3D-Printed PLDLA–TMC/PEG 400 Vascular Scaffolds with a Poly(hexamethylene Biguanide) Antibacterial Coating

Abstract

Background: Synthetic vascular scaffolds often exhibit limited mechanical performance and low hydrophilicity, which compromise early vascular integration and increase susceptibility to bacterial colonization. This study developed 3D-printed scaffolds based on poly(L-co-D,L-lactide)–poly(trimethylene carbonate) (PLDLA–TMC) with polyethylene glycol 400 (PEG) to modulate mechanical and interfacial properties and coated with poly(hexamethylene biguanide) (PHMB) to confer antibacterial activity. Methods: PLDLA–TMC scaffolds modified with PEG 400 and coated with PHMB were prepared and systematically characterized to assess their structural, thermal, mechanical, and antimicrobial properties. PHMB coatings (3%, 6%, and 12% w/w in hydroxypropyl methylcellulose, HPMC) were applied and evaluated for drug release, cytotoxicity, and activity against Staphylococcus aureus. Biocompatibility was tested in an endothelial cell and myoblast co-culture. Results: Incorporation of 2% PEG increased the tensile strength from 0.14 ± 0.10 MPa for scaffolds containing 0.5% PEG to 0.79 ± 0.12 MPa and promotes a more elastic scaffold behavior. PHMB at 12% caused cytotoxicity (7.70 ± 0.37% cell viability). The 3% PHMB coating produced a 12.5 ± 0.1 mm inhibition zone but exhibited burst release within 1 h, whereas the 6% coating maintained cell viability (72.95 ± 1.10%), produced a 13.1 ± 0.2 mm inhibition zone, and provided sustained antimicrobial release over 7 days. Scaffolds supported organized adhesion and proliferation of endothelial cells and myoblasts. Conclusions: 3D-printed PLDLA–TMC scaffolds containing 2% PEG and coated with 6% PHMB combined improved mechanical performance, sustained antimicrobial release, antibacterial activity, and biocompatibility in an in vitro vascular model.

1. Introduction

Cardiovascular diseases are the leading cause of morbidity and mortality worldwide. In 2023, the global burden reached 626 million cases, including 239 million cases of ischemic heart disease and 122 million cases of peripheral artery disease of the lower limbs. In the same year, these conditions accounted for 19.2 million deaths [1]. The progression of vascular pathologies can lead to vessel obstruction or stenosis, necessitating procedures such as bypass or vascular replacement. In these scenarios, the use of autologous grafts is limited by tissue availability and donor site trauma, whereas allogeneic grafts carry the risk of immune rejection [2].

To circumvent these limitations, artificial vascular scaffolds have been developed as an alternative when autologous or allogeneic grafts are not available. Synthetic materials currently used clinically, such as e-PTFE, PET, and polyurethane, exhibit suboptimal clinical outcomes in terms of compliance mismatch, hemocompatibility, and remodeling capacity, particularly in small-diameter grafts. These limitations have driven the investigation of biodegradable polymers capable of providing initial mechanical support and exhibiting controlled degradation over time, allowing the progressive replacement of the scaffold with neovessel tissue [3]. Current strategies for engineered vascular grafts focus on the development of bioresorbable and biofunctionalized tubular scaffolds to overcome the limitations of conventional materials and support tissue regeneration and integration [4].

The biodegradable terpolymer poly(L-co-D,L-lactic acid-co-trimethylene carbonate) (PLDLA–TMC) offers a lower glass transition temperature and superior flexibility compared to polylactic acid homopolymers, in addition to allowing for the modulation of stiffness and degradation rate by varying the lactide and trimethylene carbonate fractions [5]. In vascular scaffolds, these properties contribute to improved mechanical compatibility with the native vessel and maintenance of structural integrity during the early post-implantation phase, supporting lumen preservation and blood flow during the critical period of graft mechanical adaptation [6].

PLDLA–TMC can be processed via extrusion-based 3D printing to obtain porous structures of interest for tissue engineering; however, like other biodegradable polyesters, it exhibits limited hydrophilicity. In this context, the incorporation of low molecular weight polyethylene glycol (PEG 400) by physical incorporation has been explored as a strategy to increase surface hydrophilicity, inferred from the chemical nature of the PEG component, and modulate scaffold interfacial properties, with potential effects on material–physiological medium and material–cell interactions while maintaining the initial mechanical support provided by the continuous PLDLA–TMC matrix [7].

The incorporation of antimicrobial properties into biomaterials is of clinical importance, as infections associated with vascular materials remain a significant challenge, with reported incidence rates ranging from 1% to 6% in arterial bypass procedures. Pathogens such as Staphylococcus aureus and Staphylococcus epidermidis adhere to polymeric surfaces and form biofilms, complicating treatment and increasing the risk of early thrombosis and graft failure [8]. Given the vulnerability of biodegradable vascular scaffolds to microbial colonization, the use of antimicrobial coatings has been explored as a strategy to reduce bacterial adhesion. Among these agents, poly(hexamethylene biguanide) (PHMB) exhibits broad-spectrum activity against pathogens critical to vascular infections. While Gram-positive bacteria, such as Staphylococcus aureus and S. epidermidis, are primary agents in early-stage graft failure, PHMB also targets relevant Gram-negative strains, including Escherichia coli and Pseudomonas aeruginosa. Consequently, its incorporation as a functional coating offers a strategic approach to enhance scaffold safety without compromising their mechanical or interfacial properties [9].

In light of these mechanical, structural, and microbiological challenges, this study presents a proof-of-concept development of 3D-printed vascular scaffolds based on PEG 400-modified PLDLA–TMC and coated with PHMB to integrate architectural control, tunable mechanical properties, antibacterial functionality, and biocompatibility in vitro.

2. Materials and Methods

2.1. Polymer Synthesis and Modification

2.1.1. PLDLA–TMC Synthesis

The synthesis of PLDLA–TMC was carried out following the procedure described by Motta and Duek [10]. Initially, L-co-D,L-lactide and trimethylene carbonate (TMC) were mixed in glass ampoules at a molar ratio of 80/20 using stannous octoate (Sn(Oct)₂; Sigma-Aldrich, St. Louis, MO, USA) as a catalyst. The ampoules were then immersed in an oil bath at 130 °C for 48 h under vacuum. The resulting terpolymer was dissolved in chloroform (Sigma-Aldrich, St. Louis, MO, USA), precipitated in methanol (Sigma-Aldrich, St. Louis, MO, USA), and subsequently dried under vacuum at 45 °C for 24 h to remove any residual solvent.

