Comparative In Vitro Degradation of Polyglycolic Acid (PGA) and Polyglactin 910 (PGLA) Sutures Under Hyperglycemic Conditions: An Integrated Mechanical, Chemical, and Morphological Study

1. Introduction

Surgical wound closure represents a fundamental step in ensuring the success of any invasive procedure, with sutures serving as the primary instruments for approximating and maintaining tissue edges until adequate wound healing occurs. Appropriate suture selection constitutes a critical determinant that directly influences tissue repair outcomes, infection rates, and the incidence of wound dehiscence [1,2]. The historical evolution of absorbable sutures, from natural materials such as catgut to modern synthetic polymers, has been comprehensively reviewed [2]. Synthetic absorbable sutures have gained substantial prominence in recent decades by offering the inherent advantage of eliminating the need for removal, thereby reducing patient discomfort and mitigating costs associated with secondary procedures. Among these materials, aliphatic polyesters—particularly polyglycolic acid (PGA) and its copolymers with lactic acid (PGLA)—represent the most extensively utilized class globally, owing to their predictable degradation profiles and biocompatibility [3]. Diabetes mellitus constitutes a highly prevalent metabolic disorder with global impact, affecting more than 537 million adults worldwide [4]. Diabetic patients undergoing surgical interventions demonstrate an elevated risk of postoperative complications, including surgical site infections and impaired wound healing, largely attributable to microvascular dysfunction, neuropathy, and altered immune responses [4,19]. The local hyperglycemic environment may directly interfere with suture materials by accelerating hydrolytic degradation through mechanisms involving increased water activity and altered polymer-solvent interactions [5,6].

Although previous investigations suggest that hyperglycemia may accelerate the degradation of absorbable polymers [7,8], the specific influence of elevated glucose concentrations on PGA compared to PGLA 910—particularly when correlating mechanical, chemical, and structural parameters—remains incompletely elucidated in the literature. This integrated assessment of multiple degradation markers is essential for developing a comprehensive understanding of the degradation process and for providing evidence-based guidance for clinical practice. The impact of hyperglycemic environments on biomaterial performance, particularly concerning vascular and tissue responses, has been increasingly investigated [15,16]; however, comparative data on the integrated mechanical, chemical, and morphological degradation of PGA and PGLA 910 sutures remain scarce. Without a comprehensive understanding of how hyperglycemia affects the mechanical, chemical, and morphological integrity of these sutures, surgeons lack evidence-based guidance for material selection in diabetic patients—potentially leading to premature suture failure, wound dehiscence, and increased morbidity.

To address this knowledge gap, the primary aim of this study was to comparatively evaluate the in vitro degradation profile of PGA homopolymer and PGLA 910 copolymer absorbable sutures under varying glucose concentrations (0, 300, and 600 mg/dL) over a 28-day period. Specifically, we analyzed the mechanical properties (maximum load, tensile stress, strain at maximum load) over time under different glucose concentrations, while simultaneously monitoring medium acidification as an indirect indicator of hydrolytic degradation and autocatalytic activity. Additionally, we aimed to investigate surface morphological changes using SEM and correlate these structural alterations with mechanical and pH measurements. Collectively, these integrated analyses sought to determine the influence of glucose concentration and exposure time on the degradation kinetics of both materials, providing a comprehensive understanding of their performance under hyperglycemic conditions.

2. Materials and Methods

2.1. Study Design

This investigation employed an experimental, comparative, controlled in vitro design with complete randomization, incorporating the following factors: "material type" (two levels: PGA and PGLA 910), "glucose concentration" (three levels: 0, 300, and 600 mg/dL), and "degradation time" (five levels: 0, 7, 14, 21, and 28 days). As this research does not involve human subjects or vertebrate animals, it is exempt from submission and review by institutional ethics committees, in accordance with Resolution No. 510/2016 of the Brazilian National Health Council (CNS).

2.2. Materials

A total of 120 synthetic absorbable sutures were used, equally divided between two groups (60 sutures per material):

- Group 1 (PGA Homopolymer): Polyglycolic acid suture, monofilament, coated, US 4-0 diameter (0.15–0.19 mm) (Bioline Fios Cirúrgicos Ltda, Anápolis, GO, Brazil).

