across all geometries. This relationship is further supported by the findings in

Table 6. The tensile stress of each mesh within the comfort zone.

	Square	Sinusoidal	Tensile Stress (MPa)
	Printed Fiber Diameter	Printed Fiber Diameter	Square	Sinusoidal
Control	281.30	273.62	1.758	0.317
0.03 wt% 0.06 wt% 0.1 wt%	234.04 153.92 98.41	234.05 165.35 95.07	1.206 1.009 0.457	0.104 0.022 0.004

, which show that as fiber diameter decreases, the stress required to induce a specific deformation also decreases. This behavior is attributed to the increased surface area-to-volume ratio of thinner fibers, which makes them more sensitive to external forces.

Table 6 presents the maximum stress each mesh can support, referred to as tensile strength. The results confirm a reduction in tensile strength as the fiber diameter decreases, which is caused by an increase in HT concentration.

• HT concentration influence on the meshes’ mechanical properties

To assess whether the antistatic agent alters the properties of pure PCL, the obtained curves (PCL/HT) could be compared with those obtained from pure PCL meshes as long as they have similar diameters and the same geometry (Figure 7). The curves available for PCL meshes were obtained from [38], being the square meshes made of medical-grade PCL, with diameters of 240 and 160 and a pore size of 1.5 mm. It was not possible to compare the curve of the 0.1 wt% mesh (98.41 µm) because its diameter did not closely match any available curve.

Based on the results of the uniaxial tensile tests, the stiffness and tensile strength of the meshes were calculated and are presented in Figure 7B. By comparing the 0.03 wt% curve with the 240 µm one, it would be expected that the pure PCL mesh would support a higher load and thus have a greater tensile strength, as its diameter is larger than that of the PCL/HT mesh. However, the tensile strength of the PCL/HT mesh exceeds that of the pure PCL mesh by 0.152 MPa. The same trend is observed when comparing the 0.06 wt% curve with the 160 µm one, as the difference in tensile strength between the PCL/HT and pure PCL meshes is 0.196 MPa, with the PCL/HT mesh supporting a higher load despite its smaller diameter.

It is evident that the PCL/HT meshes not only exhibit an increase in tensile strength but also an increase in Young’s Modulus, indicating enhanced stiffness. The increased tensile strength and stiffness with the incorporation of HT suggest that the interaction between PCL and HT modifies the mechanical properties of the PCL.

3.3.3. Cytotoxicity Evaluation

Given that the meshes are designed for the application of POP repair, it is essential to evaluate not only how they would behave under load but also their cytotoxicity. The results of cellular assays (Figure 8) confirm that none of the tested meshes exhibit cytotoxicity, as all maintain cell viability higher than 90%, which is above the 70% threshold defined by ISO 10993-5:2009.

Statistical analysis between different HT concentration groups and the control group (cells in the culture medium) showed no significant differences in cell metabolic activity (p > 0.05), indicating that the tested concentrations are non-cytotoxic. Moreover, the influence of the sterilization method on the cytotoxicity of the meshes was also analyzed. Statistically significant differences were not observed between sterilization methods, indicating they are suitable for this application.

The analysis of Figure 8 confirms that none of the tested samples exhibit cytotoxicity, as all maintain cell viability above 90%, well above the 70% threshold defined by ISO 10993-5:2009. Statistical comparisons between different HT concentration groups and the control group showed no significant differences in cell metabolic activity (p > 0.05), indicating that the tested concentrations are non-cytotoxic. This result was consistent across both sterilization methods (ET + UV and UV), indicating the reliability of the findings.

Statistical analysis showed no significant differences in metabolic activity among the different HT concentrations within each sterilization method (p > 0.05), indicating that all PCL/HT meshes were non-cytotoxic. In addition, no statistically significant differences were found between the UV and EtOH + UV sterilization methods for any given concentration, including the control group. These results confirm that both sterilization protocols are equally suitable for this application.

3.4. Differential Scanning Calorimetry

To gain deeper insight into the influence of HT on PCL, particularly regarding its crystallinity and potential structural changes, a DSC test was performed to analyze the thermal properties of the material.

The addition of HT progressively alters the thermal behavior of PCL, as seen in the broadening and shifting of the melting peaks (Figure 9A). These changes suggest that HT modifies the crystallization process, leading to finer or more heterogeneous crystalline domains. The peak area analysis further supports this (Figure 9B), showing that lower HT concentrations (0.03 wt% and 0.06 wt%) reduce crystallinity, while 0.1 wt% HT results in a peak area comparable to the control samples.