The 80/20 molar ratio of lactide to TMC was chosen based on prior reports indicating a balance between mechanical flexibility and degradation rate suitable for vascular scaffolds [11].

2.1.2. Incorporation of PEG 400 into PLDLA–TMC

To obtain the PLDLA–TMC + PEG formulations, PEG 400 was first dissolved in chloroform, followed by the addition of PLDLA–TMC. The mixture was kept under magnetic stirring for 2 h, poured into a container to dry at room temperature for 48 h, and subsequently vacuum-dried to minimize residual chloroform. Two polymer blends were prepared with PEG concentrations of 0.5% and 2.0% (w/w relative to the PLDLA–TMC mass). The schematic representation of the polymer is shown in Figure 1.

2.2. Physicochemical Characterization of the Polymer

2.2.1. Gel Permeation Chromatography (GPC)

The number-average (Mn) and weight-average (Mw) molecular weights, as well as the polydispersity index (PDI), of the PLDLA–TMC/PEG blends were determined by gel permeation chromatography, providing apparent molecular weights associated with the hydrodynamic behavior of the blends in solution. The system (Waters Corporation, Milford, MA, USA) consisted of a pair of 7.8 × 300 mm Styragel HR columns (5 μm particle size), maintained at 35 °C, and equipped with a Waters 2414 refractive index detector. Tetrahydrofuran (THF) was used as the mobile phase. Polymer samples were injected at a concentration of 3.0 mg/mL and eluted at a flow rate of 1.0 mL/min. Calibration was performed using monodisperse polystyrene standards, with certified molecular weights and narrow molecular weight distributions (PDI ≈ 1.02), selected to cover the elution range of the analyzed samples. Molecular weight values were determined by interpolation of the sample elution volumes on the calibration curve.

2.2.2. Fourier Transform Infrared Spectroscopy (FTIR)

FTIR spectra of PLDLA–TMC and PLDLA–TMC + PEG samples were acquired using a Spectrum 65 spectrometer (PerkinElmer, Waltham, MA, USA) equipped with an attenuated total reflectance (ATR) accessory. Each spectrum was collected by averaging 32 scans at a spectral resolution of 4 cm⁻¹ within the range of 4000–600 cm⁻¹.

2.2.3. Differential Scanning Calorimetry (DSC)

DSC analyses of PLDLA–TMC and PLDLA–TMC + PEG samples were performed using a DSC 25 calorimeter (TA Instruments, New Castle, DE, USA). The samples were initially heated from −50 °C to 200 °C at a heating rate of 10 °C·min⁻¹ under a nitrogen flow of 20 mL·min⁻¹. Subsequently, they were cooled to −50 °C at the same rate and held at this temperature for 1 min. A second heating scan was then conducted up to 200 °C using the same heating rate and nitrogen flow conditions.

2.2.4. Thermogravimetric Analysis (TGA)

Thermogravimetric analysis of PLDLA–TMC and PLDLA–TMC + PEG samples was performed using a TA Instruments^® TGA 55 system (New Castle, DE, USA). Samples were heated from 25 °C to 400 °C at a rate of 10 °C/min under a continuous nitrogen flow of 100 mL/min.

2.2.5. Rheological Analysis

Rheological measurements of PLDLA–TMC + PEG were conducted using a TA Instruments^® DHR-2 rheometer (New Castle, DE, USA) with a 25 mm parallel-plate geometry. Molten polymer was loaded onto the parallel plates before gap adjustment. The material’s rheological properties were assessed under both oscillatory and steady-state conditions. For the amplitude sweep, the strain varied from 0.001% to 20% at a constant angular frequency of 10 rad/s, identifying the linear viscoelastic region (LVR). A strain of 0.2% was therefore used in all oscillatory tests. Frequency sweep tests were performed to determine the storage (G′) and loss (G″) moduli, with angular frequency ranging from 0.1 to 100 rad/s. Steady-shear flow tests were conducted to obtain viscosity curves with shear rates ranging from 0.001 to 1000 s⁻¹. All measurements were performed at 80 °C to ensure complete polymer melting and homogeneous flow conditions.

2.2.6. Zeta Potential (ζ)

The zeta potential (ζ) of PLDLA–TMC + PEG samples was measured using a SurPASS electrokinetic analyzer (Anton Paar GmbH, Graz, Austria) over a pH range of 3.5–10.5 at room temperature. This pH range exceeds the physiological conditions to assess surface charge behavior. The pH values were adjusted using 0.05 M HCl or 0.05 M NaOH solutions.

2.3. Scaffold Printing

PLDLA–TMC + PEG scaffolds were fabricated using a 3D bioprinter (Octopus™, 3D Biotechnology Solutions-3DBS, São José dos Campos, SP, Brazil) via extrusion-based printing. Scaffold design was performed using Slic3r software (version 1.3.0, open-source software), where printing parameters and specimen dimensions were defined: 5 mm outer diameter, 3 mm inner diameter, and 10 mm length. Scaffolds of both formulations (PLDLA–TMC + 0.5% and PLDLA–TMC + 2.0%) were printed under the same conditions. The detailed printing parameters are summarized in

Table 1. Operational parameters for extrusion-based 3D printing.

Parameters	Settings
Extrusion temperature	100 °C
Build platform temperature	40 °C
Print speed	4 mm/s
Nozzle diameter	0.4 mm

.

2.4. Tensile Mechanical Testing

To evaluate the mechanical properties of 3D-printed PLDLA–TMC + PEG scaffolds (0.5% and 2.0% PEG), tensile tests were performed following ASTM D638 [12]. Experiments were conducted in quintuplicate using an INSTRON EMIC 23-30 universal testing machine (Instron, Norwood, MA, USA) with tensile grips at 25 °C, applying a crosshead displacement rate of 10 mm/min. Stress–strain curves were obtained from the tests, allowing for calculation of the elastic modulus of the printed scaffolds.