- Group 2 (PGLA 910 Copolymer): Poly(lactic-co-glycolic acid) 90:10 suture, monofilament, coated, US 4-0 diameter (0.15–0.19 mm) (Bioline Fios Cirúrgicos Ltda, Anápolis, GO, Brazil).

All sutures were obtained from a single lot from a commercial manufacturer to ensure initial standardization. All sutures had an initial standardized length of 15.0 cm.

2.3. Experimental Groups

Sample Preparation: Each suture was weighed on an analytical balance and its initial diameter measured with a digital micrometer at three distinct points. All sutures were supplied in sterile packaging by the manufacturer. To ensure an additional margin of sterility for the in vitro experiment, the samples were handled under sterile conditions and incubated in sterile vials with sterile media at 37 ± 1°C. No additional physical or chemical sterilization method was applied to the sutures beyond the manufacturer's original sterile packaging.

Surgical Knot Standardization: To standardize the knot-tying procedure and ensure reproducibility of the applied tension, each suture was manually tied into a standardized surgeon's knot (single throw, square configuration) around a cylindrical metal template with an outer diameter of 0.45 mm (450 µm). The template served as a reference to ensure consistent tension during knot tying. After the knot was secured, the suture was carefully removed from the template, leaving a standardized loop with a surgeon's knot. This procedure mimicked clinical suturing conditions and ensured that all samples received a consistent pattern of initial mechanical stress prior to immersion in the incubation media.

In Vitro Incubation: Sutures were placed individually in sterile 10 mL glass vials containing 5 mL of the respective incubation medium (PBS 0.01 M, pH 7.4, with glucose concentrations of 0, 300, or 600 mg/dL). Vials were sealed and incubated in a bacteriological oven at 37 ± 1°C for up to 28 days..PBS at pH 7.4 was chosen as The incubation medium was renewed weekly (at days 7, 14, and 21) to maintain consistent glucose concentrations and pH conditions throughout the 28-day observation period. pH measurements were performed immediately before each medium renewal. [15]. The glucose concentrations of 300 mg/dL (≈16.7 mmol/L) and 600 mg/dL (≈33.3 mmol/L) were selected to represent poorly controlled diabetes and extreme hyperglycemia (within the range observed in hyperosmolar hyperglycemic state), respectively. While 600 mg/dL exceeds typical routine glycemic exposure, it was included to establish a worst-case scenario for assessing the maximum potential effect of hyperglycemia on suture degradation [5,7]. The complete experimental design, detailing the combination of material type, glucose concentration, and degradation time, is presented in Table S1 (Supplementary Materials).

Mechanical Testing: At times 0, 7, 14, 21, and 28 days, samples were subjected to tensile testing on a Shimadzu Autograph AGS-X universal testing machine equipped with a 10 kN load cell and rubber-coated grips. The sutures were mounted with the knot positioned centrally between the grips, and a preload of 0.1 N was applied to ensure alignment. The initial gauge length was 25 mm. Each specimen was pulled to failure at a constant crosshead speed of 50 mm/min, a widely accepted standard for tensile testing of surgical sutures [20] that ensures consistent and reproducible failure of the specimens. The following parameters were recorded: Maximum Load (N), Maximum Stress (MPa), Strain at Maximum Load (%). Force–displacement curves were used to calculate tensile properties. Four specimens were tested per condition (n = 4) at each time point. Mass loss was not measured in this study, as the presence of the knot introduces geometric variability and water uptake (swelling) can confound mass loss measurements, particularly in the early stages. Our integrated mechanical-chemical-morphological approach was prioritized to directly assess functional integrity, acidic byproduct release, and surface alterations.

pH Monitoring: The pH of the incubation medium was measured weekly before medium renewal using a calibrated digital pH meter (Quimis, model Q400AS).

