Impact of Polymer Degradation on Cellular Behavior in Tissue Engineering

Abstract

Tissue engineering frequently employs biomimetic scaffolds to direct cell responses and facilitate the differentiation of cells into specific lineages. Biodegradable scaffolds mitigate immune responses, stress shielding concerns in load bearing tissues, and the need for secondary or revision surgical procedures for retrieval. However, during the degradation process, scaffold properties such as fiber diameter, fiber porosity, fiber alignment, surface properties and mechanical properties undergo changes that significantly alter the initial properties. This review aims to comprehensively assess the impact of degradation on scaffold properties from the perspective of their effects on cellular behavior by addressing four key aspects of polymer degradation: First, we review the variables that influence scaffold degradation. Second, we examine how degradation impacts scaffold properties. Third, we explore the effects of scaffold degradation products. Finally, we investigate measures to increase tunability of degradation rate. Harnessing and incorporating these degradation mechanisms into scaffold design holds great promise for advancing the development of tissue-engineered scaffolds, ultimately improving their efficacy and clinical utility.

1. Introduction

Tissue engineering uses structure to function interactions to restore and augment tissue repair. Fundamentally, this is accomplished by first understanding and then mimicking the interactions between cells and their surrounding extracellular matrix (ECM) tissue niche. The ECM plays an essential biological role in support of the cellular fractions of tissue by providing biochemical cues, mechanical support, and a niche for colonization and homeostasis of cells.¹ Key ECM components include collagens, glycoproteins, proteoglycans, and associated regulatory proteins.² The mechanical properties of ECM vary depending on their anatomic function. For example, cartilage is required to withstand compressive loads,³ while tendons need to resist largely tensile loads.⁴ Some of these tissue engineered applications include nerve,⁵ vascular,⁶ tendon,⁷ cartilage,⁸ and bone tissue.⁹ One tissue engineering strategy that has been successfully implemented to fulfill these roles is three-dimensional (3D) biomimetic scaffolds.¹⁰ 3D scaffolds can be fabricated using natural or synthetic polymers using well established methods such as gas foaming, phase separation, 3D printing, electrospinning, and meltblowing.¹¹ Natural polymers such as collagen, gelatin, and chitosan are highly biocompatible. However, they lack adequate mechanical properties which is further hindered by their poorly controllable degradation rate,¹² which limits their application in tissue engineering.¹³ Synthetic polymers typically lack the biocompatibility of naturally occurring polymers, but are more easily tunable with regard to their mechanical properties, degradation rates, and fiber parameters that can be tailored towards a specific tissue application. A wide range of synthetic polymers for scaffolds is actively being explored to optimize their use for specific tissue applications by considering their mechanical properties and sometimes their degradation timelines. Synthetic biodegradable polymers frequently utilized in tissue engineering include members of the polyester, polyanhydride, polyphosphazene, and polyurethane chemical families.¹⁴ Specifically, commonly used polymers for these devices are poly-L-lactic acid (PLLA),⁷ poly-glycolic acid (PGA),¹⁵ poly lactic-co-glycolic acid (PLGA),¹⁶ poly-caprolactone (PCL),¹⁷ and polyhydroxybutyrate (PHB). These polymers are particularly notable for their hydrolytic degradation mechanisms, which break down ester bonds in the presence of water, leading to the formation of biocompatible byproducts that can be metabolized or excreted by the body. The associated polymer blends, molecular weights, and crystallinities of each of these specific polymers also dictate their material properties. Even though degradation properties are critical to the overall material and mechanical properties of tissue-engineered devices and are closely linked to specific material-cell interactions, they are generally not evaluated in either in vitro or preclinical studies. These properties should be more prominently incorporated into the design of 3D scaffolds for tissue engineering applications.¹⁸ The aim of this review is to first summarize the impact of scaffold design and surrounding environment on degradation rate. Then the effect of hydrolytic and enzymatic degradation of synthetic polyester scaffolds on different fiber characteristics, molecular properties, and the impact of degradation products on biologic behavior will be explored. Finally, preventive measures for degradation will be examined, and the need to incorporate degradation studies into biomaterial design even if the core polymer is well understood is highlighted.

2. Overview of Polymer Degradation Mechanisms

Biodegradability of a polymer is an essential property for a tissue engineered scaffold.^19,20 Compared to non-absorbable implants, xenografts, or allografts, biodegradable polymers offer the benefits of predictable degradation and remodeling, not requiring a subsequent surgical intervention for its removal, while also mitigating concerns about potential long-term effects at the implantation site, such as foreign body responses,²¹ stress shielding of adjacent tissues,²² or the need for extensive remodeling for complete integration. Degradation encompasses a multifaceted process leading to the reduction in the physical properties of a polymer due to the breakdown of its molecular and chemical structure over time. This process is dictated by both the polymer properties and the external surrounding environment. Among the determinants, three factors stand out: the polymer chemical properties that influence the reactivity with water, the molecular weight referring to the average size of its polymer chains, and the pH of the surrounding environment that can catalyze degradation.²³ All of these core determinants can be influenced by the biomaterial fabrication method.²⁴

2.1. Chemical Composition

In the realm of biomaterials, various classes of synthetic polymers, including polyesters, polyanhydrides, polyphosphazenes, polyurethanes, and polyhydroxyalkanoates have garnered significant attention because of their biocompatibility and biodegradability (Table 1).²⁵ The degradation of these polymers in vivo is driven by chemical hydrolysis or enzymatic processes targeting their polymer backbones. In hydrolysis, water penetrates the polymer matrix and cleaves the hydrolysable bonds breaking the long polymer chains into smaller water-soluble oligomers that diffuse out of the matrix.²³ Synthetic polymers can be specifically designed for accelerated degradation using enzymatic pathways, where enzymes such as lipases, esterases, and proteinases are adsorbed onto the polymer surface to catalyze the hydrolysis of the polymer backbone. Following hydrolysis, the resulting oligomers are naturally removed by the organism itself. For example, in an in vivo environment, PLGA oligomers are further decomposed into lactic and glycolic acid monomers and metabolized through the citric acid cycle or the Cori cycle. The citric acid cycle is a metabolic pathway that takes place in the mitochondria of most aerobic cells and is involved in converting lactic acid into ATP, which serves as an energy source for various tissues.^65,66 The Cori cycle, also known as the lactic acid cycle, is a metabolic pathway in the liver that converts lactic acid back into glucose and releases it into the bloodstream as an energy source.⁶⁷

Table 1:

Comparison of the different synthetic polymer classes used in tissue engineering. Poly(sebacic anhydride) (PSA), Poly-[bis (p-carboxy-phenoxy) propane anhydride] (PCPP), Poly(terephthalic acid anhydride) (PTA), Poly(ester urethane)urea (PEUU), Poly(ether ester urethane)urea (PEEUU), poly[(methylphenoxy)(ethyl glycinato) phosphazene] (PMEGP), poly[(ethylalanato)1.4(imidazolyl)0.6phosphazene] (PEIP), Poly(organophosphazene)-doxorubicin (DOX), Poly[(ethyl alanato)(1)(p-methyl phenoxy)(1)] phosphazene (PNEA-mPh), poly(3-hydroxybutyrate) (P3HB), poly(4-hydroxybutyrate) (P4HB), poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx), poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV).