2.5. Antimicrobial Coating

2.5.1. Scaffolds Coated with HPMC/PHMB

PLDLA–TMC + PEG scaffolds (ring-shaped, 5 mm outer diameter, 3 mm inner diameter, and 2 mm length; 5.5 ± 0.2 mg) were immersed in an aqueous solution containing 2.0 g of hydroxypropyl methylcellulose (HPMC) types K4M and K100M (3:1 ratio of Methocel–Colorcon, Cotia, Brazil) in 60 mL of deionized water, supplemented with PEG 400 (0.10 g, 5% w/w of HPMC mass-Sigma-Aldrich, St. Louis, MO, USA). The PHMB was obtained by lyophilizing an aqueous solution (Polihexam Líquido, Helianto, Cotia, Brazil). The resulting solid was incorporated at 3%, 6%, or 12% of the HPMC mass (0.06 g, 0.12 g, or 0.24 g), and a control group was prepared without PHMB. The chemical structures of the coating components are illustrated in Figure 2.

The scaffolds were immersed in the coating solution for 2 min at room temperature. This procedure was repeated once to ensure a uniform layer. Following the drying stage, the samples were sterilized via UV-C irradiation (254 nm) for a total of 2 h. To ensure uniform decontamination of the ring-shaped structures, the scaffolds were inverted at the midpoint of the process, allowing each face to be oriented toward the light source for 1 h.

2.5.2. In Vitro Release Study

PLDLA–TMC + PEG scaffolds coated with 3%, 6% or 12% of PHMB were immersed in 2 mL of phosphate-buffered saline (PBS, pH 7.4, Sigma-Aldrich, St. Louis, MO, USA) at 37 °C under orbital shaking at 75 rpm. Aliquots of 1 mL were collected at predetermined points: 5, 10, 15, 20, 25, and 30 min; 1 and 12 h; and 1, 2, 3, 5, and 7 days and an equal volume of fresh PBS was added to maintain a constant volume. PHMB concentration in the medium was quantified using Coomassie Brilliant Blue-based colorimetric reagent (Sigma-Aldrich, St. Louis, MO, USA), which interacts electrostatically and via hydrophobic interactions with the cationic PHMB, producing a measurable color change. A calibration curve was constructed using PHMB standards with the same reagent to enable direct quantification of the released antimicrobial. Release profiles were expressed as PHMB concentration (µg/mL) as a function of time. All experiments were performed in triplicate.

Release kinetics were evaluated by fitting the experimental concentration data to zero-order, first-order, Higuchi, and Korsmeyer–Peppas models using DDSolver (an Excel add-in; China Pharmaceutical University, Nanjing, China). For the Korsmeyer–Peppas model, the release exponent (n) was obtained directly from DDSolver nonlinear fitting.

2.5.3. Cell Viability

Endothelial cells (Code: 0345) obtained from the Rio de Janeiro Cell Bank (BCRJ) were seeded at a density of 1 × 10⁴ cells per well in 96-well plates and allowed to adhere overnight in Dulbecco’s modified Eagle medium (DMEM; Sigma-Aldrich, St. Louis, MO, USA; Cat. No. D6046) supplemented with 10% fetal bovine serum (CultiLab, Campinas, SP, Brazil) and 1% Penicillin-Streptomycin (Gibco, Carlsbad, CA, USA). For the subsequent indirect cytotoxicity assay (ISO 10993-5:2009) [13], scaffold extracts were first prepared in accordance with ISO 10993-12:2012 [14]. Coated scaffolds (5.5 mg each) from the 3%, 6%, and 12% PHMB groups, as well as the uncoated control, were individually incubated in 2 mL of the culture medium for 24 h at 37 °C. The resulting mass-to-volume ratio was 2.75 mg of scaffold per mL. The medium was filtered and used as the corresponding extract.

The standard culture medium was then replaced with 200 µL of the corresponding scaffold extract. Cells were incubated for an additional 24 h. After incubation, 200 µL of MTT solution (0.5 mg/mL) was added to each well and incubated for 1.5 h at 37 °C. Formazan crystals were solubilized with 200 µL of DMSO (Sigma-Aldrich, St. Louis, MO, USA), and absorbance was measured at 570 nm using a microplate reader (SpectraMax^® iD5, Molecular Devices, San Jose, CA, USA).

Cell viability was expressed as a percentage relative to the control group (cells without scaffold). Data were analyzed using GraphPad Prism 8.0 (GraphPad Software, San Diego, CA, USA). Results are presented as mean and standard deviation (SD) and were compared using one-way analysis of variance (ANOVA) followed by Tukey’s post hoc test for multiple comparisons. Differences were considered statistically significant at p < 0.05.

2.5.4. Antimicrobial Susceptibility Test

Scaffolds coated with 3% and 6% PHMB were tested against Staphylococcus aureus (ATCC 25923). For the agar diffusion assay, bacterial suspensions (1 × 10⁶ CFU/mL) were spread on Mueller–Hinton agar plates (Becton Dickinson-Difco, Franklin Lakes, NJ, USA). Scaffolds (ring-shaped, 5 mm outer diameter, 3 mm inner diameter, and 2 mm length) were placed on the agar surface and incubated at 37 °C for 24 h. Zones of inhibition were measured in millimeters. All tests were performed in triplicate.

2.6. In Vitro Vascular Model Development

3D-printed PLDLA–TMC + PEG scaffolds (0.5% and 2.0% PEG) were seeded with myoblasts (Code: 0141) and endothelial cells (Code: 0345) obtained from the Rio de Janeiro Cell Bank (BCRJ). Cells were expanded in 75 cm² culture flasks and maintained in DMEM (Sigma-Aldrich, St. Louis, MO, USA; Cat. No. D6046) supplemented with 10% fetal bovine serum (CultiLab, Campinas, SP, Brazil) and 1% Penicillin-Streptomycin (Gibco, Carlsbad, CA, USA). Scaffolds were sterilized by UV-C irradiation (254 nm, 1 h) and pre-incubated in DMEM at 37 °C for 24 h. Following 14 days of co-culture, constructs were fixed with 4% paraformaldehyde (Sigma-Aldrich, St. Louis, MO, USA) immunostained with CD31 (Thermo Fisher Scientific, Cat. No. 14-0311-82) for endothelial cells, followed by an Alexa Fluor 488-conjugated secondary antibody (excitation/emission: 488/519 nm), and α-smooth muscle actin (α-SMA; Thermo Fisher Scientific, Cat. No. 14-9760-82) for myoblasts, followed by an Alexa Fluor 647-conjugated secondary antibody (excitation/emission: 650/668 nm). Nuclear labeling used DAPI with FluroshieldTM antifade mounting medium. Samples were analyzed by laser scanning confocal microscopy (Zeiss LSM 710, Carl Zeiss, Jena, Germany), with Z-stack images acquired to a depth of up to 100 μm.