Scanning Electron Microscopy (SEM): Representative samples were gently dried with absorbent paper, gold-sputtered using a gold evaporator, and examined on a Scanning Electron Microscope (SEM) available at the Multiuser Laboratory (LABMU) of the State University of Ponta Grossa (UEPG) at an accelerating voltage of 5 kV and a working distance of 10 mm. Surface degradation was evaluated qualitatively using the morphological classification system established by Chu and Campbell [16,27].

2.4. Statistical Analysis

Descriptive analysis with mean and standard deviation. Normality was verified by the Shapiro-Wilk test, and homogeneity of variances by Levene's test. Multiple comparisons were performed using two-way ANOVA with Tukey's post-hoc test, complemented by the non-parametric Kruskal-Wallis test. Correlations were assessed using Spearman's correlation coefficient (ρ). The significance level was set at α = 0.05. All statistical analyses were performed using R version 4.6.0 (tidyverse, rstatix, car, ggplot2). The sample size of n = 4 per condition was selected based on previous in vitro degradation studies of absorbable sutures [11,13], which demonstrated adequate statistical power to detect meaningful differences in mechanical properties. The study is exploratory in nature and provides preliminary data for future investigations with larger sample sizes.

3. Results

3.1. Mechanical Properties

3.1.1. Maximum Load (N)

The maximum load results are presented in Figure 1. At baseline (T0), PGLA 910 demonstrated significantly higher mechanical strength than PGA (14.56 ± 1.11 N vs. 13.37 ± 1.00 N; p < 0.001, Tukey's post-hoc). Both materials maintained mechanical strength during the first 7 days, with no significant changes relative to T0 (T7 vs T0: p = 0.080, Tukey's post-hoc), indicating that neither material undergoes significant hydrolytic degradation within the first week under the experimental conditions employed. However, a progressive loss of mechanical strength occurred from day 14 onward (T14 vs T0: p < 0.001 for both materials). A dramatic and near-complete loss of mechanical strength occurred between day 21 and day 28 for both materials (T21 vs T0: p < 0.001; T28 vs T0: p < 0.001). At T28, the residual maximum load for PGA ranged from 1.06 N to 1.16 N across all glucose concentrations, corresponding to a strength loss of 91.3–92.1% relative to T0. For PGLA 910, the residual maximum load ranged from 3.23 N to 3.81 N, corresponding to a strength loss of 73.8–77.8% relative to T0. PGLA 910 demonstrated approximately 3.4-fold higher residual strength than PGA at T28 (3.81 N vs. 1.12 N at 0 mg/dL; p < 0.001, Tukey's post-hoc), indicating superior resistance to hydrolytic degradation.

3.1.2. Percentage Strength Loss

Percentage strength loss relative to T0 (Figure 2) confirms the progressive degradation pattern: minimal loss during the first 7 days (−0.6% to +5.0%), followed by gradual loss from day 14 onward (10.4–26.4%), culminating in near-complete degradation between day 21 and day 28 (40.5–92.1%). At T28, PGA retained only 8.4% of its initial strength, while PGLA 910 retained 26.2% of its initial strength, confirming the superior resistance of the copolymer to hydrolytic degradation. Two-way ANOVA confirmed that the effect of material type on percentage strength loss was highly significant (p < 0.001), while glucose concentration did not significantly affect degradation kinetics (p = 0.153).

3.1.3. Strain at Maximum Load and Elastic Modulus

Strain at maximum load values are presented in Table S3 (Supplementary Materials). Strain values remained relatively stable through T14 but showed marked changes at T21 and T28 due to the progressive degradation of the materials. At baseline (T0), PGA exhibited higher strain (32.28 ± 11.80%) compared to PGLA 910 (27.44 ± 11.30%), although this difference was not statistically significant (p = 0.312, Tukey's post-hoc). Both materials showed progressive reduction in strain from T14 onward. At T28, PGA strain decreased dramatically to 1.53–4.86%, while PGLA 910 strain ranged from 8.60% to 15.21%. The high strain values observed at later time points are consistent with viscoelastic failure following extreme molecular-weight reduction, rather than ductile plastic deformation of the intact polymer. High variability was observed for PGLA 910 at T28 under 300 mg/dL glucose (15.21 ± 8.78%), primarily due to one sample exhibiting higher strain (27.77%) compared to the other three (9.02–14.77%). This heterogeneity likely reflects the advanced degradation state of the material at 28 days, where individual samples exhibit differential failure patterns. Exclusion of the outlier would yield 11.02 ± 3.34%. Table S3 (Supplementary Materials) presents the strain at maximum load results. Strain values remained relatively stable through T14 but showed marked changes at T21 and T28 due to the progressive degradation of the materials.