Polymer	Structure	Examples	Applications	Hydrolytic Degradation Rate
Polyester		PLLA7 PGA15 PLGA9 PCL26	Musculoskeletal7–9,27 Cardiovascular28 Neural26 Cutaneous29 Drug delivery30 Suture31	PLA Suture: 28–70 days32,33 PLA Films: >80% in 25 weeks34 PCL Films: >80% in over 2 years35
Polyanhydride		PSA36 Maleic anhydride37 PCCP38 PTA38	Drug delivery38 Bone37	Drug carrier: 30% in 8–30 hours39 Implants: 80% in 5–10 days40
Polyurethane		PEUU41,42 PEEUU42	Cardiovascular41 Neural43 Sutures44 Drug delivery45	Film: 20% in 8 weeks42,46 Porous Scaffold: 20 – 40% in 8 weeks42 Implants: 80% in 3 months47
Polyphosphazene		PMEGP48 PEIP49 PNEA-mPh50 DOX51	Musculoskeletal48,50 Nerve49 Drug delivery51	Films: 60% in 1–11 weeks52,53
Polyhydroxyalkanoate		P3HB54,55 P4HB56,55 PHBHHx57 PHBV57	Neural54 Cardiovascular58 Musculoskeletal55,59,60 Sutures57 Drug delivery61 Cutaneous62	P3HB Films: over 2 years P4HB Films: 80% in 8–52 weeks35,63 P4HB Fibers: 3% in 30 days64

Degradation can be categorized into two main types: bulk and surface degradation (Figure 1). Bulk degradation is a process where the polymer degrades uniformly throughout its entire volume. In contrast, surface degradation primarily involves polymer degradation on the surface while the bulk remains intact. Polymers experience both bulk and surface degradation, but typically one degradation process dominates over the other, depending on the rate of fluid flow within the bulk material. There are two kinetic processes contribute to degradation: 1) Hydrolytic degradation via chain scission, and 2) diffusion of water into the bulk of the scaffold.²⁵ The rates of these two processes determine the dominant degradation type and are the primarily determined by the chemical composition of the polymer.

Figure 1:

Process of surface and bulk degradation at the macroscopic and molecular scale at three stages of increasing degradation time. Fibers in surface degradation will result with the thinning of fibers due to chain scission on the molecular chains at the surface of the scaffold. Fibers in bulk degradation will experience chain scission throughout the bulk of the scaffold resulting in the deterioration of the entire volume of the scaffold.

The type of polymer backbone determines the reactivity of the polymer with the kinetics of hydrolysis.⁶⁸ For instance, poly(anhydrides) consist of relatively unstable anhydride linkages compared to the more stable ester bonds present in poly(caprolactones) and other poly(esters). As a result, poly(anhydrides) degrade significantly faster (50% mass loss after 1 to 5 days depending on molecular weight and environment).^69–71 Moreover, water uptake of a polymer is associated with the chemical composition of the polymer which determines if the polymer is primarily hydrophilic or hydrophobic. Water uptake is the ability for a material to absorb water from its surroundings and is influenced by the hydrophilicity,⁷² and it is strongly correlated with the organization of crystalline domains within the polymer,⁷³ and with the kinetics of water movement within the scaffold. For example, the increased water uptake in PLLA compared to PCL is correlated with more amorphous domains consisting of entangled and disordered polymer chains in PLLA compared to PCL.^73,74 As a result, more rapid degradation in mechanical properties is observed in PLLA compared to PCL.

Structural classifications play a significant role in influencing degradation rates. In this regard, aromatic ring structures are typically more resistant to hydrolysis compared to aliphatic structures composed of linear or branched carbon chains.⁷⁵ One example of an aromatic polymer explored in the field of bone tissue engineering is polyetheretherketone, valued for its impressive mechanical properties. However, it exhibits minimal degradation over extended periods, which is undesirable in some situations. To address this limitation, it is often blended with polyglycolic acid (PGA) to promote a controlled and rapid degradation profile, while maintaining many of its mechanical properties.⁷⁶

Crosslinking is a widely investigated technique that slows the degradation of polymers. During crosslinking, the introduction of covalent bonds between polymer chains effectively restricts the entry of water, leading to a considerable reduction in water uptake and an increase in polymer molecular weight. In the context of tissue engineering, crosslinked polyurethane extends degradation times and improves mechanical properties, showcasing its potential for controlled degradation in various applications.^77,78 In conclusion, the degradation behavior of synthetic polymers is influenced by factors such as chemical composition, structural classification, water uptake, and techniques like crosslinking. Understanding these factors enables the design of biomaterials with tailored degradation profiles to meet specific tissue engineering needs.

2.2. Molecular Weight and Crystallinity

Molecular weight is another parameter that influences degradation rate. The greater the molecular weight, the more chain scissions need to take place to create smaller water-soluble oligomers. Furthermore, higher molecular weight molecules exhibit reduced chain mobility which reduces water diffusion.⁷⁹ Crystallinity plays a major role in changing the accessibility of the polymer.^80,81 Amorphous regions are more accessible for diffusion of water compared to crystalline regions due to their disordered structure, in which polymer chains are arranged randomly, creating gaps and voids. Polymer crystallinity can be adjusted by blending copolymers that combine two or more monomers in order to improve functionality and degradability.^16,73 PLGA, a copolymer derived from PLLA and PGA, has been extensively investigated in tissue engineering. Copolymerization of PLLA with PGA at a ratio of 50:50 composition led to faster degradation rates compared to a ratio of 95:5.⁸² This is because PGA is a highly crystalline polymer compared to PLLA.⁸³ Introducing PLLA reduces the degree of crystallinity which leads to increased rates of hydration and hydrolysis.⁸⁴ Another factor that can influence polymer crystallinity is the scaffold fabrication technique and its associated parameters. A study found that PLLA fiber scaffolds fabricated using 3DMB technology exhibited higher crystallinity compared to those made using the standard meltblowing technique.⁸⁵ This difference is likely due to variations in molecular realignment during the two processes. 3DMB produces fiber scaffolds with greater alignment compared to standard meltblowing. This process results in fibers experiencing greater shear forces, which promote more extensive polymer chain realignment.^86–88 Furthermore, another study observed faster polymer crystallization at lower extruder temperatures, further highlighting the sensitivity of polymer crystallization to fabrication parameters.⁸⁸ Together, molecular weight, crystallinity, and copolymer composition collectively govern polymer degradation rates, offering multiple avenues to tailor scaffold properties for specific tissue engineering applications.