3. Results and Discussion

3.1. Physicochemical Properties of the PLDLA–TMC/PEG Polymers

3.1.1. Molecular Weight Distribution

The number-average Mn, Mw, and PDI values for the PLDLA–TMC, PLDLA–TMC + PEG 0.5%, and PLDLA–TMC + PEG 2.0% polymers are summarized in

Table 2. Values of number-average molecular weight (Mn), weight-average molecular weight (Mw), and polydispersity index (PDI) obtained by gel permeation chromatography (GPC).

Sample	Mn (g/mol)	Mw (g/mol)	PDI
PLDLA–TMC	60,943	130,506	2.14
PLDLA–TMC + PEG 0.5%	38,562	99,408	2.57
PLDLA–TMC + PEG 2.0%	42,838	140,691	3.28

PLDLA–TMC: poly(L-co-D,L-lactide)–poly(trimethylene carbonate); PEG: polyethylene glycol.

.

GPC-derived molecular weight values correspond to apparent values for the PLDLA–TMC/PEG blends, reflecting their hydrodynamic behavior in solution. For PLDLA–TMC, the Mn and Mw values are consistent with ring-opening polymerization-derived terpolymers and indicative of a moderately broad but well-controlled molecular weight distribution [15]. An average reduction of approximately 20% in the molecular weight of PLDLA–TMC was observed after PEG incorporation, particularly in the PLDLA–TMC + PEG 0.5% sample. This molecular weight reduction was accompanied by an increase in PDI to 2.57, suggesting a broadening of the molecular weight distribution as differences in hydrodynamic volume between components increase the apparent heterogeneity measured by GPC [16].

The reduction in Mn and Mw observed after PEG incorporation, together with the increase in PDI, is consistent with the heterogeneous nature of PLDLA–TMC/PEG blends. Since PEG is physically incorporated, with no evidence of covalent bonding to the PLDLA–TMC backbone, differences in hydrodynamic volume between blend components influence chromatographic elution behavior, resulting in broader apparent molecular weight distributions. Accordingly, the variations in Mn, Mw, and PDI can be attributed to blend heterogeneity and hydrodynamic effects rather than to PEG-induced degradation or chain scission of the PLDLA–TMC terpolymer [17].

Increasing the PEG content up to 2.0 wt% led to a further increase in PDI, indicating enhanced system heterogeneity, which has been associated in the literature with PEG-induced phase separation effects or preferential elution behavior of higher-molecular-weight fractions, as reported by Kulinski & Piorkowska [18].

From a materials perspective, the observed evolution in molecular weight distribution with increasing PEG content has important implications for downstream processing and performance. While the reduction in Mₙ at low PEG concentration may adversely affect mechanical strength, the higher Mw and broader distribution at 2.0 wt% PEG may enhance melt elasticity, processability, and degradation tunability.

3.1.2. FTIR Spectroscopy Analysis

The FTIR spectra confirmed the presence of the main functional groups of the PLDLA–TMC terpolymer before and after PEG incorporation (Figure 3). Regarding the characteristic bands of the poly(lactic acid) family, the following were identified: C–O stretching at 1081 cm⁻¹, C–O–C stretching at 1185 cm⁻¹, C=O stretching of the ester group at 1745 cm⁻¹, CH stretching at 2945 cm⁻¹, and CH₃ stretching at 2995 cm⁻¹. The bands associated with the presence of TMC were observed at 1745 cm⁻¹ (C=O stretching of the ester group), 1247 cm⁻¹ (asymmetric O–C–O stretching), and 790 cm⁻¹ (asymmetric O=C–O deformation) [19,20]. Upon PEG incorporation, additional absorption bands became evident in the spectra of the modified formulations. In particular, bands at 1452 cm⁻¹ and 874 cm⁻¹ were observed and attributed to PEG-related vibrational modes, including the –CH₂ bending and C–O–C skeletal vibrations [21]. The presence of these bands, combined with subtle changes in band intensity in the fingerprint region (1200–1000 cm⁻¹), indicates the successful incorporation of PEG into the polymeric system.

Notably, no new absorption bands or significant peak shifts were detected, suggesting that PEG is physically incorporated without chemically altering the PLDLA–TMC backbone, and that the subsequent changes in interfacial properties are associated with PEG-induced physical rearrangements and hydration effects.

3.1.3. DSC Thermal Analysis

The DSC thermograms obtained from the first and second heating cycles of PLDLA–TMC and PLDLA–TMC containing PEG at 0.5% and 2.0% revealed the presence of a single glass transition (Tg) and the absence of any melting or crystallization events, confirming the fully amorphous nature of the systems (Figure 4).

The Tg values were determined from the second heating cycle (

Table 3. Glass transition temperature (Tg) of PLDLA–TMC before and after PEG incorporation.

Sample	Tg (°C)
PLDLA–TMC	40.0
PLDLA–TMC + PEG 0.5%	37.0
PLDLA–TMC + PEG 2.0%	37.0

PLDLA–TMC: poly(L-co-D,L-lactide)–poly(trimethylene carbonate); PEG: polyethylene glycol.

). The Tg values of PLDLA–TMC, PLDLA–TMC + 0.5% PEG, and PLDLA–TMC + 2.0% PEG were 40 °C, 37 °C, and 37 °C, respectively. These results suggest a trend of decreasing Tg with PEG addition, while the presence of TMC units already provides a baseline reduction in chain rigidity, consistent with a plasticizing effect of PEG at the concentrations investigated.

Although PEG is widely recognized as an efficient plasticizer for poly(lactic acid)-based systems, typically reducing Tg by increasing free volume and enhancing chain mobility [22], its effectiveness is strongly dependent on molecular weight, concentration, miscibility, and the strength of intermolecular interactions with the polymer matrix. In the present case, the minimal change in Tg indicates that PEG is molecularly dispersed within the PLDLA–TMC matrix without significantly disrupting the dominant intermolecular interactions governing the amorphous phase. This behavior can be attributed to the relatively low PEG contents, which are insufficient to markedly increase segmental mobility or alter the cooperative motion of PLDLA–TMC chains. Moreover, the presence of TMC units, known to impart flexibility to the copolymer backbone, may already provide a baseline reduction in chain rigidity, thereby attenuating any additional plasticization effect introduced by PEG [23].

Importantly, the absence of PEG-related thermal transitions further supports the hypothesis of good miscibility and lack of PEG-rich domains, which would otherwise be expected to manifest as distinct melting or glass transition events. This homogeneous amorphous structure is particularly advantageous for applications involving thermal processing or 3D printing, as it ensures dimensional stability without compromising the thermal window of the material.