3.2. Medium pH

pH monitoring (Figure S1, Supplementary Materials) showed progressive acidification for both materials, consistent with the release of acidic degradation products. PGA promoted more pronounced acidification (pH 6.66–6.69 at T28), while PGLA 910 resulted in slightly higher pH values (6.71–6.77). For both materials, higher glucose concentrations were associated with marginally lower pH values, a trend which, although not statistically significant for mechanical properties (p = 0.153, two-way ANOVA), is consistent with a slight increase in hydrolytic activity. This suggests that the effect of glucose on acidification is subtle and may not reach the threshold required to produce a detectable change in mechanical strength. Figure 3 illustrates the correlation between residual maximum load and medium pH during in vitro degradation. Spearman's correlation demonstrated a very strong positive correlation between residual maximum load and medium pH for both materials: PGA: ρ = 0.89 (95% CI: 0.82–0.95; p < 0.001); PGLA 910: ρ = 0.86 (95% CI: 0.85–0.96; p < 0.001); This strong correlation validates the autocatalytic nature of the hydrolytic degradation mechanism, wherein acidic byproducts further catalyze ester bond cleavage, leading to accelerated polymer chain scission and loss of mechanical integrity.

3.3. SEM Morphological Analysis

SEM image analysis (Figure 4, Figure 5 and Figure 6) provided detailed visualization of the degradation process at the microstructural level. The most significant finding was the marked difference in surface morphology between the two materials. PGA Homopolymer: T0: Smooth, homogeneous surface without apparent structural defects (Figure 4a); T7 (600 mg/dL): Initial surface erosions with small depressions and shallow fissures (Figure 5a); T14 (600 mg/dL): Deeper fissures, rough surface topography with incipient delamination areas (Figure 5b); T21: Generalized erosions, plate-like delamination, and deep fissuring (Figure 5c); T28 (600 mg/dL): Highly degraded surface with extensive crater formation, deep fissures, plate-like delamination, and substantial mass loss (Figure 5d); PGLA 910 Copolymer: T0: Smooth, regular surface (Figure 4b); T7 (300 mg/dL): Rough surface with microporosities and small cavities (Figure 5e); T14 (600 mg/dL): Moderately eroded surface with some roughness (Figure 5f); T21 (600 mg/dL): Superficial fissures, but overall structure preserved (Figure 5g); T28 (600 mg/dL): Relatively preserved surface with only moderate roughness and minimal fissuring, demonstrating superior resistance to surface degradation compared to PGA (Figure 5h).

Figure 5. Baseline surface morphology of PGA and PGLA 910 sutures. (a) PGA homopolymer (T0) (×184, 200 µm); (b) PGLA 910 copolymer (T0) (×125, 500 µm). Both sutures exhibit smooth, regular, and intact surfaces without apparent structural defects.

Figure 5. Baseline surface morphology of PGA and PGLA 910 sutures. (a) PGA homopolymer (T0) (×184, 200 µm); (b) PGLA 910 copolymer (T0) (×125, 500 µm). Both sutures exhibit smooth, regular, and intact surfaces without apparent structural defects.

Figure 6. Time-dependent surface degradation of sutures under severe hyperglycemic conditions (600 mg/dL glucose). (a–d) Scanning electron microscopy (SEM) images of PGA sutures after (a) 7 days, (b) 14 days, (c) 21 days, and (d) 28 days of exposure. (e–h) SEM images of PGLA 910 sutures after (e) 7 days, (f) 14 days, (g) 21 days, and (h) 28 days of exposure. Images were acquired at the following magnifications and scale bars: (a) 216×, 200 μm; (b) 395×, 200 μm; (c) 125×, 500 μm; (d) 250×, 200 μm; (e) 159×, 500 μm; (f) 125×, 500 μm; (g) 250×, 200 μm; and (h) 250×, 200 μm.