2.3. Surface to volume ratio of scaffolds

The surface to volume ratio (SVR) affects the degradation rate. In tissue engineering, a higher SVR generally correlates with a faster degradation rate by enhancing the exposure between the scaffold and its surrounding environment, which facilitates hydrolysis and enzymatic degradation.⁸⁹ However, this relationship can be misleading, as degradation is also heavily influenced by scaffold design. One study found that, in porous PLGA scaffolds, those with a low SVR ratio degraded more rapidly compared to those with a high SVR ratio.⁸ The accelerated degradation was not driven by the water-polymer interactions but rather by the accumulation of acidic degradation byproducts within the scaffold, creating an autocatalytic environment. Fiber parameters such as porosity and diameter, associated with SVR ratio, also impact degradation rates where PLGA scaffolds exhibit increased weight loss with larger fiber diameters (4–6 mm) compared to smaller diameters (1.1–1.3 mm).⁹⁰ Therefore, the SVR ratio is a key factor in designing drug release profiles, especially when developing complex geometries through 3D printing.^91,92

2.4. pH of Surrounding Environment

The pH of the surrounding environment also affects degradation rates through catalysis of the degradation reaction. Specifically, aliphatic polyesters can produce acidic degradation products which act as a catalyst (Figure 3). Submerging PLA in PBS resulted in a significant pH drop from 7.4 to 2.1 over a span of 60 days, and this decline in pH accelerated the decrease in its molecular weight.⁹³ This self-perpetuating process is known as autocatalysis, where acidic pH leads to protonation of the polymer, which increases its reactivity with water. In contrast, PLLA and PGA may be more susceptible to deprotonation due to their observed accelerated degradation in alkaline conditions.^94,95 Furthermore, autocatalysis can occur in polymers if there is pathologic accumulation of acidic degradation products that lower the pH of the surrounding environment.^96,97 Further details about the impact of the acidic environment on cell-seeded scaffolds and on scaffolds in a cellular environment are discussed in Section 3.6.

Figure 3:

Schematic representation of (A) acid and (B) alkaline hydrolysis processes for PLA. The acid catalysis process generates an acidic byproduct, while the alkaline catalysis process produces an alkaline byproduct. In both cases, the byproduct can further accelerate the reaction through autocatalysis, creating a self-sustaining cycle that increases degradation process of PLA. Reprinted from McKeown et al.⁹⁸ with permission from Sustainable Chemistry

2.5. Degradation Model

All of these critical factors are incorporated into an erosion model by von Burkesroda et al.⁹⁹ that predicts whether surface or bulk erosion will dominate. Understanding whether a polymer undergoes surface or bulk erosion is crucial for determining which polymer to use for a specific application. For example, prediction of the dominant degradation mechanism is critical in the development of controlling drug release strategies. Surface erosion may lead to a more controlled and sustained drug release, while bulk erosion may result in a rapid burst release of the drug. The erosion model quantifies an erosion number (ε), defined as the ratio of the rate of water diffusion (expressed by the equation in the numerator) to the velocity of degradation (expressed by the equation in the denominator), as follows:

$$\epsilon =\frac{\frac{1}{4D_{\mathit{\text{eff}}}}{〈x〉}^2\pi }{\frac{1}{\lambda }\left(\text{In}[〈x〉]-\text{In}\left[\frac{\sqrt[3]{M_n}}{N_A(N-1)\rho }\right]\right)}$$

Where x is the distance water needs to travel in the function matrix, M_n is the average molecular weight, N is the average degree of polymerization which is the number of monomers per polymer chain, N_A is Avogadro’s number, ρ is the density of the polymer, λ is the degradation rate constant of the polymer bonds, D_eff is the effective diffusion coefficient of water inside a polymer, and ε is the dimensionless erosion number.

For polymer matrices, ε falls into three categories. If ε >1, bond hydrolysis occurs more rapidly than water diffusion, thus surface degradation is expected to occur. If ε <1, the rate of water diffusion exceeds the rate of bond hydrolysis thus polymers will undergo bulk degradation. If ε = 1, both surface and bulk degradation takes place at similar rates, and the dominant process cannot be predicted using the model. Although this erosion model can identify the dominant erosion mechanism, it comes with several limitations. Specifically, the model exclusively considers hydrolysis and does not consider other degradation mechanisms such as enzymatic degradation. Additionally, this model relies on simplified assumptions, such as uniform degradation velocity within the polymer matrix, and structural characteristics such as crystallinity are not incorporated. Moreover, external factors like mechanical stimulation and fluid flow were not considered. Nonetheless, this model is useful in the early development of novel drug delivery strategies but does not replace the need for experiments specifically designed to evaluate degradation in individual biomaterials.

2.6. External Factors That Impact Degradation Rate

The degradation rate of scaffolds in tissue engineering is influenced by a variety of external factors, including mechanical stimulation, environmental conditions, and biologic responses (Table 2). Understanding these factors is crucial for optimizing scaffold design to ensure that degradation aligns with tissue regeneration.

Table 2:

Summary of external factors that impact scaffold degradation rate. ↑ represents increase in that specific external factor.

