Biodegradable polymer scaffolds have been used to construct new tissues or organs by culturing cells on them. The researchers conducted a 9-week degradation, structure, and performance study of a fibrous nonwoven polyglycolide (PGA) scaffold (2 mm thick, 10 mm diameter) in tissue culture medium under 37 mixed conditions. After 3 days of in vitro culture, there was a slight increase in the mass of the scaffold due to hydration and adsorption (56%), but then decreased. After 9 weeks, only 125% quality. Measured with differential scanning calorimetry (DSC), the melting point of the stent increased from 218 in the first three weeks1 down to 1860℃。In the subsequent time (6 and 9 weeks), no melting peaks were found. Within the first 11 days, the crystallinity of the stent doubled, but then decreased. The glass transition temperature (36-39) of the degraded scaffold is lower than that of the dry start scaffold (400), which may be due to the plasticization of absorbed water and other low molecular weight molecules. A PCA scaffold without cells completely loses its structural integrity and mechanical strength within 2 to 3 weeks. In contrast, the neocartilage construct, which was regenerated by the PGA scaffold and bovine chondrocytes, maintained structural integrity throughout the in vitro culture study. After 12 weeks of in vitro culture, the biomechanical properties of the neocartilage construct reached the same order of magnitude as normal bovine cartilage. 1. IntroductionHuman organ transplants have saved or improved countless lives. However, organ transplantation is severely limited due to a severe shortage of donors. The use of biodegradable synthetic polymers as scaffolds to transplant cells and regenerate new tissues or organs is a revolutionary approach to tissue loss and organ failure. Polymer scaffolds can be used as a guide for three-dimensional tissue regeneration. Biodegradable polymers are absorbed after the purpose of temporary support, allowing for a natural replacement of tissues or organs. Synthetic biodegradable polymers control structural variables and scaffold properties such as molecular structure, molecular weight, hydrophilicity, degradation rate, mechanical properties, batch-to-batch consistency, and processing flexibility. Polyethylene lactide (PGA) fiber nonwoven scaffolds have been used to engineer a variety of tissues, including liver, cartilage, bone, tendon, intestinal, and urethral tissues. In order to understand its mechanism and better control tissue engineering, it is important to understand the mass changes, micro- and macrostructure, and mechanical properties of PGA nonwoven scaffolds during culture. In this paper, an in vitro study of a model PGA nonwoven scaffold was performed to address these issues. 
2. Materials and methods1.BracketsPGA resin has an intrinsic viscosity greater than 11 dl/g。Round PGA non-woven discs (2 mm thick, 10 mm diameter) with staple fiber length of 2-25 inches, fiber diameter of 15 microns, bulk density of 72 mg cubic centimeters (porosity of 954%）。2.In vitro degradationIn order to study the scaffold degradation process while studying in vitro cartilage tissue engineering, the in vitro degradation study of the PGA scaffold employed exactly the same culture conditions as in the study of the biomechanical properties of engineered cartilage, but without the cell seeding process. PGA non-woven discs are sterilized with ethylene oxide. Sterile non-woven medium for discs (DMEM, containing 4500 mg L glucose, 10% fetal bovine serum, 10 mM HEPES, 01 mm MEM non-essential amino acids, 04 mM L-proline, 50 mg L L-ascorbic acid, 50 U cc penicillin, and 50 g cc streptomycin) in an incubator with % CO2 overnight. The pre-wet discs are then transferred to another set of 6-well plates and placed on an orbital shaker at 75 rpm, one discs per well, and 5 mL of medium per well. Change the medium (7 ml per well) for the first time after 3 days. Thereafter, change the medium (7 ml per well) every other day. After the PGA nonwoven disc collapses, change the medium by aspirating the supernatant (1,000 rpm for 10 min) and resuspend the short fibers in 7 ml of medium per well. Degraded scaffolds were sampled at different times and lyophilized for two days prior to structural characterization. The mass of the remaining stents was normalized to the mass of the pre-wet stents (fully lyophilized). 3.Structural characterizationThermal analysis was performed using a differential scanning calorimeter with a heating rate of 20 minutes. The melting point tm is determined from the peak temperature of the internal thermal melting peak. The glass transition temperature tg is determined by the midpoint method on a heating thermometer. The degree of crystallinity is calculated from the enthalpy of the internal thermal melting peak. 4.Measurement of mechanical propertiesThe mechanical properties were measured using a closed compression chamber filled with culture medium and a Vitrodyne V-1000 mechanical tester. The thickness is measured with a micrometer. Before measuring, the specimen is left in fresh medium at room temperature for 1 to 2 hours. The compressive modulus and aggregation modulus were determined at room temperature by confinement compression (10 ms) and confined compressive stress relaxation (20% strain for one hour). The calculation method is the same as that for in vitro engineered neocartilage using a PGA non-woven stent. Apparent permeability is an important physical property that reflects the mass transfer behavior (nutrients, metabolites, etc.) in the scaffold, which is determined by a duplex model. III. Results and DiscussionAfter the first 3 days of incubation, there was a slight increase in the mass of the PGA nonwoven scaffold (Figure 1), which may be due to hydration and adsorption, but the mass has decreased since then. After 9 weeks of incubation, the mass was only 125%。Overall, the mass of the scaffold decreased exponentially, and the data correlated with the first-order degradation kinetics of the scaffold mass (degradation rate constant k=-3.).78x10^(-2)/day, r[gf]b2[/gf]=0.943) fits perfectly. 