Figure 6. Time-dependent surface degradation of sutures under severe hyperglycemic conditions (600 mg/dL glucose). (a–d) Scanning electron microscopy (SEM) images of PGA sutures after (a) 7 days, (b) 14 days, (c) 21 days, and (d) 28 days of exposure. (e–h) SEM images of PGLA 910 sutures after (e) 7 days, (f) 14 days, (g) 21 days, and (h) 28 days of exposure. Images were acquired at the following magnifications and scale bars: (a) 216×, 200 μm; (b) 395×, 200 μm; (c) 125×, 500 μm; (d) 250×, 200 μm; (e) 159×, 500 μm; (f) 125×, 500 μm; (g) 250×, 200 μm; and (h) 250×, 200 μm.

Figure 7. Effect of glucose concentration on the surface morphology of suture materials after 28 days of incubation. Scanning electron microscopy (SEM) images of (a–c) PGA sutures and (d–f) PGLA 910 sutures. Images were obtained at 250× magnification for (a), (b), (d), and (e) (scale bars = 200 μm); at 125× magnification for (c) (scale bar = 500 μm); and at 80× magnification for (f)(scale bar = 1 mm).

Figure 7. Effect of glucose concentration on the surface morphology of suture materials after 28 days of incubation. Scanning electron microscopy (SEM) images of (a–c) PGA sutures and (d–f) PGLA 910 sutures. Images were obtained at 250× magnification for (a), (b), (d), and (e) (scale bars = 200 μm); at 125× magnification for (c) (scale bar = 500 μm); and at 80× magnification for (f)(scale bar = 1 mm).

3.4. Statistical Analysis and Correlations

Two-way ANOVA demonstrated significant effects of time (p < 0.001), material (p < 0.001), and their interaction (p = 0.008) on maximum load. Glucose concentration alone did not significantly affect degradation (p = 0.153). The complete ANOVA results are presented in Table S5 ( Supplementary Materials ). The Kruskal-Wallis test confirmed the significance of the time factor (p < 0.001) for both materials, corroborating the parametric findings ( Table S6 ). Tukey's post-hoc test revealed that T7 vs T0 was not significant (p = 0.080), indicating maintenance of mechanical strength during the first week. All time points from T14 onward showed significant differences from T0 (p < 0.001). Complete Tukey results are presented in Table S8 . Spearman's correlation demonstrated a very strong positive correlation between residual maximum load and medium pH: PGA (ρ = 0.89; p < 0.001) and PGLA 910 (ρ = 0.86; p < 0.001), as illustrated in Figure 3 , validating the autocatalytic nature of the hydrolytic degradation mechanism ( Table S9 ).

4. Discussion

The present study systematically and integratively investigated the in vitro degradation profile of two of the most widely utilized absorbable suture materials—PGA homopolymer and PGLA 910 copolymer—under varying glucose concentrations. By combining mechanical, chemical (pH), and morphological (SEM) analyses, this exploratory study provides a multi-faceted perspective on the degradation process and the potential impact of hyperglycemia.

Mechanical Performance and Degradation Kinetics

Both materials maintained mechanical strength during the first 7 days, with no significant changes relative to T0 (T7 vs T0: p = 0.080). This finding indicates that neither material undergoes significant hydrolytic degradation within the first week under the experimental conditions employed. However, a progressive loss of mechanical strength occurred from day 14 onward, culminating in near-complete degradation between day 21 and day 28. The progressive degradation pattern observed in this study is consistent with the autocatalytic hydrolytic mechanism reported in the literature [11,22]. Szabelski and Karpiński [14] observed that SafilQuick+ (PGA) and MonosynQuick (glyconate) sutures lost statistically significant tensile strength within 9–12 days, while Novosyn (PGLA 910) and Monoplus (PDS) maintained strength throughout the study period. Our findings extend this observation by demonstrating that under hyperglycemic conditions, PGLA 910 retains 3.4-fold higher residual strength than PGA at 28 days, further reinforcing the importance of polymer selection in patients with metabolic disorders. The acceleration of degradation after day 14 suggests a critical threshold effect, wherein the accumulation of acidic degradation products catalyzes further ester bond cleavage, leading to rapid polymer chain scission and loss of mechanical integrity [7,8,27].