Factors	Impact	Higher degradation rate when…
Mechanical stimulation	Formation of micro fractures leading to increased SVR ratio	↑ Frequency / Number of cycles ↑ Stretch
Fluid flow	Shear force induced degradation	↑ Fluid flow
pH	Catalyzes breakdown of polymer chains	↑ Acidic/alkaline environment
Cell metabolism	Promote enzyme activity	↑ Cell metabolism ↑ Enzymatic activity
Inflammation	Promote macrophage activity	↑ Macrophage mediated degradation

2.6.1. Dynamic Mechanical Stimulation

Dynamic mechanical stimulation impacts degradation rates in vitro, but unpredictable and depend on interactions between multiple other factors. One study observed negligible differences in mechanical strength, average molecular weight, and mass between static and compressive loading in phase separated PLGA scaffolds.¹⁰⁰ However, this was due to static conditions having an increased accumulation of acidic degradation products, so it was believed that the dynamic loading alleviates the acidity by creating fluid flow in the surrounding environment. Supporting this, the average molecular weight and mass of solvent-cast, particulate-leached PLLA scaffolds decreased more rapidly under dynamic compressive loading than under static loading in a flow chamber, where the acidic environment was effectively removed.¹⁰¹ However, this effect may be more relevant for scaffolds with relatively low porosity. When compared to highly porous 3D printed fiber scaffolds, dynamically loaded scaffolds exhibited greater degradation than those under static conditions.¹⁰² This increased degradation was likely due to the formation of microcracks caused by dynamic loading.¹⁰³ Finally, fluid flow impacts degradation rates by inducing shear stress. Increasing fluid flow leads to faster mass and molecular weight loss in PLGA films.¹⁰⁴ Increasing fluid flow enhanced transport of water into the polymer matrix and directly increased shear forces which induced the breakage of polymer chains on the surface. In tissue engineering, these considerations are essential as the site of implantation is a dynamic environment. In the case of SVR ratios, the addition of fluid flow or mechanical stimuli may flush out the acidic products and improve degradation properties in scaffolds with low compared to high SVR ratios.

2.6.2. Cell Metabolism and Behavior

Cell metabolism and behavior also impact degradation of the scaffold. Cells secrete enzymes such as matrix metalloproteinases (MMPs) which play an essential role in ECM homeostasis, remodeling, and degradation, enhancing the ability of cells to migrate in the 3D environment.¹⁰⁵ Interestingly, cells dynamically adjust their secretion of MMPs and tissue inhibitors of metalloproteinases (TIMPs) in response to the stiffness of their physical microenvironment. For example, studies have shown that human mesenchymal stem cells (hMSCs) seeded in poly(ethylene glycol) hydrogels modulate their MMP secretion based on the stiffness of the surrounding material.^106,107 This adaptive response helps reduce physical barriers of the cells by increasing the degradation rate of the scaffold, facilitating improved cell migration and tissue integration. Another enzyme known to degrade polyesters is lipase, a non-specific esterase secreted by the pancreas. Lipase plays a key role in the resorption of PHB after its implantation in the human body, and as a result, PHB fiber scaffolds have been shown to undergo a significant reduction in fiber diameter over a period of 6 months in vitro in the presence of lipase.¹⁰⁸ Additionally, cells can also generate acidic byproducts including lactic acid during anaerobic glycolysis and carbonic acid resulting from carbon dioxide production during metabolism. These acidic byproducts may contribute to accelerated scaffold degradation. Interactions between cell to material can further impact degradation through generation of mechanical stresses on the scaffold during cell adhesion, migration, and spreading.^109,110 In summary, cell-mediated processes such as enzyme secretion, metabolic byproduct generation, and mechanical interactions play a pivotal role in scaffold degradation, highlighting the dynamic interplay between cellular activity and material properties in tissue engineering applications.

3. Impact of Scaffold Degradation on Scaffold Properties

With multiple properties of the scaffold being affected by degradation, for each potential application, it is critical to understand how parameters such as fiber, mechanical, and molecular properties are influenced by degradation (Figure 4),¹¹¹ since these scaffold parameters in turn influence cell behavior, proliferation, and differentiation. For instance, slower degrading scaffolds are needed for massive rotator cuff tendon tears compared to small rotator cuff tendon tears to provide extended mechanical support for the rotator cuff tendon.¹¹² Mismatching slow degrading scaffolds for small rotator cuff tendons tears could lead to stress shielding, resulting in less robust rotator cuff tendon during remodeling.¹¹³ Instead, the progressive loss of structural integrity and mechanical properties of the scaffold that occurs during degradation should be promptly compensated by the progressive increase in function of newly formed tissue. Therefore, the degradation properties of a scaffold are of critical importance for understanding biomaterial selection and design for long-term success in a tissue engineering application.

Figure 4:

Schematic diagram of the morphological and structural changes of fiber scaffolds at early and late stages of hydrolytic degradation at both the (A) scaffold level, and (B) fiber level. The early stage, scaffolds (A) experience fiber swelling, fiber crimping, and scaffold shrinkage. This leads to increased fiber diameter, decreased alignment, and decreased pore size. At the late stage, scaffolds experience surface degradation, fiber breakage, and continued shrinkage. This leads to decreased fiber diameter, decreased alignment, and decreased pore size. Within the fibers (B), amorphous regions are more susceptible to hydrolytic degradation compared to crystalline regions, and the cleavage of amorphous regions are followed by the rearrangement of the polymer fibers resulting with the disappearance of amorphous gaps and the collapse of crystalline regions. This leads to increased crystallinity during degradation.

3.1. Impact of Degradation on Fiber Diameter

Fiber diameter is the thickness or width of fibers used in tissue engineering scaffolds. The diameter of the fibers typically ranges from nano (<1μm) to micro (>1μm) scale depending on application, and significantly impacts cellular behavior and tissue formation within the engineered construct. Neural progenitor cells increased oligodendrocyte differentiation on 283 nm fibers but increased neuronal differentiation on 749nm fibers.¹¹⁴ In the same study, neuronal progenitor cells stretched multi-directionally on 283 nm fibers but extended along a single fiber axis on larger fibers > 750nm.¹¹⁴ In other studies, vascular smooth muscle cells had significantly higher cell viability after 12 days of culture on 0.75μm diameter fiber scaffolds compared to 2, 4, and 6μm diameter fiber scaffolds.¹¹⁵

Degradation of synthetic polymers undeniably exerts a significant influence on fiber diameter, and this is governed by various critical factors. Degradation of fiber diameter initiates by surface degradation of the polymer which causes a decrease in fiber diameter.^116–119 However, the impact on fiber diameter is initially not as pronounced due to the absorption of water in synthetic polymer scaffolds which causes swelling of the scaffold fibers.¹²⁰ In this regard, the fiber diameter of different PLGA blends was either increased or unchanged over a 12-week incubation period in PBS.¹²¹ Another study that incubated PLGA in simulated body fluid, DMEM, and artificial saliva observed fiber diameter swelling after two weeks at varying degrees depending on solution.¹¹⁶ After initial scaffold hydration however, the fiber diameter will ultimately show decrease.^116–119 Finally, additional parameters that impact fiber diameter degradation are SVR ratio and polymer composition.¹¹⁹ SVR ratio can be changed by adjusting the fiber diameter size and porosity of the polymer. This influences the rate of water absorption and increasing fiber diameter slows degradation rates. Together, these data suggest that fiber diameter initially increases at the early stages of exposure to physiological conditions due to swelling but then diameter gradually decreases due to degradation.

