The artificial urine (AU) was formulated with the following concentrations: urea (12.13 g/L), sodium chloride (3.17 g/L), ammonium chloride (0.81 g/L), calcium chloride (0.35 g/L), magnesium sulfate heptahydrate (0.5 g/L), sodium bicarbonate (0.17 g/L), sodium sulfate (0.13 g/L), sodium dihydrogen phosphate (0.5 g/L), and sodium hydrogen phosphate (0.06 g/L), yielding a pH of 5.85 at 25 °C.

For the subsequent in vitro degradation experiments, novel ureteral stents were precisely cut into 1 cm segments from 5 cm tubes and weighed meticulously. The polymer samples were immersed in 5 mL of the AU solution, which was replenished on a weekly basis. Degradation assays were conducted in a shaking incubator at 37 °C with continuous agitation for 24 h daily. Samples were withdrawn at predetermined intervals, thoroughly rinsed with distilled water, and dried under vacuum to a constant weight. The calculation of mass loss was conducted in accordance with the stipulated formula: Mass loss = W i − W t W i × 100 % Water uptake = W w − W t W i × 100 %

W_i is the initial weight of the PTG copolymer material before degradation, W_t is the dry weight of the PTG copolymer material after degradation, and W_w is the wet weight of the PTG copolymer material after degradation.

The lipase solution, sourced from Aspergillus oryzae lipase, was meticulously prepared by mixing the enzyme with PBS (pH 7) in a 1:1 ratio, resulting in an AP solution.

Prior to the in vitro degradation analysis with artificial urine, novel ureteral stent tubes were precisely cut into 1 cm segments from 5 cm tubes and weighed with care. The polymer samples were then immersed in 1 mL of the lipase solution, with the solution refreshed daily to preserve the enzyme’s activity. Degradation experiments were conducted in a shaking incubator at 37 °C, with continuous agitation for 24 h daily. At established time points, samples were extracted from the degradation medium, rinsed with distilled water, and excess moisture was removed using filter paper. Subsequently, the samples were dried under vacuum to a constant weight, with mass loss calculated as previously described.

Figure 4a illustrates the macroscopic appearance of four distinct PTG ureteral stent groups at various fragmentation time points in artificial urine. PTG4060 demonstrated complete fragmentation and a loss of supportive function by the fourth week, while PTG5050 exhibited the same by the fifth week. In contrast, PTG6040 and PTGGDG maintained their structural integrity until the sixth week. Despite the formation of cracks and the shedding of surface fragments, the PTG stents retained their overall tubular structure. This phenomenon is primarily attributable to the moderate degradation rate of the intermediate PTG5050 layer, which buffers the rapid degradation of the outer stent layer, while the slower degradation of the inner PTG6040 layer provides essential support. Consequently, by the sixth week, the macroscopic appearances of PTGGDG and PTG6040 were notably similar. Such characteristics are beneficial for ensuring ureteral patency during clinical placements, preventing sudden collapses of PGA polyester materials due to intrinsic degradation, which can lead to blockages from fragmented materials and subsequent hydronephrosis, causing iatrogenic kidney dysfunction. Uriprene™, a material developed previously, presented challenges related to unpredictable degradation rates. This resulted in the abrupt disintegration of the material and temporary obstructions in animal trials, leading to adverse outcomes (Chew et al., 2013). In response to the challenges identified, Wang et al. developed a PCL/PLGA ureteral stent. This was produced using electrospinning technology and was designed to degrade progressively from the distal bladder to the proximal kidney. The stent was developed to facilitate drainage and minimise ureteral obstruction (Wang et al., 2015). The PTGGDG ureteral stent that has been developed is uniquely designed to degrade from the exterior inward, with the PTG4060 layer degrading first. The configuration of the PTG6040 is such that it maintains its structural integrity while facilitating urine flow, thereby aiding in the clearance of surface degradation debris and potentially mitigating the formation of stone. This design may indeed signify a promising trajectory for the future development of biodegradable ureteral stents.