Comparative Degradation

The most significant finding of this study is the superior resistance of PGLA 910 to hydrolytic degradation compared to PGA. At T28, PGA retained only 1.12 N (8.4% of T0 strength), while PGLA 910 retained 3.81 N (26.2% of T0 strength). Although both materials underwent near-complete degradation, PGLA 910 demonstrated a 3.4-fold higher residual strength than PGA at T28, confirming its superior resistance to hydrolytic degradation. This difference is primarily attributable to the distinct chemical and physical structural characteristics of each polymer. PGA, being a homopolymer with higher crystallinity (approximately 50%), possesses a more ordered molecular structure [3,27]. For hydrophilic polyesters such as PGA and PGLA, water penetrates the bulk material within hours; hence, hydrolysis is predominantly bulk-mediated from the outset [11]. The surface cracking observed by SEM is a consequence of bulk degradation: as the interior degrades and loses mass, the outer shell collapses and fissures. This is not surface erosion (as seen in polyanhydrides) but rather a morphological manifestation of heterogeneous bulk hydrolysis. The more pronounced surface alterations observed in PGA reflect this mechanism, as the polymer surface becomes progressively compromised while the bulk material may retain mechanical integrity. In contrast, PGLA 910, as a copolymer containing 10% lactic acid, exhibits a more amorphous structure, which may facilitate water penetration throughout the material volume [12]. However, both polymers undergo bulk erosion; the differences in surface morphology (more fissured PGA vs. more preserved PGLA) can be attributed to differences in degradation kinetics, crystallinity, and the rate of diffusion of oligomers out of the polymer matrix, rather than fundamentally different erosion mechanisms. Interestingly, despite this theoretical predisposition to bulk degradation, PGLA 910 demonstrated superior surface preservation compared to PGA. This may be attributed to differences in degradation kinetics, where PGLA 910 undergoes more uniform degradation throughout the material volume, potentially resulting in less dramatic surface changes [13]. These findings are consistent with recent comparative studies on absorbable suture degradation. Szabelski and Karpiński [14] reported that PGLA-based sutures maintained higher mechanical integrity than PGA-based counterparts under hydrolytic conditions, corroborating the superior performance of the copolymer observed in the present study. Similarly, the progressive degradation pattern and the critical role of pH in accelerating polymer chain scission align with the work of von Burkersroda et al. [22,23] and more recent investigations on polyurethane-based surgical adhesives [10], which highlighted the importance of buffering capacity in preclinical degradation testing.

Effect of Hyperglycemia

The presence of glucose in the incubation medium did not significantly affect degradation of either material, as evidenced by the non-significant effect of glucose concentration in the ANOVA (p = 0.153). This finding suggests that glucose alone, in the concentrations tested (0–600 mg/dL), does not substantially alter the hydrolytic degradation kinetics of these suture materials under the experimental conditions employed. However, pH monitoring showed progressive acidification for both materials, consistent with the release of acidic degradation products. The strong Spearman correlation (ρ > 0.85) between residual force and pH confirms that the degradation mechanism is predominantly governed by autocatalytic hydrolysis, wherein acidic byproducts further catalyze ester bond cleavage. The near-complete degradation by 28 days suggests that the hydrolytic process is significantly more rapid than previously reported in the literature [11,33].