3.2. Impact of Degradation on Fiber Alignment

Fiber alignment refers to the orientation and arrangement of fibers in a scaffold used for tissue engineering. It involves aligning the fibers in a specific direction to mimic the natural architecture of the tissue of interest to promote the desired cellular behavior and tissue development. Based on the application, fiber alignment ranges from random to multiaxial to uniaxial alignment. Vascular tissue engineered scaffolds uses highly aligned fibers to match the arrangement of fiber bundles in blood vessels.^110,122 On the other hand, musculoskeletal tissues such as tendons may prefer to use a gradient of aligned fibers to recapitulate the microarchitecture of tendon, combined with unaligned fibers to recreate the multiaxial or non-aligned microarchitecture of the fibrocartilage at the tendon-to-bone interface.^123,124

In aqueous environments, fiber alignment is initially impacted by hydration which increases the tortuosity of the fibers due to scaffold macroscale shrinkage and fiber swelling.¹²⁵ The level of deformation depends on the polymer where a macroscale size reduction of between 55–80% was observed in PLGA (75:25),^126,127 ~3% was observed in PCL, and ~30% was observed in collagen.¹²⁵ This fiber buckling may be due to constraints on the ability of fibers to freely expand during hydration resulting in an overall decrease in fiber alignment. Over time, polymer fibers progressively shrink as degradation occurs due to hydrolysis-induced chain scission of stretched amorphous chains. This process, known as cleavage-induced crystallization, leads to the release of internal stresses and the collapse of crystalline lamellar stacks.^64,108,127 Several methods have been explored to prevent significant scaffold shrinkage. One approach involves introducing a polymer with a highly crystalline chemical structure. In the case of Poly D,L-lactic acid (PDLLA) macroscale shrinkage of 82% occurs, but in contrast the more crystalline PLLA structure shrunk by only 9%. This substantial difference was attributed to the amorphous nature of PDLLA, primarily due to the presence of D-lactic acid units in its structure, which disrupts the uniform and organized polymer chain arrangement.¹²⁷ Another method of reducing shrinkage is by thermal annealing which further decreased the shrinkage of PLLA to 0.¹²⁷ By subjecting the material to temperatures near its glass transition temperature, thermal annealing facilitates controlled and gradual relief of internal stresses within the material. This process leads to a more stable and improved material microstructure. As degradation progresses, the alignment of fibers is also compromised by fiber breakage, which disrupts the previously continuous fiber structure and results in gaps or breaks in the once-aligned pattern.⁸¹ Consequently, any scaffold, polymer, or environmental/cellular properties that enhance the degradation rate will inevitably accelerate the disruption of fiber alignment. These findings indicate that fiber alignment initially decreases due to fiber swelling and crimping induced by scaffold shrinkage. In the later stages, alignment continues to decrease due to ongoing shrinkage and fiber degradation.

3.3. Impact of Degradation on Fiber Pore Size

Fiber porosity refers to the presence of interconnected void spaces or pores within a scaffold. It is a measurement of the open space available within the structure, and adequate porosity is important for cell migration and infiltration, and to allow for ECM deposition throughout the scaffold. Both fiber diameter and alignment influence porosity. Larger fiber diameter results in larger pore sizes and lower porosity.^115,128 This is due to reduced fiber density and packing efficiency because there is less space available for a given volume of fibers. Random fiber alignment was computationally¹²⁹ and experimentally¹³⁰ tested and resulted in smaller pore sizes compared to aligned fibers. In biological applications, optimal pore size varies for different cell types.¹³¹ A study by Han et al. testing three different cell lineages on scaffolds with different pore sizes observed greatest cell viability on 200μm pore size scaffolds for bone marrow mesenchymal stem cells, 100 and 200μm pore size scaffolds for chondrocytes, and 300μm size scaffolds for tendon stem cells.¹⁰⁹

Pore size is linked to the packing density of fibers, with increased packing density associated with smaller pore size. Consequently, pore size consistently decreases during degradation due to scaffold shrinkage increasing the packing density of fibers.^121,132 Interestingly, fiber diameter impacts fiber packing density with increased fiber diameter resulting in reduced packing density.¹³³ However, despite the degradation process decreasing fiber diameter, there is no observed correlation between fiber diameter and pore size changes during degradation.¹²¹ In relation to degradation rate, pore size also influences the degradation rate of scaffolds. Smaller pore sizes and lower scaffold porosity are associated with decreased degradation rates.^134,135 This phenomenon primarily stems from the surface-to-volume ratio, which limits the diffusion of water into the scaffold. Therefore, in the initial stages, the pore size decreases as a result of scaffold swelling. As the process progresses, the pore size continues to diminish due to scaffold shrinkage.

3.4. Impact of Degradation on Mechanical properties

Mechanical properties play a pivotal role in guiding cellular responses. Factors like substrate stiffness and modulus significantly influence cell behavior. For instance, when examining cell morphology, mesenchymal stem cells (MSCs) exhibit reduced spreading, stress fibers, and proliferation on soft substrates compared to their behavior on hard substrates.¹³⁶ Furthermore, cell differentiation can be directed by these mechanical cues. Engler et al. demonstrated that MSCs can be induced into neuronal, muscle, or bone lineages depending on whether they are cultured on soft, medium, or stiff hydrogel matrices.¹³⁷ In terms of cell migration, epithelial and fibroblastic cells increase motility on flexible substrates in comparison to rigid ones.¹³⁸ Notably, while the mechanical properties of individual fibers are generally intrinsic to the polymer used, as expected, altering the polymer composition such as molecular weight can influence these mechanical properties.¹³⁹