Figure 1In vitro degradation of PGA nonwoven scaffolds in 37 culture broths. The microstructure of the scaffold varies with culture time. For the first 11 days, the melting point (tm) of the stent decreased slowly (decline rate k=-0.).40 days, r[gf]b2[gf]=099, Figure 2), after which the rate of decline was accelerated by about 7-fold (k=-2.).79 days, r[gf]b2[ gf]=100）。In the subsequent time (6 and 9 weeks), no melting peaks were found. The amorphous region is susceptible to erosion by water molecules and degrades first. We believe that in the first 11 days, degradation (mass loss) occurs mainly in the amorphous stage. The initial linear decrease in melting point is independent of the total mass, which may be due to the plasticizing effect of absorbed water and other low molecular weight molecules, as well as a slight degradation of the crystalline phase (reduction in grain size). This is also evidenced by the decrease in TM values during pre-wetting (Table I). From day 11 onwards, the crystalline phase begins to play an important role in degradation. The main reason for the rapid decline in TM after the first 11 days is the decrease in crystal size. This is consistent with an increase in crystallinity over the first 11 days followed by a decrease (Figure 3). After that (6 and 9 weeks), due to the massive mass loss, the lattice is severely damaged and no longer behaves as a regular crystalline phase, so the melting peak cannot be detected. 
Figure 2The melting point of the PGA scaffold decreases with the extension of the in vitro degradation time. 
Table IEffect of pre-wetting on DSC results of PGA non-woven stents. 
Figure 3The crystallinity of the PGA scaffold varies with the time of degradation in vitro. The glass transition temperature of the scaffold drops abruptly during pre-wetting (Table I) and continues to decrease during the first 3 days of incubation (Figure 4), mainly due to the plasticizing effect of absorbed moisture and low molecular weight molecules. The rapid increase from days 3 to 14 (Figure 4) may be primarily due to the increase in crystallinity, i.e., the remaining polymer chains in the amorphous region (which contributes to the TG) are more limited due to the increased ratio of the crystalline phase (physical cross-linking) to the amorphous phase. Thereafter, the TG decreases slowly, possibly due to a decrease in the molecular weight of the remaining PGA molecules. 
Figure 4The relationship between the glass transition temperature and the in vitro degradation time of the PGA nonwoven scaffold. As the microstructure changes, the thickness and mechanical properties of the PGA nonwoven scaffold also change with the extension of the in vitro degradation time. After 3 days of incubation, the scaffold thickness increased by approximately 11% (Figure 5), probably due to the adsorption swelling effect of water and various components in the medium, consistent with a slight increase in scaffold mass. From day 3 to day 14, the thickness slowly decreases to approximately the same (2% lower) thickness as the pre-wet stent. After 2 to 3 weeks of incubation, the scaffold eventually disintegrates into fragmented short fibers. The closed compression modulus decreases over time (Figure 6). In the first 3 days, the polymerization modulus increased by about 28% (Figure 6), which may also be due to the effect of initial adsorption. Thereafter, the polymeric modulus decreases over time. The apparent permeability of the stent increases exponentially over time (r[gf]b2[gf]=0.).98, fig. 7). 
Figure 5The thickness of the PGA scaffold as a function of the time to degradation in vitro. 
Figure 6Relationship between the confinement modulus (MO) and aggregation modulus (mA) of PGA nonwoven scaffolds and in vitro degradation time. 
Figure 7Apparent permeability of PGA scaffolds as a function of in vitro degradation time. Changes in thickness and mechanical properties are related to the degradation and microstructural changes of the scaffold. Based on our findings, we speculate that crystallization has a greater effect on the relative equilibrium mechanical properties than amorphous, while the total mass (porosity) of the scaffold has an impact on the mass transfer performance of the medium through the scaffold. The total mass of the stent decreases exponentially (figure 1), resulting in an exponential increase in apparent permeability (figure 7). During the first 11 days, degradation occurs mainly in the amorphous phase, so the polymeric modulus (a type of equilibrium modulus) decreases slowly. Thereafter, the crystallization relative degradation plays an important role, so the polymerization modulus decreases faster. These changes in mechanical properties are consistent with the following results: crystallinity increases over the first 11 days and decreases thereafter; TM decreases slowly over the first 11 days (mainly due to plasticizing effects of adsorbed water and other media components) and then accelerates (mainly due to a decrease in crystal size). The initial adsorption (mass increase in the first 3 days) has an effect on some structural parameters and mechanical properties (e.g., decreased glass transition temperature, increased thickness, and increased aggregation modulus). Contrary to the fact that PGA scaffolds lose their structural integrity and mechanical strength within two to three weeks, PGA-chondrocyte constructs (2 million bovine chondrocytes per scaffold) grow into neocartilage in vitro with biomechanical properties of the same order of magnitude as normal cartilage within 12 weeks (Table II). These results suggest that biodegradable PGA nonwoven scaffolds can serve as a template for regenerating mechanical functional tissues. 
Table iiMechanical properties of PGA nonwoven stents, engineered neocartilage with scaffolds (in vitro culture for 12 weeks, surface removed), and normal bovine cartilage. Fulin Plastics** implantable grade bioresorbable PGA non-woven fabric, the sizes are as follows:
Length: 30 cm Width: 20 cm Thickness Range: 1 mm - 10 mm Density Range: 40 - 100 mg CC Standard density tolerance is 10% of the target value. Density is measured as the sum of the entire PGA nonwoven sheet and can be cut into a variety of geometries, including discs and squares as small as 2 mm.
All stent products are individually packaged in heat-sealed foil pouches and packaged with desiccant in a nitrogen environment. Note: All products provided are non-sterilized Email: li@fulinsujiaocom Company address: 810, Building 1, Sujin International, Zhangmutou Town, Dongguan City, Guangdong Province