Morphological Analysis and Degradation Mechanisms

SEM analysis was fundamental for elucidating how degradation occurs at the microstructural level and for identifying the superior resistance of PGLA 910 to surface degradation. For PGA, degradation manifested as progressive surface erosions (T7–T14), progressing to fissures and plate-like delamination (T21), culminating in extensive crater formation and mass loss at T28. This pattern is consistent with a predominantly surface erosion mechanism, where hydrolysis proceeds from the external surface inward, leading to progressively more severe surface alterations [11]. For PGLA 910, degradation was substantially less apparent on the surface, with only moderate roughness and minimal fissuring observed even at T28. This suggests that PGLA 910 undergoes a more uniform degradation process throughout the material volume, resulting in less dramatic surface changes while maintaining comparable mechanical strength [12]. SEM also revealed a clear effect of hyperglycemia on morphology. For the same material and time condition, increasing glucose concentration resulted in a rougher surface with more fissures, deeper erosions, and more extensive delamination, providing strong visual evidence of hydrolysis acceleration. Interestingly, while glucose concentration alone did not significantly affect mechanical properties in the ANOVA (p = 0.153), the morphological changes observed suggest that hyperglycemia may influence surface characteristics before mechanical compromise becomes detectable. This apparent discrepancy may be explained by the fact that mechanical failure requires a critical threshold of polymer chain scission, whereas surface morphological changes—such as fissuring and delamination—can occur at earlier stages of degradation. The polymer surface is directly exposed to the surrounding medium, making it more susceptible to the effects of glucose-induced alterations in water activity and polymer-solvent interactions, while the bulk material may retain its mechanical integrity until a critical accumulation of chain scission events occurs.

Clinical Implications

The results have direct clinical implications for the surgical management of diabetic or hyperglycemic patients [28]. The finding that PGLA 910 demonstrates superior resistance to hydrolytic degradation compared to PGA suggests that PGLA 910 may offer a greater safety margin for patients with diabetes or uncontrolled hyperglycemia when extended mechanical integrity is required. However, in vivo validation is essential before definitive clinical recommendations can be made, as the inflammatory and biochemical responses to degradation products (glycolic vs. lactic acid) must also be considered. However, clinicians should be aware that both materials undergo near-complete degradation within 4 weeks, and alternative suture materials should be considered for procedures requiring prolonged mechanical support [12]. The superior surface preservation of PGLA 910 may translate to better resistance to bacterial colonization [21], reduced foreign body reaction, and more predictable degradation kinetics in the clinical setting. However, suture selection should consider a balance between strength, flexibility, degradation rate, and surface stability, taking into account the patient's metabolic status and the specific requirements of the surgical procedure. The development of functional sutures with antimicrobial properties, as explored by Lou et al. [10,25], represents a promising direction for future research in diabetic patients.

Translation to Clinical Practice

The in vitro findings of this study provide a mechanistic foundation for clinical decision-making, but translation to patient care requires careful consideration of several factors. First, the absence of a significant glucose effect on mechanical degradation (p = 0.153) suggests that under the controlled in vitro conditions of this study, the direct chemical effect of glucose on suture hydrolysis is limited. However, this finding does not exclude an important clinical role of hyperglycemia, as chronic elevations in glucose contribute to the formation of advanced glycation end-products, oxidative stress, microvascular dysfunction, immune dysregulation, impaired tissue repair, and increased susceptibility to infection [5,25]. In the clinical setting, hyperglycemia is the driving factor of diabetic complications through altered metabolic pathways that are activated due to high glucose concentrations, and it is impossible to separate hyperglycemia from its consequent complications. Second, the 3.4-fold higher residual strength of PGLA 910 at 28 days indicates that this material may offer a greater safety margin in patients with delayed wound healing, a common feature in diabetic populations [14]. The progressive degradation pattern observed—maintenance of strength during the first week followed by rapid decline after day 14—aligns with the window of highest mechanical demand in most surgical wounds [11,12,14]. This suggests that both materials provide adequate support during the critical early healing phase, but PGLA 910 offers extended integrity that may be particularly valuable in patients at risk for wound dehiscence [13]. To advance these findings toward clinical application, a stepwise translational approach is recommended, as previously outlined for preclinical biomaterial testing [9,28]: (1) Ex vivo validation using human diabetic tissue explants to assess suture-tissue interactions under controlled conditions; (2) In vivo evaluation in diabetic animal models to confirm degradation kinetics and tissue response in a physiologically complex environment [16]; (3) Prospective clinical trials comparing suture materials in diabetic patients undergoing similar surgical procedures, with endpoints including wound healing time, infection rates, and dehiscence incidence [14,31]. Such studies would not only validate the present findings but also establish evidence-based guidelines for suture selection in the growing diabetic population worldwide [14].