The degradation process of scaffolds significantly impacts their mechanical properties. During the initial stages of degradation of PLLA, PCL, and 50:50 PCL:PLLA blend, modulus and ultimate tensile strength initially increase, accompanied by a decrease in yield strain. This change is associated with heightened crystallinity of the polymers due to cleavage-induced crystallization resulting in a more brittle material.¹¹⁷ Increased crystallinity leads to a more ordered atomic structure, stronger bonds, reduced molecular mobility, and improved load-bearing capacity, all contributing to enhanced stiffness.¹⁴⁰ Further into degradation, the mechanical properties decrease due to erosion of the overall structure of the scaffold. In a polyesteretherurethane implant mesh, the yield strain and stress decreased significantly over the course of 6 weeks in PBS primarily due to fiber breakage.¹⁴¹ Additionally, PHB fiber scaffolds significantly decreased in yield strain and stress after 30 days in lipase containing PBS.⁶⁴

The mechanical properties of a scaffold are significantly influenced by the properties of its constituent fibers. Fiber diameter exhibits an inverse relationship with the scaffold’s modulus.¹⁴² The alignment of fibers also plays a crucial role, showing an asymmetrical U-shaped relationship.^143,144 Moreover, porosity demonstrates an inverse relationship with the modulus.¹⁴⁵ Thus, it is important to emphasize that any degradation in these fiber properties directly impacts the overall mechanical properties of the scaffold. Furthermore, dynamic tensile forces can accelerate the degradation process.¹⁴⁶ This acceleration is attributed to greater diffusion, increased shear stress, and friction. Additionally, polymers under tensile stresses may exhibit creep properties, further contributing to the degradation.¹⁴⁷ Collectively, these factors highlight the complex interplay between fiber properties, degradation, and mechanical performance in scaffolds. Therefore, during the early stages of degradation, the mechanical stiffness of the scaffold increases as a result of cleavage-induced crystallization. However, as degradation progresses, the mechanical properties of the scaffold decline due to fiber degeneration and breakage.

3.5. Impact of Degradation on Surface properties

Surface properties of a polymer play a pivotal role in cellular to material interactions,¹⁴⁸ by modifying cell adhesion and communication. In tissue engineering, cell adhesion is largely determined by surface wettability rather than surface functional groups, surface density, or cell type.¹⁴⁹ The cellular adhesion protein, fibronectin, prefers to adhere to moderately hydrophilic surfaces (water contact angles ranging from 20° to 70° depending on cell type),^149,150 but not super hydrophilic surfaces (less than 5°). The hydrophilicity of the polymer depends on its chemical, and synthetic scaffolds generally exhibit hydrophobic properties. PCL has a non-polar aliphatic backbone consisting of repeating units of six carbon atoms. In contrast, PLLA has a more polar structure due to the presence of shorter repeating carbon chains in its backbone. This polar nature makes PLLA more hydrophilic compared to PCL. PLLA surfaces have an observed water contact angle of 75°−85°¹⁵¹ whereas PCL surfaces have a water contact angle of 120° to 140°.¹⁵² In response to hydrophobic surfaces, methods of modifying polymer surface to improve wettability of polymers have been explored. A common method to enhance wettability of scaffolds is by ethanol pretreatment.^7,153 Ethanol pretreatment of PGA/PCL composite scaffolds enhanced chondrocyte development and surface adhesion.¹⁵³ Another method used to improve wettability is plasma treatment.¹⁵⁴ Etching plasma patterns on hydrophobic surfaces leads to the continued growth of cells on top of each other on the plasma patterns, instead of expanding their growth across hydrophobic areas.¹⁵⁵

Surface topography can also modulate the activity of cells interacting with an implant. Surface roughness of a material can influence its hydrophilicity. Increasing roughness raises the surface energy of the scaffold, which is associated with its wettability. The roughness also allows for greater effective surface area for improved mechanical attachment of cells. However, there is a critical limit in roughness for good cellular growth and proliferation.^156,157 At high roughness ratio, the elastic energy of a cell hinders its ability to completely coat the surface grooves by satisfying Cassie-Baxter state. In this state, cells sit over the tips of the rough structures leading to only point-contact with the cell which significantly reduces cell to substrate interaction. This critical limit is also dependent on the elasticity of the cell.

How degradation impacts surface roughness of the scaffold is currently an underexplored area. Initial evidence suggests that degradation significantly influences surface properties of scaffolds. The formation of micropores was observed on polyurethane films after 21 days in culture due to degradation.¹⁵⁸ Poly(sebacic anhydride) films developed significant surface roughness over the course of 48 hours.¹⁵⁹ Furthermore, longitudinal cracks were observed in poly(lactic-co-ε-caprolactone) fiber scaffolds after 168 days in culture medium.¹⁶⁰ Over time, the degradation process will increase surface area of the scaffold and accelerate degradation rate. However, cell to material interaction remains unknown as there are currently no studies that investigate how degradation impacts the hydrophilicity and roughness ratios of the scaffold surface. Thus, as degradation proceeds, surface roughness of the scaffold will increase, potentially altering cellular adhesion, proliferation, and overall tissue integration, which underscores the need for further research in this critical area.

3.6. Scaffold Degradation Products

During the process of hydrolysis in polymer scaffolds, degradation byproducts are generated. Aliphatic polyesters produce acidic byproducts. These byproducts create an autocatalytic environment that expedites degradation. However, they also have the potential to interact with the surrounding environment and cells, leading to modifications in their behavior. One notable effect is the significant impact that pH alteration in the surrounding medium can have on cells. For instance, the degradation of PLGA over a 28-day period substantially decreased the pH, resulting in reduced cell viability, decreased cell mobilization, and angiogenesis in mouse aortic smooth muscle cells.¹⁶¹ Moreover, reducing the pH to 6.9 was found to inhibit bone formation in osteoblasts by impeding mineralization.¹⁶² When cultured with poly(D,L-lactic acid) (PDLLA), osteoblasts increased alkaline phosphatase (ALP) activity in response to the acidic products.^158,162 ALP activity increases the pH of the immediate vicinity of osteoblasts due to the release of hydroxide ions during the dephosphorylation process.¹⁶³ As such, this process improves the interfacial pH and creates an environment conducive to their optimal survival. Osteoblasts tend to thrive in alkaline conditions, prompting efforts to alkalize the surface pH of biomaterials in the field of bone tissue engineering.¹⁶⁴ Another noteworthy consequence of pH fluctuations is their potential to trigger an immune response. Acidic environments induce the production of inflammatory mediators (IL-1ß, COX-2 and iNOS), altering the phenotype of macrophages towards a pro-inflammatory, phagocytic activity.¹⁶⁵ Additionally, PLGA degradation initiates a cascade of inflammatory reactions that shift macrophages towards a pro-inflammatory (M1) polarization state.¹⁶⁶ Due to the influence of acidic environments on cell behavior, approaches to mitigate detrimental pH alterations are under evaluation. In one study, nanophase titania was incorporated into PLGA to create a pH-buffering effect on the polymer surface. This incorporation effectively prevented the pH from decreasing, raising it from 2.2 in pure PLGA to 4.2 when using a blend consisting of 70% nanophase titania and 30% PLGA.¹⁶⁷