As demonstrated in Figure 5a, the PTG ureteral stent demonstrated a substantial decrease in weight during the degradation process in artificial urine. It is imperative to note that the composition of the copolymer in the ureteral stent exerts a pivotal influence on the hydrolysis rate within artificial urine. It is noteworthy that the weight change in the PTG ureteral stent demonstrated relative stability during the initial 14-day period. A significant escalation in mass loss was observed following this period, particularly evident in the PTG4060 variant. Following a 28-day period of degradation, there was a significant increase in weight loss for the PTG4060 ureteral stent, from 13.40% ± 0.95% at 10 days to 44.10% ± 2.83%. In addition, by the 28th day, the PTG4060 ureteral stent exhibited signs of fragmentation and subsequent collapse. The PTG5050 ureteral stent demonstrated signs of fragmentation on the 35th day, resulting in a mass loss of 18.11% ± 0.33%. In contrast, the PTG6040 variant demonstrated a prolonged degradation period, not fracturing until the 42nd day, with a corresponding mass loss of only 12.59% ± 0.31%, which is significantly lower than that observed in PTG4060. The degradation rate of PTG in artificial urine is primarily influenced by the proportion of GA in the stent, as the acetyl units present in GA are particularly prone to hydrolytic attack (Zurita et al., 2006). This finding suggests that increasing the GA content in the copolymer may enhance the degradation of PTG ureteral stents.

The PTGGDG ureteral stent underwent fragmentation after 42 days, resulting in a substantial weight reduction of 55.96% ± 1.70%. This phenomenon can be attributed to the composition of the stent’s outermost layer, which consists of PTG4060. Upon contact with AU, PTG4060 preferentially engages in hydrolytic degradation, which in turn exposes the innermost layer, PTG6040. Consequently, the degradation duration of the PTGGDG stent is considerably longer than that of the PTG4060 ureteral stent, resulting in a greater mass loss compared to PTG6040. An extended presence of stents in the body can elevate the risk of urinary tract infections and enhance the rate of surface encrustation (El-Hayek et al., 2017). The recommended retention time for ureteral stents is approximately 3 weeks, as this can significantly minimize the occurrence of urinary tract infections (Wang et al., 2022). Typically, the clinical implantation duration for ureteral stents is around 4–8 weeks. In contrast, the PTGGDG ureteral stent is capable of degrading within 6 weeks into small particles that can be evacuated through the urine, thereby meeting the clinical demands for ureteral stents.

Furthermore, the water absorption rates of the PTG ureteral stents were monitored throughout the degradation process, as illustrated in Figure 5b. As the degradation time increased, a rising trend in water absorption rates was exhibited by all four groups of PTG ureteral stents. By day 28, the water absorption rate of the PTG4060 ureteral stent had reached 43.79% ± 3.82%, whereas the PTG6040 group had only achieved a water absorption rate of 17.71% ± 0.31% at the time of fracture on day 42. This disparity can be attributed to the heightened hydrophilicity of PTG copolymers with higher GA content, as demonstrated in Figure 3a. The absorbed water molecules can infiltrate the material’s interior, inducing a plasticizing effect on the copolymer, which increases the mobility of the polymer chain segments and allows for enhanced hydration, promoting accelerated degradation (Edlund and Albertsson, 1999). This observation is consistent with the trend in mass loss of the PTG ureteral stents.

Figure 5c presents the relative GA content of the PTG6040 ureteral stent in AU. Over time, the relative GA content of PTG6040 exhibits a decline, which confirms that the PGA segments are preferentially degraded during the degradation process in AU.

Urinary stones are intricately linked to infections, and the pH level of urine plays a critical role in influencing the pathogens present in the urinary tract as well as the formation and progression of stones, thereby significantly affecting patient health (Khan et al., 2025). For example, an elevated pH can result in the precipitation of struvite stones, often triggered by urease produced by bacteria associated with urinary tract infections, such as those from the Proteus genus, or as a consequence of the presence of implanted ureteral stents (Thomas and Tolley, 2008). Conversely, a lower pH can promote uric acid precipitation in an insoluble form, which substantially raises the risk of stone formation (Kamel et al., 2002). Hence, the design of ureteral stent materials should consider the effects of their degradation products on urinary pH. Previous studies from our team have indicated that PTMC does not generate acidic degradation products during its degradation process (Yang et al., 2015). However, the hydrolysis of GA leads to the formation of acidic carboxylic products, resulting in a decrease in pH (Díaz-Celorio et al., 2013). Importantly, the PTG ureteral stent that we have developed maintains a pH range between 5.5 and 6.0 during degradation (Figure 5d), which aligns with the normal physiological range of urine pH. Although the hydrolysed degradation products of the PTG ureteral stent have been shown to reduce urine pH, they are unlikely to cause significant acidosis or tissue irritation when used in urology.

Figure 6a shows an SEM image of the stent prior to degradation. Figures 6b,c present SEM microphotographs of the surfaces of PTG4060 and PTGGDG ureteral stent samples at different degradation periods. As time progresses, the surfaces of both PTG4060 and PTGGDG samples become rougher after 1 week of hydrolysis. Cracks begin to appear in the second week, and the number of cracks significantly increases by the third week. There are more fragments in PTG4060 than in PTGGDG at the third week. This is attributed to the fact that the former degrades more quickly and experiences greater weight loss (Figure 5a). Additionally, the three-layer structure of PTGGDG provides some support from the innermost layer, PTG6040, which reduces the formation of fragments.