Limitations of the Study

This study has several limitations that should be acknowledged. First, the use of a surgeon's knot configuration, while clinically relevant and representative of the weakest point in a suture, introduces a stress concentrator that may influence the failure mode and mask the true degradation kinetics of the bulk material. Straight-pull tensile testing would be more appropriate for characterizing intrinsic polymer hydrolysis. Second, as an in vitro study, this investigation does not account for the complex biological environment encountered in vivo, including inflammatory responses, enzymatic activity, and tissue interactions. Third, mass loss was not measured, and the sample size (n = 4 per condition) was relatively small, reflecting the exploratory nature of this study. Fourth, the use of PBS as an incubation medium, while providing a controlled environment, does not replicate the complex in vivo environment in the presence of diabetes mellitus, which includes inflammatory mediators, enzymatic activity, immune cells, and dynamic tissue interactions. Additionally, the sterile in vitro conditions do not account for the potential effects of bacterial colonization or the host inflammatory response, which could influence the degradation kinetics and mechanical performance of the sutures in a clinical setting. Fifth, maintaining stable hyperglycemic conditions for up to 28 days does not reflect the clinical fluctuations of glycemic states in patients with diabetes, where glucose levels vary considerably over time. It should be noted that no additional sterilization methods, such as gamma radiation, were applied to the sutures beyond the manufacturer's original sterile packaging. However, if such methods were used in clinical practice, they could potentially affect the mechanical properties of the materials, and this should be considered in future studies. Future investigations should incorporate larger sample sizes, additional metabolic factors such as advanced glycation end-products and inflammatory mediators, and intermediate time points (e.g., day 10, day 24) to precisely characterize degradation kinetics. Finally, the recommendation that PGLA 910 may be preferable for diabetic patients requires in vivo validation in diabetic animal models and prospective clinical trials before it can be translated into clinical practice.

5. Conclusions

The present investigation establishes that both PGA and PGLA 910 sutures maintain mechanical integrity during the first 7 days of in vitro degradation (T7 vs T0: p = 0.080), confirming their adequacy for wound closure during the early healing phase. From day 14 onward, however, both materials exhibit progressive strength loss, with near-complete degradation by day 28—PGA retaining only 8.4% and PGLA 910 retaining 26.2% of initial tensile strength at T28—demonstrating that neither material is suitable for applications requiring mechanical support beyond 3 weeks under the experimental conditions employed.

PGLA 910 consistently outperformed PGA across all time points, exhibiting 3.4-fold higher residual strength at T28 (3.81 N vs. 1.12 N; p < 0.001), attributable to its copolymeric structure and more uniform degradation kinetics. These findings were obtained using a clinically relevant surgeon's knot configuration, which replicates the weakest point of the suture and enhances the translational value of the results.

Glucose concentration alone did not significantly influence degradation (p = 0.153), indicating that polymer chemistry, rather than ambient glucose levels, is the primary determinant of hydrolytic resistance. The strong positive correlation between mechanical strength loss and medium acidification (ρ = 0.86–0.89; p < 0.001) corroborates autocatalytic hydrolysis as the dominant degradation mechanism for both materials.

For patients with diabetes mellitus or uncontrolled hyperglycemia, PGLA 910 offers extended mechanical integrity compared to PGA, which may be advantageous when prolonged wound support is required. However, clinicians should consider that the choice of suture material in diabetic patients must be multifactorial, taking into account not only mechanical properties but also the inflammatory response to degradation products and the patient's overall metabolic status. Both materials undergo near-complete degradation within 4 weeks, necessitating alternative suture selection for procedures requiring prolonged support.

The translational pathway for these findings requires sequential validation: ex vivo confirmation in human diabetic tissue explants; in vivo assessment in diabetic animal models; and prospective clinical trials comparing wound healing outcomes and complication rates between PGA and PGLA 910 in diabetic populations. Future investigations should incorporate additional metabolic factors such as advanced glycation end-products and inflammatory mediators, and include intermediate time points (e.g., day 10, day 24) to precisely characterize degradation kinetics.