3.1.4. TGA Thermal Stability Analysis

The thermogravimetric and derivative thermogravimetric curves of the PLDLA–TMC and PEG-containing formulations are presented in Figure 5, while the corresponding onset degradation temperatures (T_onset) and temperatures of maximum mass loss (T_max) are summarized in

Table 4. Onset temperature of mass loss (T_onset) and temperature of maximum mass loss (T_max) for PLDLA–TMC before and after PEG incorporation.

Sample	Tonset (°C)	Tmax (°C)
PLDLA–TMC	281.4	296.0
PLDLA–TMC + PEG 0.5%	103.0/283.9	326.0
PLDLA–TMC + PEG 2.0%	113.0/278.1	318.0

PLDLA–TMC: poly(L-co-D,L-lactide)–poly(trimethylene carbonate); PEG: polyethylene glycol.

. Pure PLDLA–TMC exhibits a single, well-defined mass loss event, with T_onset at 281.4 °C and T_max at 296.0 °C, which is characteristic of a homogeneous degradation process dominated by random chain scission of ester bonds in the amorphous terpolymer backbone [24,25].

Upon PEG incorporation, the thermal degradation profile is modified, leading to the appearance of a two-step mass loss behavior. For both PEG-containing systems, an initial mass loss event is observed at lower temperatures (T_onset 103.0–113.0 °C), followed by the main degradation step associated with the PLDLA–TMC matrix. This early mass loss is attributed to the thermal decomposition of PEG segments, in agreement with the intrinsically lower thermal stability of PEG compared to aliphatic polyesters [26].

The addition of PEG does not significantly affect the T_onset of the samples compared to pure PLDLA–TMC, indicating that the initial thermal stability of the terpolymer backbone is maintained. However, T_max values shifted to higher temperatures relative to pure PLDLA–TMC (326 °C for 0.5% PEG and 318 °C for 2.0% PEG, compared to 296 °C for pure PLDLA–TMC), suggesting that PEG modifies the degradation kinetics, requiring higher thermal energy to reach the maximum mass loss rate. This effect is likely associated with intermolecular interactions between PLDLA–TMC and PEG, which may restrict chain mobility and delay the main degradation process. Supporting this, the reduced intensity of the DTG peaks for PEG-containing samples indicates a slower decomposition rate, consistent with the interaction between PLDLA–TMC and PEG [27,28]. From a processing standpoint, although PEG incorporation introduces an early mass loss event, the temperatures employed for processing remain below those associated with significant thermal degradation of the polymer matrix.

Taken together, DSC and TGA analyses indicate that at the investigated concentrations, PEG does not induce a classical plasticization of the PLDLA–TMC matrix, as evidenced by the minimal changes in Tg values while still modifying the thermal degradation kinetics by introducing an additional low-temperature mass loss event and subtly shifting the temperature of the maximum degradation rate without compromising the intrinsic thermal stability of the polymer backbone.

3.1.5. Rheological Behavior

The PLDLA–TMC + PEG 0.5% and PLDLA–TMC + PEG 2.0% samples were subjected to oscillatory rheological tests to investigate their structural properties, particularly viscoelasticity. Initially, an amplitude sweep was performed to identify the linear viscoelastic region (LVR) of the materials. Determining the LVR is essential for studying the polymer’s structural properties, as within this region, stress and strain are proportional, meaning that the applied stress is insufficient to cause structural breakdown of the material [29]. Consequently, the results obtained can be directly correlated to the structural properties of the polymer under study. For both materials, a strain of 0.2% was selected, and based on these values, frequency sweep tests were subsequently performed (Figure 6).

As reported in the literature for the neat PLDLA–TMC terpolymer (80:20 composition), a predominantly viscous response is typically observed in the melt state, characterized by a loss modulus (G″) higher than the storage modulus (G′) across the 0.1 rad. s⁻¹ frequency range [29]. In the present study, both moduli increased with rising frequency, indicating a clear frequency-dependent relationship. The PLDLA–TMC + PEG 0.5% sample followed the trend of the neat polymer, exhibiting predominantly viscous behavior with G″ consistently higher than G′. In contrast, the PLDLA–TMC + PEG 2.0% sample displayed a transition toward predominantly elastic behavior, as evidenced by the crossover of the storage and loss moduli at higher frequencies. This shift in viscoelastic properties is directly related to the varying PEG concentrations. The results indicate that PEG incorporation at 2.0% increased the material’s structural resistance, as reflected by the modulus crossover, likely due to enhanced chain entanglement under higher frequency conditions. From the perspective of vascular model applications, this combination of increased resistance and flexibility is highly relevant, as such devices require these properties for successful implantation [30].

Based on the viscoelastic characterization of both materials, steady-shear rheological tests were conducted to evaluate their viscosity behavior. Viscosity measures a material’s resistance to a given type of flow [30]. The shear rate-dependent viscosity, η(γ), is a critical rheological property in polymer processing for 3D printing, as processing parameters such as temperature, pressure, and printing speed directly depend on it [31]. The viscosity curves obtained from the steady-shear tests for PLDLA–TMC + PEG 0.5% and PLDLA–TMC + PEG 2.0% are shown in Figure 7.

Overall, the results indicate that both materials exhibited pseudoplastic behavior. Pseudoplasticity is common in molten polymers and arises from the unfolding and alignment of macromolecules under applied shear [32]. At low shear rates, the velocity gradient is insufficient to unfold the macromolecules, and the viscosity remains high. As the shear rate increases, macromolecules gain mobility, disentangle, and viscosity decreases. At high shear rates, most entanglements are undone, and the macromolecules align in the flow direction, leading to a further reduction in viscosity [33].

However, distinct differences between the two materials were observed. The PLDLA–TMC + PEG 2.0% sample exhibited higher viscosity than PLDLA–TMC + PEG 0.5%. Moreover, increased material resistance was evident for PLDLA–TMC + PEG 2.0%, as the steady-shear test could not be completed across the full shear rate range due to the higher resistance encountered during the measurement. Considering that the steady-shear tests were conducted at 80 °C, these results indicate that the 3D printing of PLDLA–TMC + PEG 2.0% would require temperatures above 80 °C to ensure sufficient viscosity reduction at the shear rates relevant for printing (10²–10³ s⁻¹).