The curve for 0.1 wt% HT reveals differences when compared to the other concentrations, possibly highlighting optimal stability and crystallinity, which corresponds with improved mechanical properties. This implies that 0.1 wt% HT might act as a threshold concentration where HT exerts the most substantial influence on the thermal and crystalline characteristics of the polymer without oversaturating or disrupting its structure. This suggests that it may function as a nucleating agent, facilitating the development of uniform or enhanced crystalline regions.

Despite these alterations, the melting temperatures across all samples remain consistent, indicating that HT does not compromise the polymer’s thermal stability, which is essential for biomedical applications. Furthermore, the broadening of melting peaks at higher HT concentrations suggests a range of crystalline and amorphous areas, likely contributing to increased tensile strength and stiffness. These results illustrate that HT can be used to optimize the thermal and mechanical properties of PCL meshes, making them suitable for applications such as POP repair.

4. Discussion

The optimization of the MEW printing parameters (temperature, collector speed, and applied voltage) was fundamental to ensure the formation of continuous, well-defined fibers. During the calibration process, it was observed that lower temperatures or voltages led to polymer accumulation at the nozzle, whereas excessively high voltages or collector speeds resulted in fiber discontinuities. The optimal parameters found—200 °C, 1200 mm/min collector speed, and 3.43 kV—are consistent with those reported in the literature for high-resolution PCL printing [39], supporting their suitability for MEW-based fabrication.

The incorporation of the antistatic agent HT improved jet stability and fiber deposition during printing. The amphiphilic nature of HT enhances surface conductivity by facilitating charge dissipation, reducing jet whipping, and material accumulation at the nozzle. These effects enabled a reduction in fiber diameter of up to 66%, while maintaining excellent control over mesh geometry. These findings align with prior studies using antistatic agents in melt electrospinning and thermoplastics, where increased ionic conductivity was shown to reduce fiber diameter and improve print fidelity [25].

Interestingly, although the HT-loaded meshes had smaller fiber diameters, they displayed enhanced mechanical properties—specifically, increased tensile strength and stiffness. This is in contrast to traditional behavior in PCL systems, where thinner fibers often result in reduced mechanical resistance. Our results suggest that HT not only improves printability but also modifies the internal structure of PCL, potentially increasing crystallinity or chain alignment. This hypothesis is supported by DSC results, which showed broadened melting peaks at higher HT concentrations, indicating heterogeneous or altered crystalline domains.

In a recent study [40], a polymer-type antistatic agent called TMQ, prepared from TPEG, MAH, and a quaternary ammonium salt (QAS), was mixed with polypropylene (PP) to evaluate its effects. Incorporating 15 wt% TMQ led to a 38.6% reduction in tensile strength and decreased crystallinity, while also lowering surface resistivity, indicating improved surface conductivity. In a separate study [41], 3, 6, and 9 wt% of Irgastat P 18 was added to PP/OMMT nanocomposites, causing a slight decrease in tensile strength with increasing additive concentration. Another investigation [42] on the ionic antistatic agent Allyl Trimethylammonium Chloride (TMA) in PMMA found that 1.25 wt% TMA reduced surface resistance by a factor of five and increased tensile strength by 38%. Compatibility with PMMA was further enhanced by adding Acrylic Acid (AA), which improved hydrogen bonding through increased molecular polarity.

The previous studies showed that the incorporation of antistatic agents in polymers modifies their properties, namely tensile strength, surface resistivity, and crystallinity. In this study, the tensile strength increased by 14.42% and 24.10% for 0.03 wt% and 0.06 wt% of HT, respectively, when compared with PCL meshes with similar diameters. There was also an increase in Young’s Modulus of 38.68% and 55.59% for 0.03 wt% and 0.06 wt% of HT, respectively. These results suggest that the interactions between PCL and HT promote hydrogen bonding between molecular chains, and the crystallization of the polymer may also be affected. When the crystallization level of a polymer increases, the intermolecular forces are stronger, so the movement of the polymer chains is reduced, and, therefore, the stiffness of the polymer increases [40]. By verifying an increase in the Young’s Modulus values of the PCL/HT meshes, it can be concluded that the polymer’s stiffness increased with HT incorporation, suggesting that the polymer’s crystallization was also modified.

Ultimately, it is necessary to consider whether it is advantageous or not to incorporate HT in PCL meshes. On the one hand, its incorporation will control the charge’s distribution on the material surface, resulting in a more stable and controlled jet that is stretched uniformly, leading to thinner fibers. On the other hand, this HT incorporation resulted in increased tensile strength and stiffness. For the specific case of POP repair, it is ideal to have a mesh that provides good support and can withstand high loads while still providing flexibility, avoiding problems like stress shielding.