Degradation products yielding ketones, as seen in polymers like PCL and polyhydroxybutyrate (PHB), at physiological levels, potentially offer a valuable alternative energy source through ketogenesis particularly for the brain when glucose levels are depleted within the body.¹⁶⁸ Therefore, these polymers, producing ketone degradation products such as PHB, are commonly preferred choices for nerve tissue engineering due to their positive metabolic effects.¹⁶⁹ Moreover, ketones have demonstrated potential anti-inflammatory properties by exerting antioxidant effects within macrophages. In a noteworthy study, PHB degradation products upregulated anti-inflammatory genes (NF-κBIA, MAP3K8, and TLR5) while concurrently downregulating pro-inflammatory genes (TNFSF6, TNF-α, PI3K, NF-κB, and TLR1), thereby enhancing the immune response.¹⁷⁰

Lactic acid, a degradation byproduct of PLA and PLGA, shares similarities with ketones in that it can act as an alternative energy source for the body and possesses antioxidant properties. Therefore, it is metabolized as a harmless byproduct. Lactic acid can be utilized as an energy source in the Cori cycle.^65,66 Alternatively, it can undergo oxidation to become pyruvate which is then oxidized to acetyl-CoA, serving as fuel for the citric acid cycle within the mitochondria.⁶⁷ Additionally, beyond these roles, lactic acid may play a protective role in shielding cells from damage caused by naturally occurring free radicals. When oxidized, lactic acid generates pyruvate, a molecule able to scavenge hydrogen peroxide and superoxide radicals.¹⁷¹ Given these effects, exposure to lactic acid at concentrations ranging from 0.005 to 0.5 mg/mL over a period of five days increases the total DNA content and the expression of nestin, a marker associated with neurogenesis, in neural cells.⁶⁷ During this study, media was replaced after 2–3 days, so the impact of acidic environment on cell behavior was likely not observed.

Although PHB degrades into 3-hydroxybutyric acid, a natural metabolite in the human body, its degradation does not alter the local pH, making it highly biocompatible. This stability is due to its very low degradability.¹⁷² In a 30-day comparison of mass loss in the presence of lipase, PHB fiber scaffolds lost only about 3% of its mass, while PCL fiber scaffolds lost approximately 14%.⁶⁴ Because of this slow degradation rate, PHB is particularly well-suited for applications in bone tissue engineering.

When considered together, cellular activity throughout the body drives a continuous cycle of scaffold degradation and cellular adaptation until the implanted scaffold is fully absorbed into the surrounding tissue (Figure 5).

Figure 5:

Diagram illustrating the biodegradation cycle of synthetic scaffolds in vivo. Synthetic scaffolds degrade through hydrolysis, enzymatic activity, or macrophage-mediated degradation. The degradation products undergo metabolic processes including gluconeogenesis (Cori cycle), oxidation to pyruvate (Krebs cycle), and ketogenesis. These pathways then serve as alternative energy sources for cells which further contribute to scaffold degradation.

4. Preventative Measures to Slow Degradation Rate

The degradation rate of synthetic scaffolds is a critical factor in tissue engineering, influencing both the mechanical properties of the scaffold and the regenerative capacity of the surrounding tissues. Various preventative measures can be employed to modulate this degradation rate, ensuring that it aligns with the rate of tissue regeneration and lessen detrimental impacts on fiber scaffold properties (Table 3).

Table 3:

Summary of scaffold modifications and how they impact scaffold degradation. ↑ represents increase and ↓ represents decrease in the respective scaffold property.

Property	Impact	Lower degradation rate when…
Molecular weight	Determines the number of molecular chains that need to be broken	↑ Molecular weight
Fabrication parameters and scaffold geometry	Alters scaffold properties tied to degradation (e.g., crystallinity and porosity)	↑ Crystallinity ↓ Porosity ↓ SVR Ratio
Polymer blending	Combines degradation properties of two or more distinct polymers.	Depends on blended polymers
Treatment (e.g., annealing and crosslinking)	Enhances intermolecular bonds and polymer stability	↑ Crosslinking ↑ Annealing
Surface modifications and coatings	Modifies surface properties (e.g., water affinity and biocompatibility)	↓ Hydrophilicity ↓ Biocompatibility

4.1. Modifying scaffold raw material

One approach involves modifying the raw material used to synthesize the scaffold. For instance, certain polymers like PCL³⁵ exhibit a higher degradation rate compared to PLA.³² Additionally, polymer molecular weight plays a crucial role in determining degradation behavior. Higher molecular weight tends to decrease degradation rates.^173,174 This occurs because increased molecular weight will result in longer polymer chains, requiring a greater extent of chain cleavage for degradation to occur over time.⁹⁷ Another option is polymer blending. This process combines two or more polymers of different characteristics to create a composite material with tailored properties. One study blending polydioxanone (PDO), a fast degrading polymer, with poly(lactic-co-caprolactone) (LCL), a slower degrading polymer, observed increasing LCL content decreased hydrolytic degradation rate of the polymer blend.¹⁷⁵ Similarly, blending PGA with PLA decreased the degradation of PGA scaffolds.¹⁷⁶ Thus, blending two or more polymers together can influence the degradation of the other. In summary, scaffold degradation can be modulated through material selection such as molecular weight adjustments and polymer blending.

4.2. Adjusting scaffold fabrication parameters and design

Another method is to adjust parameters during scaffold fabrication. For instance, compared to melt processing, the high temperature, shear rate, and longer dwell times of injection molding were detrimental to the final material, resulting in a reduction in molecular weight.¹⁷⁴ Another study looking into different extruder temperatures using the meltblowing process observed changes in surface roughness and polymer crystallinity, both of which are linked to scaffold degradation.¹⁷⁷ Increasing the printing temperature and speed of 3D printed PLA scaffolds increases crystallinity, enhancing compressive strength and improving thermal resistance. These changes resulted in slower enzymatic degradation rate of the scaffolds.¹⁷⁸ Notably, crystallinity has a more pronounced impact on degradation than polymer molecular weight. For instance, PLA films with molecular weight of 3×10⁶ g/mol and crystallinity of 5% demonstrated a faster degradation rate compared to films with molecular weight of 3×10⁵ g/mol and crystallinity of 30%.¹⁷⁹ Scaffold geometry and design also significantly influence degradation rates. A study comparing various 3D-printed PLA micropatterns - including lattice, hexagonal, square wave, gyroid, and triangular designs - with a consistent porosity of 75% and a fiber diameter of 0.2 mm found that hexagonal patterns exhibited the highest degradation. This was attributed to localized acidic buildup from degraded polymer chains and fiber spacing affecting recrystallization.¹⁸⁰ Ultimately, fabrication parameters such as temperature, speed, and scaffold design significantly impact degradation behavior. Notably, crystallinity and geometry are crucial elements that affect both the rate and extent of scaffold degradation.