3.1.6. Surface Charge

The zeta potential (ζ) of PLDLA–TMC formulations containing PEG at 0.5% and 2.0% was evaluated as a function of pH, and the results are presented in Figure 8. For comparison, the processed PLDLA–TMC formulation without PEG exhibited a zeta potential of approximately −6.3 mV at near-physiological pH (7.4) [34], indicating a low-magnitude negative surface charge when compared with PEG-containing formulations, which showed more negative values of −31 mV for PLDLA–TMC/PEG 0.5% and −37 mV for PLDLA–TMC/PEG 2.0%.

Both formulations exhibited negative ζ values over the entire pH range investigated (pH 3–10), with increasingly negative potentials at higher pH values, reflecting the progressive deprotonation of ionizable surface groups typical of polyester-based systems [35]. At acidic pH values, the ζ potential assumes moderately negative values (approximately −15 to −20 mV), indicating partial protonation of surface groups and limited surface charge density. As the pH increases, deprotonation becomes more pronounced, leading to a substantial increase in negative surface charge and a corresponding decrease in ζ potential, reaching values below −35 mV for PEG 0.5% and approximately −42 mV for PEG 2.0% at alkaline pH. This progressive shift highlights the strong pH sensitivity of the surface electrokinetic behavior of the formulations.

A clear PEG concentration-dependent effect was observed, with the 2.0% PEG formulation consistently exhibiting more negative ζ potential values than the 0.5% PEG formulation. Although PEG is non-ionic, its presence can promote surface hydration and polymer chain rearrangement at the interface, leading to increased exposure of negatively charged groups and altered electrokinetic response [36].

The absence of an isoelectric point within the investigated pH range indicates that the scaffold surface remains predominantly negatively charged under physiologically relevant conditions. In the context of a solid scaffold, this strong surface charge likely reflects a high density of active sites and the formation of a stable hydration layer at the polymer–liquid interface, which can influence protein adsorption and cell–material interactions [37].

From a biological perspective, surface charge plays a critical role in protein adsorption, cell–material interactions, and in vivo behavior. Moderately strongly negative surfaces are often associated with reduced nonspecific protein adsorption and attenuated inflammatory responses, which may be beneficial for implantable or injectable biomaterials. However, other studies with PLDLA–TMC have also shown that surface negativity does not inhibit cell growth [30,38].

FTIR analysis confirmed PEG incorporation into the PLDLA–TMC system without chemical modification of the polymer backbone. Consistently, zeta potential measurements showed a PEG content-dependent shift toward more negative values, indicating modification of the surface electrokinetic properties. Together, these results support the presence of PEG at or near the scaffold interface, associated with surface hydration and polymer chain rearrangement.

3.2. Mechanical Performance of Printed Scaffolds

The mechanical properties of polymeric materials are of great importance due to the wide range of applications that require specific mechanical performance. Parameters such as the elastic modulus can serve as a basis for comparing the mechanical behavior of different materials [39]. Therefore, to evaluate the mechanical properties of the 3D-printed PLDLA–TMC + PEG 0.5% and PLDLA–TMC + PEG 2.0% scaffolds, tensile tests were performed in accordance with the ASTM D638 standard [12], as illustrated in Figure 9.

For the PLDLA–TMC + PEG 0.5% samples, the initial region of the stress–strain curves exhibited elastic behavior, with an almost linear increase in stress with strain, reflecting the elastic resistance of the samples. The maximum stress reached was low (0.15 ± 0.10 MPa), suggesting that the material containing 0.5% PEG has relatively limited strength. After the peak stress, the material showed a decline in resistance, corresponding to the yielding and subsequent softening regions. This behavior indicates that the material loses part of its structural integrity after reaching maximum stress, resulting in plastic deformation or embrittlement.

For PLDLA–TMC + PEG 2.0%, the initial region was similar, with a linear increase in stress with strain. However, the peak stress was higher (0.79 ± 0.12 MPa). This increase in mechanical strength indicates that a higher PEG content improved the tensile resistance of the material. After the peak, the stress decline was more gradual compared to the PLDLA–TMC + PEG 0.5% curves. Thus, PLDLA–TMC with 2.0% PEG exhibited substantially higher mechanical resistance, with peak stresses approximately four times greater than those of the 0.5% PEG scaffolds. Additionally, the 2.0% PEG scaffolds showed higher deformation capacity and, consequently, greater ductility. In contrast, the 0.5% PEG scaffolds experienced a sharper stress drop after the peak, suggesting faster embrittlement or yielding, whereas the 2.0% PEG material maintained more controlled deformation.

Figure 10 presents the Young’s modulus as a function of PEG content. The scaffold containing 2.0% PEG exhibited the highest Young’s modulus, indicating increased stiffness relative to the formulation with 0.5% PEG. This response is consistent with enhanced intermolecular interactions between the PLDLA–TMC matrix and PEG at 2.0% incorporation, which likely contributed to more efficient stress transfer within the polymer network when compared with the lower PEG content.

Overall, scaffolds containing 2.0% PEG demonstrated a more balanced combination of strength and deformability, which is crucial for applications requiring both mechanical resistance and flexibility, without premature fracture, such as in vascular grafts.

3.3. Antimicrobial Properties of the Coated Scaffolds

3.3.1. PHMB Release Kinetics

The PHMB release kinetics from HPMC-coated PLDLA–TMC scaffolds containing 2% PEG were dependent on the drug load (Figure 11). The PLDLA–TMC (80/20) scaffolds with 2% PEG in the polymer matrix and coated with an HPMC layer containing 3% PHMB (PEG2/PHMB3) reached a release plateau within 1 h (310.59 ± 15.74 µg/mL; 94.1% of the incorporated mass), indicating a burst-dominated release. Scaffolds coated with HPMC containing 6% PHMB (PEG2/PHMB6) exhibited a gradual release and reached a plateau at 7 days (658.74 ± 55.83 µg/mL; 99.81%). Scaffolds coated with HPMC containing 12% PHMB (PEG2/PHMB12) showed sustained release over 7 days, reaching 1299.41 ± 61.83 µg/mL (98.44%), consistent with near-complete release.

Higher PHMB loading appears to promote more effective retention within the HPMC matrix due to the formation of a denser and more viscous hydrated gel layer, which slows diffusion and shifts the release profile toward anomalous transport. This behavior aligns with reports on hydrophilic HPMC-based delivery systems, where variations in the drug-to-polymer ratio modulate both release rate and dominant transport mechanisms. Similar effects have been observed in HPMC K100M microspheres loaded with fexofenadine, in which higher drug content led to sustained release associated with a thicker hydrated barrier layer [40].