At the third week of degradation, we compared the surface morphologies of the four groups of samples. As illustrated in Figure 6d, the SEM microphotographs of different PTG ureteral stent samples at week three reveal their distinct surface morphologies. As the GA content in the PTG ureteral stents increases, greater corrosion is exhibited on the surfaces of the samples. As the degradation time increases, the gradual precipitation of inorganic salts from artificial urine occurs, resulting in interactions with the material surface. The PTG6040 specimen displays indications of surface encrustation, which suggests that elevated TMC content may augment the likelihood of such encrustation. This phenomenon is likely attributable to the increased presence of polar groups on the surface, thereby facilitating crystalline formation. It is noteworthy that the PTG6040 surface displays no significant corrosion, indicating that its degradation rate is lower than that of the other groups (Figure 5a). Conversely, the PTGGDG group demonstrated surface erosion by the third week, with no indications of encrustation. This reduction in encrustation risk may be attributed to the degradation of the outer layer, which has effectively diminished the potential for stone formation. This finding confirms that the multilayer structure of PTGGDG can effectively prevent encrustation while maintaining an ideal degradation rate.

As the degradation time of the stent increases, the relative molecular weight of the PTG ureteral stent samples gradually decreases (Figure 7a). It is noteworthy that the relative molecular weight of the PTG ureteral stent undergoes a precipitous decline within the initial 14 days, while the associated mass loss remains comparatively negligible during this same period (see Figure 4a). This observation is in accordance with the intrinsic degradation patterns that are generally observed in polyester materials (Li et al., 2022). The findings of this study indicate that the PTG ureteral stent undergoes intrinsic degradation when immersed in artificial urine.

Further analysis of the GPC data reveals that the molar mass dispersity (Đ_M) during hydrolytic degradation in AU shows an increasing trend (Figure 7b). This rise is due to the accumulation of low molecular weight degradation products before they dissolve (Díaz-Celorio et al., 2013).

The trend of the Tg during the degradation process of PTG ureteral stents in AU is illustrated in Figure 7c. Theoretically, as the PGA segments degrade preferentially (as depicted in Figure 5c), the Tg of the copolymer should gradually decrease over the course of degradation. As demonstrated in Figure 7c, the Tg values for PTG4060, PTG5050, and PTG6040 are consistent with the anticipated trend, suggesting a uniform behaviour among these copolymers during the degradation process.

Figure 8a illustrates the mass loss of four groups of PTG ureteral stent samples over time in AP solution. The extent of mass loss correlates with the content of the TMC segment. Notably, PTG6040 shows a mass loss of 66.12% ± 1.60% upon complete degradation by day 10, while PTG4060 only exhibits a mass loss of 8.64% ± 1.33% during the same period. Additionally, the hydrolytic mass loss recorded for PTG6040 at day 10 is relatively low, at 7.78% ± 0.20%. These results indicate that a higher concentration of TMC segments is associated with greater mass loss, and the degradation facilitated by lipase occurs at a rate faster than that of hydrolysis.

The introduction of TMC segments can significantly accelerate the enzymatic degradation rate of PTG copolymers, and the degradation rate of PTG can be modulated by increasing the TMC content in the polymer. In Table 2, referencing our previous study (Yang et al., 2015), we calculated the rate constant k for the degradation process of ureteral stents made from PTG copolymers: R t 2 = R t 1 − k t 2 − t 1

In the equation, R _t represents the weight loss rate of the PTG copolymer at time t . The findings reveal that the degradation constants for PTG4060, PTG5050, PTG6040, and PTGGDG during lipase degradation are 0.48, 1.56, 5.06, and 4.29, respectively, with correlation coefficients exceeding 0.9, indicating a strong fit. Overall, the apparent enzymatic degradation rates of PTG show significant variability. Notably, as the number of GA segments increases, the degradation rate of PTG decreases. This is attributed to the fact that lipase-mediated degradation of TMC primarily occurs through surface erosion, which is influenced by the chain segments available for enzyme contact as well as the mobility of these segments. The incorporation of GA units is likely to enhance segment rigidity, thus reducing the accessibility of TMC-rich regions to enzyme active sites and consequently diminishing the number of effective cleavage sites. This observation is consistent with the findings of Liu et al. (2012). Overall, the results indicate that the degradation of PTG ureteral stents in lipase solution adheres to a controllable surface erosion mechanism that aligns with first-order kinetic equations.