Incorporating the adequate HT concentration in PCL meshes enables achieving the ideal mesh stiffness and tensile strength. Since the influence of HT on fiber diameter is already understood, selecting an appropriate HT concentration allows the printed PCL/HT mesh to achieve a diameter capable of supporting loads comparable to native tissue. Furthermore, this would increase the mesh’s stiffness to the point that the mesh mimics the vaginal tissues’ characteristics and behavior. Regarding the stiffness and the stress shielding problem, it is also a fact that, unlike synthetic meshes, it is expected that these meshes will be absorbed after 2 to 3 years of implantation, so even if the mesh is slightly stiffer than the native tissue, it would not be such a problem as it is with the synthetic meshes.

These meshes are designed for POP repair, making it crucial to assess both their mechanical behavior under load and their cytotoxicity. Cytotoxic test results showed that all tested concentrations of HT (up to 0.1 wt%) maintained cell viability above 90%, fulfilling the ISO 10993-5 standard for non-cytotoxic biomaterials. While other studies [43] using pellets reported cytotoxicity at 0.1 wt%, our results with printed mesh structures suggest that the final product’s geometry and surface area can mitigate adverse effects. These findings are in line with previous research using PCL scaffolds and confirm the safe application of low-dose HT in biomedical contexts.

The evaluation of two sterilization methods (UV and EtOH+UV) showed no statistically significant differences in cell viability across all HT concentrations and the control group. Horakova et al. [28] found that neither method affected the morphology, molecular weight, or thermal properties of electrospun PCL nanofibers, though EtOH soaking resulted in higher cell viability. Łopianiak et al. [44] reported similar findings on polyurethane scaffolds but noted that neither method alone ensured complete sterilization. Although prior studies have reported higher biocompatibility with ethanol pre-treatment [28], differences in soaking time and the absence of post-rinse steps may explain the uniform outcomes observed in our study. In this context, both sterilization methods proved effective and suitable for MEW-based PCL/HT meshes.

The incorporation of the antistatic agent HT into PCL improved mesh performance for POP repair by enabling the fabrication of finer, more uniform fibers, enhancing print resolution and mechanical properties such as tensile strength and stiffness. These improvements result in meshes that better mimic the mechanical behavior of vaginal tissue. Additionally, thinner fibers reduce the total amount of implanted material, which may minimize the foreign-body response and improve tissue integration. Combined with confirmed biocompatibility, these features highlight the potential of antistatic melt-electrowritten biodegradable meshes as advanced candidates for POP treatment.

5. Conclusions

This study investigates PCL biodegradable meshes as a potential solution for POP. It explores how incorporating the antistatic agent HT into PCL meshes influences their mechanical properties, fiber structure, and biocompatibility. HT incorporation not only leads to a significant reduction in fiber diameter but also enhances the meshes’ structural integrity and stiffness when compared against PCL meshes with similar diameters. The cytotoxicity analysis shows that all PCL/HT meshes remain non-toxic and that the tested sterilization methods effectively sterilize the meshes. DSC tests indicate that HT maintains the polymer’s thermal stability. Additionally, the broadened melting peaks observed at higher HT levels indicate varied crystallinity, leading to enhanced tensile strength and stiffness. These findings favor the applicability of PCL/HT meshes for POP repair, offering mechanical properties comparable to vaginal tissue while avoiding stress shielding and remaining non-toxic.

However, further research is warranted to address specific limitations of this work. Antimicrobial activity was not evaluated, which is relevant for ensuring sterility in clinical applications. The range of HT tested concentrations was also limited, and the mechanical and cytotoxicity tests were conducted in simplified in vitro settings that do not fully mimic the complexity of the in vivo environment. Additionally, long-term degradation and biointegration studies should be performed in the future. Despite these limitations, the results demonstrate the practical potential of PCL/HT meshes in POP repair. The improved mechanical performance, thinner fibers, and confirmed biocompatibility make these meshes strong candidates for clinical translation. Further work should include Scanning electron microscopy (SEM), X-ray diffraction (XRD), and rheological analyses to better characterize structural changes and degradation, as well as in vitro and in vivo studies to confirm tissue integration and biological response. Furthermore, real-time analysis of deformation processes during mechanical testing and microstructural evaluation of fracture surfaces (e.g., via SEM) should be included in future studies to better understand local failure mechanisms, fiber slippage, and pore evolution under load.

Future studies will incorporate uniaxial mechanical testing of PCL/HT meshes under hydrated conditions using a humidity chamber recently acquired for our testing system. This setup will allow testing in fluid media (e.g., PBS at 37 °C), enabling more accurate simulation of the in vivo vaginal environment and better comparison with native tissue behavior. Future research endeavors can also further explore the structural changes in PCL fibers with HT incorporation. SEM, rheological analysis, and XRD tests are also recommended to gain deeper insights into the material’s behavior and performance. In vitro and in vivo studies should also be performed before real-world applicability to study PCL/HT bioactivity, its degradation mechanism, and its impact on tissue regeneration.