PHMB release profiles were fitted to classical kinetic models to elucidate predominant transport mechanisms (

Table 5. Correlation coefficients (R²) obtained from fitting the release profiles to kinetic models over the entire experimental period (0–168 h).

Sample	Zero Order	First Order	Higuchi	Korsmeyer– Peppas
PEG2/PHMB3	0.32	0.55	0.72	0.81
PEG2/PHMB6	0.58	0.82	0.89	0.94
PEG2/PHMB12	0.72	0.88	0.97	0.99

Samples consist of PLDLA–TMC (80/20)/2% PEG 400 scaffolds coated with HPMC containing 3%, 6%, or 12% PHMB, respectively.

). For PEG2/PHMB3, global fitting over 0–168 h resulted in limited correlations (R² = 0.72 for Higuchi, 0.81 for Korsmeyer–Peppas) due to the rapid establishment of the plateau. Analysis restricted to the initial 0–1 h interval showed a strong correlation with the Higuchi model (R² = 0.985), indicating that early release is predominantly Fickian diffusion.

PEG2/PHMB6 displayed a better overall fit to the Korsmeyer–Peppas model (R² = 0.94, n = 0.64), reflecting anomalous transport governed by combined diffusion and polymer relaxation. For PEG2/PHMB12, the Korsmeyer–Peppas model yielded R² = 0.99 and N = 0.71, indicating a greater contribution of HPMC matrix erosion and relaxation to the release rate. Early-stage kinetic analysis (0–1 h) confirmed Fickian-dominated diffusion for all formulations, with N values close to 0.5 (

Table 6. Correlation coefficients (R²) obtained from fitting the release profiles to kinetic models over the entire experimental period (0–1 h).

Sample	Zero Order	Higuchi	Korsmeyer– Peppas
PEG2/PHMB3	0.916	0.985	0.989
PEG2/PHMB6	0.875	0.981	0.983
PEG2/PHMB12	0.931	0.996	0.997

Samples consist of PLDLA–TMC (80/20)/2% PEG 400 scaffolds coated with HPMC containing 3%, 6%, or 12% PHMB, respectively.

).

These results demonstrate a transition from predominantly diffusive release in the 3% formulation to anomalous transport in higher PHMB-loaded scaffolds, highlighting the role of the HPMC coating in modulating antimicrobial release through the formation of a hydrated gel barrier following initial diffusion, as also observed by Felipin et al. [41].

The sustained release of PHMB from PEG2/PHMB6 and PEG2/PHMB12 scaffolds over 7 days maintained elevated drug concentrations within the matrix, supporting prolonged antimicrobial activity compared with the rapid burst observed in PEG2/PHMB3.

3.3.2. Cytotoxicity of Scaffold Extracts

The indirect cytotoxicity of extracts from PLDLA–TMC scaffolds with 2.0% PEG, coated with HPMC containing different PHMB proportions, was evaluated using the MTT assay. The cell viability results are presented in Figure 12.

The PEG2 group, corresponding to extracts from scaffolds with a PLDLA–TMC core with 2.0% PEG, coated with HPMC without PHMB (PEG2), exhibited a cell viability significantly higher than the viable control (122.50 ± 1.80%; p < 0.001), indicating the absence of cytotoxic effects associated with the polymeric matrix.

The incorporation of 3% PHMB into the coating of PLDLA–TMC/PEG 2.0% scaffolds (PEG2/PHMB3) resulted in a reduction in cell viability to 95.68 ± 1.94%, remaining above the cytocompatibility threshold defined by ISO 10993-5. Increasing the PHMB concentration to 6% (PEG2/PHMB6) led to a further decrease in viability, reaching 72.95 ± 1.10%, a value close to the lower limit recommended by the standard.

In contrast, extracts from the PLDLA–TMC/PEG 2.0% scaffolds coated with 12% PHMB (PEG2/PHMB12) showed residual cell viability (7.70 ± 0.37%), statistically indistinguishable from the unviable control (p = 0.271), indicating a pronounced cytotoxic effect. Statistical analysis confirmed a PHMB concentration-dependent response, with significant differences among all evaluated groups (p < 0.001), except between PEG2/PHMB12 and the unviable control.

Cell viability results align with the current literature on PHMB in biomedical matrices, which reports a concentration-dependent cytotoxicity profile. In chitosan/PEG hybrid hydrogels, the incorporation of 3% PHMB slightly reduced fibroblast viability compared to PHMB-free hydrogels (~94% vs. ~100%), while higher concentrations, such as 10%, decreased viability to ~45% after 24 h, indicating a clear dose-dependent effect [42]. Consistently, our findings suggest that PHMB concentrations above 6% compromise cytocompatibility and should be avoided in vitro models.

3.3.3. Antimicrobial Activity Against Staphylococcus aureus

PHMB exhibits strong antimicrobial activity against staphylococcal species, including Staphylococcus aureus and Staphylococcus epidermidis, with minimum inhibitory concentration (MIC) values reported in the literature in the low microgram per milliliter range (0.5–1 µg·mL⁻¹), as determined by broth microdilution assays [43,44]. These values provide a reference for interpreting the antimicrobial activity observed in the agar diffusion assays. Accordingly, in the present study, antimicrobial activity is discussed in terms of the effectiveness of PHMB when associated with the developed scaffold system.

The positive control, gentamicin, exhibited the highest antimicrobial activity against Staphylococcus aureus, with an average inhibition diameter of 22.0 ± 0.3 mm. Both PHMB-containing formulations demonstrated measurable antimicrobial effects. The scaffold coated with 3% PHMB (PEG2/PHMB3) produced an inhibition zone of 12.5 ± 0.1 mm, while the 6% PHMB scaffold (PEG2/PHMB6) generated a zone of 13.1 ± 0.2 mm. The PLDLA–TMC/PEG scaffold without PHMB (PEG2) showed no inhibition zone. Representative images of the agar diffusion assay are shown in Figure 13.