Our study also recorded the water absorption characteristics of PTG ureteral stents throughout the enzymatic degradation process, as illustrated in Figure 8b. The water absorption rate of the PTG ureteral stents shows a gradual upward trend with increasing degradation time. It is noteworthy that PTG6040 achieves a water absorption rate of 30.40% ± 2.40% by day 10. In contrast, PTG4060 demonstrated the lowest water absorption rate, reaching only 15.68% ± 1.30% by day 13. Nevertheless, this rate is still higher than its hydrolytic water absorption rate of 13.80% ± 1.50% measured on day 14. This outcome suggests that lipase-mediated degradation remains the dominant process in the AP solution.

The relative GA content of the PTG6040 ureteral stent in AP is illustrated in Figure 5c. The relative GA content of PTG6040 increases with the extension of time. This confirms that TMC segments are preferentially degraded during the degradation process in AP.

Figure 4b illustrates the macroscopic fragmentation of four groups of PTG ureteral stents at various time points in AP solution. By day 13, both PTGGDG and PTG5050 ureteral stents had completely fractured and lost their structural support. PTG6040 also completely fragmented and lost its supporting function by day 10. In contrast, PTG4060 did not show complete fragmentation by day 13. Additionally, the color of the sample surfaces transitioned from white to yellow, which may be attributed to the accumulation of degradation products from enzymatic or oxidative processes on the surface of the samples.

Figure 9a shows an SEM image of the stent prior to degradation. Figures 9b,c present SEM micrographs of the surfaces of PTG4060 and PTGGDG ureteral stents immersed in AP solution across various degradation periods. With the progression of time, the surfaces of both PTG6040 and PTGGDG samples became increasingly rough after just 1 day in lipase. By day 4, cracks began to appear, and by day 7, a significant increase in the number of cracks and pits was observed. Notably, PTG6040 developed more cracks and deeper pits compared to PTGGDG by day 7. This difference can be attributed to the faster degradation rate and greater mass loss of PTG6040 in the lipase environment (illustrated in Figure 8a). The observed increase in surface roughness over time further supports the presence of a surface erosion mechanism during degradation, which aligns with findings related to pure PTMC, as reported in our prior work (Yang et al., 2015).

On the 7th day of degradation, we compared the surface morphology of the four sample groups. Figure 9d presents SEM micrographs of various PTG ureteral stent samples taken at day 7, illustrating the differences in their surface morphologies. Increased TMC content in the PTG ureteral stents was associated with greater surface corrosion. The SEM findings suggest that a higher concentration of TMC segments leads to a slower degradation rate of the PTG ureteral stents, which aligns with the degradation rate constants obtained from the mass loss data.

The changes in the relative molecular weight (Mn) and the polydispersity index (ĐM) of the samples during degradation are shown in Figures 10a,b. It can be observed that the molecular weight of the PTG ureteral stents decreases sharply within 6 days. During degradation, lipase rapidly attacks the PTMC segments, and the degradation products gradually disperse into the enzyme solution. However, when the rate of dispersion of degradation products into the AP solution is slower than the degradation process, a significant amount of degradation products will be retained in the samples, leading to a decrease in molecular weight accompanied by mass loss. For example, the amount of enzyme used is relevant since enzymes can act as surfactants to promote the diffusion of small molecular segments into the solution (Li et al., 2019; Yang et al., 2015). Although GA undergoes intrinsic degradation, the crystalline regions in PTG can hinder the diffusion and distribution of enzymes within the samples. Therefore, surface erosion degradation and incomplete intrinsic degradation coexist in PTG in the AP solution. The accumulation of small molecular degradation products on the sample surface results in a slow increase in ĐM. This is consistent with our previous findings on the enzymatic degradation mechanism of poly (2,2′-dimethyltrimethylene carbonate-co-ε-caprolactone) (Zhang et al., 2023).

Figure 10c depicts the trend of the Tg of PTG ureteral stents throughout their degradation in AP. Theoretically, as the PTMC segments preferentially degrade, the Tg of the copolymer should steadily increase during degradation, As shown in Figure 7c, the Tg values for PTG4060, PTG5050, and PTG6040 exhibit an upward trend from the beginning of degradation until day 4, at which point a subsequent decrease is observed. This initial rise in Tg is primarily due to the preferential degradation of TMC segments by lipase. Following this, the hydrolytic degradation of GA in the PBS solution occurs, leading to a decrease in Tg as the GA chain segments disintegrate. These observations reinforce the notion that PTG undergoes both surface erosion degradation and incomplete intrinsic degradation within the AP solution. Furthermore, the plasticization effect caused by the accumulation of small molecular degradation products post-degradation contributes to the later decline in Tg, consistent with findings from our earlier studies (Zang et al., 2025).