4.3. Scaffold Treatment Techniques

Additional techniques to slow degradation include annealing and crosslinking. Annealing is a treatment process aimed at enhancing the strength and stiffness of polymer scaffolds. For instance, the annealing process of PLGA demonstrated increased degree of crystallinity with annealing times. This correlated with decreased degradation rates. However, samples that were exposed to longer annealing processes exhibited faster degradation rates which were associated with the formation of voids during annealing. This process subsequently increased average water uptake into the sample.¹⁸¹ Annealing temperature also impacts degradation outcomes. Annealing electrospun PDO scaffolds at low temperatures (65°C) resulted in reduced rate of degradation whereas higher temperatures (75°C and 85°C) resulted in increased or negligible degradation rates compared to untreated PDO scaffolds.¹⁸² Crosslinking is a widely used technique that stabilizes the polymer by extending the network structures of polymeric chains. Polymer crosslinking reduces the degradation rate.^183–185 Finally, surface modifications can also impact degradation rate. One example that increases degradation rate is plasma treatment which enhances wettability by increasing oxygen-containing bonds and surface roughness which causes degradation rates to increase.¹⁸⁶ Another study coated the surface of PCL 3D printed scaffolds with egg shell nanoparticles (a type of bioceramic) and observed increased degradation rate.¹⁸⁷ Surface coatings, such as demineralized ECM, are commonly used to improve scaffold biocompatibility. These coatings may also promote the presence of bioactive components, like cell-produced enzymes, further accelerating degradation. Therefore, techniques such as annealing, crosslinking, and surface modifications offer versatile strategies to control scaffold degradation rates, with each method presenting unique advantages and trade-offs depending on the desired application and material properties.

4.4. Limitations of Degradation Slowing Preventative Measures

There are several limitations when applying preventative measures against scaffold degradation in biomedical applications. One concern is the change in mechanical properties of the material. Processes such as crosslinking,^183–185 annealing,¹⁸² and polymer blending¹⁷⁵ can increased mechanical properties such as stiffness. While this may be advantageous in certain applications, it could pose challenges in biomedical applications that require flexible materials such as vascular grafts, stents, and neural constructs. Additionally, increased mechanical properties combined with a slower degradation rate could cause stress shielding of the tissue due to the inappropriate material degradation alongside the regeneration of tissue.¹⁸⁸ This effect is more prominent in load bearing tissues such as bone,¹⁸⁸ tendons,¹¹³ and ligaments. Finally, the resulting modifications could necessitate implant removal in clinical applications, potentially leading to increased treatment costs and the need for removal surgery. The requirement for removal surgery could heighten the risk of complications such as tissue damage and infection.

5. Future Directions

The existing body of research reveals certain gaps that warrant additional exploration. (1) Understanding of how surface degradation influences cell behavior remains elusive. Surface topography plays an important role in cell to material interaction,¹⁴⁸ and further insight on this property will inform more effective tissue engineering strategies. (2) Assessing the impact of degradation on pH in vivo, particularly in fast-degrading polymers and their subsequent effect on the immune response, remains challenging as fluid flow can play a pivotal role in maintaining the desired pH levels in an in vivo environment. (3) Determining whether degradation and mechanical properties inherently hinder cell activity is a complex task to ascertain. In some instances, the degraded scaffold could be replaced by the production of ECM, resulting in minimal cell-scaffold interaction and the development of superior mechanical properties in the presence of polymer degradation. (4) Although the positive effects of ketones and lactic acid on cell behavior are known, the potential risks of prolonged exposure to elevated levels, or impaired clearance, should not be underestimated. These byproducts may induce cellular complications by activating harmful pathways that lead to cellular damage.¹⁸⁹ (5) Long term degradation studies should be conducted to encapsulate the evolution of scaffold properties over time. The majority of investigations consist of short-term degradation studies in vitro that may not capture the full spectrum of changes that can occur during the late stages of degradation in vivo. (6) Bridging the gap between synthetic scaffold integration and real physiological conditions remains a significant challenge. Current research often relies on static culture systems, neglecting the crucial influence of fluid flow and mechanical stimuli that occur in a physiological setting. Considering the dynamic nature of the host environment is essential for accurately modeling scaffold degradation and its impact on cell responses. Incorporating fluid flow and mechanical factors into studies will provide a more realistic picture of how scaffolds interact with surrounding tissues.

6. Conclusion:

The literature highlights how scaffold degradation has a multifaceted impact on the scaffold’s fiber, mechanical, and extracellular properties, which, in turn, play a pivotal role in signaling cues that regulate cellular processes such as differentiation, proliferation, adhesion, and migration. In this context, the literature presents a comprehensive examination of how degradation affects each property. Fiber diameter initially decreases due to surface degradation, then increases as a result of swelling, and ultimately decreases again with continued surface erosion after the fibers become saturated. Fiber alignment continues to diminish owing to cleavage-induced crystallization, which induces scaffold shrinkage. Fiber pore size also decreases progressively as a consequence of scaffold shrinkage. Mechanical properties of the fibers exhibit an increase in stiffness due to crystallization. Finally, it is essential to recognize the impact of acidic degradation products, particularly prominent in aliphatic polyesters. A deeper understanding of these effects - particularly the limited knowledge surrounding how cells respond during degradation and to strategies that modify it - is crucial for optimizing biomaterial design and performance in tissue engineering and regenerative medicine.

Figure 2:

Schematic illustrating the molecular realignment process at the (A) fiber and (B) molecular level. The fiber alignment within non-woven fiber scaffolds can be increased by modifying fabrication parameters such as air flow, collector surface speed, and temperature. These adjustments increase shear stress on the individual fibers which causes fiber stretching and increased crystallinity of the polymer due to molecular realignment induced crystallization.