The agar diffusion assay performed over 24 h confirmed the effective antimicrobial activity of PHMB-coated scaffolds. Under these conditions, both 3% and 6% PHMB formulations produced clear inhibition zones, demonstrating that the amount of PHMB released during the initial phase is sufficient to inhibit Staphylococcus aureus growth. The similar inhibition diameters observed for both concentrations indicate that within the first 24 h, antimicrobial efficacy is already achieved, while differences in long-term release profiles become relevant at later stages. When compared with the literature-reported MIC values (≈0.5–1 µg·mL⁻¹), both 3% and 6% PHMB formulations maintained PHMB levels above the MIC, whereas the PEG2/PHMB6 system exhibited a more sustained release profile over time. The antimicrobial results are consistent with studies on PHMB-containing biomaterials. In porous PHMB/silk fibroin scaffolds, PHMB/SF ratios above 2/100 produced clear inhibition of Staphylococcus aureus, while lower concentrations showed no antibacterial effect [45].

When the microbiological and cytocompatibility results are analyzed together, a clear therapeutic window for PHMB incorporation into the scaffold system can be identified. The agar diffusion assay demonstrated that both 3% and 6% PHMB-loaded scaffolds produced comparable inhibition zones against S. aureus, indicating that effective antimicrobial activity is already achieved at lower PHMB concentrations, likely due to the high intrinsic potency of PHMB and its initial release during the early exposure phase. In contrast, cytocompatibility assays revealed a concentration-dependent decrease in cell viability at higher PHMB loadings. Taken together, the findings here indicate that increasing the PHMB content beyond the minimum effective antimicrobial concentration does not provide additional microbiological benefit, while significantly compromising cellular viability. Therefore, PHMB concentrations within the lower tested range represent an optimal balance between antimicrobial efficacy and cytocompatibility, supporting their suitability for scaffold-based biomedical applications.

3.4. In Vitro Vascular Model Assessment

Following material characterization, scaffolds containing 0.5% and 2.0% PEG were seeded with cells in a spatially organized and biomimetic manner. Endothelial cells were seeded onto the inner lumen of the scaffolds to emulate the tunica intima, while myoblasts were cultured on the outer region to mimic the cellular organization of the tunica media of native blood vessels (Figure 14B). After 14 days of culture, confocal microscopy analyses were performed to assess cell distribution, coverage, and spatial organization within both the luminal and outer regions of the scaffolds (Figure 14A).

The confocal micrographs revealed strong α-SMA staining in the scaffold regions seeded with myoblasts for both PLDLA–TMC + PEG 0.5% and PLDLA–TMC + PEG 2.0% scaffolds (Figure 14A). α-SMA is a widely used marker for identifying cells with a contractile phenotype, such as myoblasts, indicating differentiation and potential for muscle tissue formation [46]. In contrast, CD31 staining was less pronounced in the regions containing myoblasts (Figure 14A). CD31, also known as platelet endothelial cell adhesion molecule-1 (PECAM-1), is a well-established marker of endothelial cells and vascular structures [47]. The low CD31 expression in these regions suggests limited interaction between myoblasts and endothelial cells, with both cell types remaining predominantly within their initially seeded compartments.

In the inner cavity of the scaffolds, where endothelial cells were seeded, CD31 expression was more prominent in PLDLA–TMC + PEG 0.5% scaffolds compared to PLDLA–TMC + PEG 2.0% scaffolds (Figure 14A). The enhanced CD31 signal observed in scaffolds with lower PEG content may be associated with a more favorable microenvironment for endothelial cell adhesion, proliferation, and organization, potentially influenced by differences in microporosity or cell–material interactions [48]. Conversely, α-SMA staining was virtually absent in the regions seeded with endothelial cells (Figure 14A), as expected, since endothelial cells do not exhibit the contractile characteristics typical of smooth muscle or myogenic cells.

From a vascular tissue engineering perspective, these findings indicate that both PLDLA–TMC + PEG 0.5% and PLDLA–TMC + PEG 2.0% scaffolds effectively support cell adhesion, growth, and lineage-specific marker expression. Despite some degree of cellular proximity throughout the scaffold, cells largely remained within their originally designated regions, as schematically illustrated in Figure 14B. Such spatial organization is desirable for in vitro vascular models, as it supports biomimetic tissue architecture and the maintenance of cell-specific phenotypic functions. Regarding potential applications in vascular tissue engineering, the results suggest that both scaffold types represent promising candidates, highlighting the importance of developing innovative approaches to overcome current model limitations and advance therapies for vascular diseases.

Building on the mechanical, structural, antimicrobial, and in vitro biocompatibility performance demonstrated here, future research will focus on validating the hemocompatibility, degradation profiles, and sterilization effects of these scaffolds to further ensure their clinical safety and efficacy.

4. Conclusions

This study successfully developed and validated an in vitro vascular model based on 3D-printed PLDLA–TMC scaffolds modified with polyethylene glycol and functionalized with a PHMB-based antibacterial coating. The incorporation of PEG proved to be an effective strategy to modulate the physicochemical, rheological, and mechanical properties of the polymeric system without compromising its thermal stability or processability by extrusion-based 3D printing. In particular, the formulation containing 2.0% PEG exhibited a marked improvement in tensile strength, elastic behavior, and deformation capacity, yielding mechanical properties more compatible with the requirements of vascular applications.

The application of an HPMC/PHMB coating enabled controlled antimicrobial release while preserving scaffold integrity. PHMB loading resulted in a clear concentration-dependent biological response: high PHMB content (12%) induced pronounced cytotoxicity, whereas lower concentrations maintained acceptable cytocompatibility. Although both 3% and 6% PHMB exhibited antibacterial activity against Staphylococcus aureus, the 3% formulation was characterized by a rapid burst release, leading to a transient antimicrobial effect. In contrast, the 6% PHMB coating provided a more sustained release profile, ensuring effective bacterial inhibition while preserving cytocompatibility, thereby representing a more suitable concentration for antimicrobial scaffold functionalization.

The in vitro vascular model further demonstrated the ability of the scaffolds to support spatially organized co-culture of endothelial cells and myoblasts, reproducing key aspects of native vascular architecture. Confocal analyses confirmed lineage-specific marker expression and preferential cellular localization, with endothelial cells lining the luminal surface and myoblasts occupying the outer region of the scaffold. This organized cellular distribution highlights the suitability of the developed system for mimicking vascular tissue structure and function under in vitro conditions.

Taken together, these findings demonstrate that PLDLA–TMC scaffolds containing 2.0% PEG and coated with 6% PHMB provide an optimized balance between mechanical performance, sustained antimicrobial functionality, and biocompatibility. The integrated strategy presented herein offers a robust platform for the development of advanced in vitro vascular models and establishes a rational framework for future studies in vascular tissue engineering and translational research.