Biomedical Applications of Biodegradable Polyesters

Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons by Attribution (CC-BY) license (http://creativecommons.org/licenses/by/4.0/).

Abstract

The focus in the field of biomedical engineering has shifted in recent years to biodegradable polymers and, in particular, polyesters. Dozens of polyester-based medical devices are commercially available, and every year more are introduced to the market. The mechanical performance and wide range of biodegradation properties of this class of polymers allow for high degrees of selectivity for targeted clinical applications. Recent research endeavors to expand the application of polymers have been driven by a need to target the general hydrophobic nature of polyesters and their limited cell motif sites. This review provides a comprehensive investigation into advanced strategies to modify polyesters and their clinical potential for future biomedical applications.

1. Introduction

The current market for regenerative implantation surgeries, therapeutic cell culturing and tissue repair is approximately US $23 billion, and it is anticipated to reach US $94.2 billion by the end of 2025 [1]. Synthetic biodegradable polyesters are considered the most commercially competitive polymers for these applications as they can be produced reproducibly in a cost-effective manner with a wide range of characteristics. Polyesters are also biocompatible, and biodegradable polymers are used for the manufacturing of different medical devices, such as sutures, plate, bone fixation devices, stent, screws and tissue repairs, as their physicochemical properties are suitable for a broad range of medical applications [2,3,4,5]. Polyesters are also used commercially in controlled drug delivery vehicles [6,7].

In all of the current commercial products, polyesters act as a biologically inert supporting material as a mesh or a drug-releasing vehicle. For more advanced medical and regenerative applications, polyesters are modified to tackle issues such as low cell adhesion, hydrophobicity, and inflammatory side-effects [8,9]. Consequently, the modification of polyesters has been one of the major research topics in the fields of material engineering and polymer science.

In this review, the properties of polyesters and the modification methods that have been implemented to improve some of the shortcomings of this class of polymers are discussed. Specifically, this review covers the applications and modifications of the most commonly used polyesters such as polylactic acid (PLA), poly(lactic-co-glycolic acid) (PLGA), poly(ε-caprolactone) (PCL), poly-3-hydroxybutyrate (or poly-β-hydroxybutyric acid, PHB), poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV), poly(propylene carbonate) (PPC), poly(butylene succinate) (PBS) and poly(propylene fumarate) (PPF).

2. Synthesis of Polyesters

Polyesters are produced predominantly by using random polymerization, ring opening polymerization, and the block copolymerization techniques. For instance, PCL is produced by the ring opening polymerization of the ε-caprolactone using a catalyst such as an octoate [10]. The synthesis methods have been extensively reviewed in detail by many researchers; therefore, these synthesis approaches are not discussed in detail in this review [11,12,13,14,15]. The vast majority of the polyesters are derived from carbohydrate petroleum-based sources. Therefore, in recent decades, there has been a drive to find alternative sustainable polymers. Among all the polyesters, only PPC, PHB and PLA come from renewable sources.

PPC is produced in commercial scale from the ring opening reaction between CO₂ and propylene oxide in the presence of an active catalyst such as zinc glutarate [16]. Similar ring opening polymerization mechanisms that are used to synthesise PPC and PCL are also used to synthesise PLA. The synthesis of PLA is a multi-step fermentation process starting with the biosynthesis of lactic acid. Lactic acid is then converted to its cyclic lactide foam and then polymerized via a metal catalyst [17,18].

PHB entirely is biosynthesized by an efficient fermentation process with different molecular weight (from 200 to 1500 kDa) using diazotrophic bacteria of acetobacter and Rhizobium genus [19]. PHB is primarily a product of carbon assimilation and it is employed by microorganisms as a form of energy storage molecules. The polycondensation of two molecules of acetyl-CoA leads to the formation of acetoacetyl-CoA that can be reduced to hydroxybutyric-CoA and polymerize PHB. However, the biosynthesis process of PHB is chirally selective and the resulting polymer typically has a polydispersity of around 2 or higher [20].

3. Properties of Polyesters

Linear aliphatic polyesters are mostly hydrophobic biodegradable polymers [21]. Their tunable physical and mechanical properties have extended their applications in the biomedical field [22]. It is easy to process these materials into desired structures with minimal risks of toxicity, immunogenicity, and infection. The main differentiating characteristics of polyesters are their mechanical performance and degradation behaviors that are discussed extensively as follows.

3.1. Mechanical Strength

In regenerative medicine, the mechanical property of a polymer plays a vital role in the selection of a biomaterial for any application. A robust biomaterial that does not mimic the mechanical strength of the targeted tissue interferes with the natural regeneration mechanism, and, ultimately, is a drawback for the damaged tissue repair [23]. The mechanical performance of bone, cartilage and cardiovascular tissues that are mostly treated with polyester-based implants are summarized in Table 1 . In addition, this table outlines the mechanical performance of different polyesters and some medical devices. Medical devices such as screws and meshes are designed from polymers with the ultimate elongation strength of 200 MPa to fix cortical bones with the compression strength of 100–200 MPa.

Table 1

Mechanical properties of the biodegradable polyesters and a few tissues and commercially available biomaterials.

There are numerous medical applications for polyester due to their broad range of mechanical properties. For instance, PGA has a relatively brittle structure as its ultimate strain is 30%. Therefore, PGA is not a desirable polyester for the fabrication of medical meshes as they are normally under high tensile strain. On the other hand, PPC displays a very flexible structure as its ultimate elongation at break is nearly 330%, which is at least five-fold higher than other polyesters. However, PPC may deform under elongation as this polymer displays very low tensile modulus, e.g., 22 MPa. Therefore, PPC is not a favorable candidate for the fabrication of medical screws, sutures, and meshes that are under constant tensile stress. PLGA and PLA posse significantly higher tensile modulus and strength compared to PPC. PLA displays the highest tensile stress (σ_m= 55 MPa) and favorable ultimate elongation at breakage (ε_m = 30%–240%); hence, it has been broadly used for the fabrication of devices that are under constant tensile stress and high elongation.

3.2. Degradation

An essential element in biomedical applications of polymers is the development of a temporary physical and mechanical support for the regeneration of newly formed tissues over time. Information about the degradation rate of a polymer is imperative for the design of various medical devices. For instance, a slow degradation rate of PLA provides the opportunity for the production of long-term orthopedic implants such as plates and screw [39,40,41]. However, PGA-based biomaterials are mainly used for the fabrication of sutures and drug delivery carriers due to their fast degradation [42,43]. Moreover, the rate of the degradation of polymers needs to be balanced to assure that the implanted device or the scaffold can provide the required mechanical strength for the regeneration of the newly formed tissue over time. For instance, in one case, a PLA-based implant, after an arthroscopic surgery, failed to regenerate the tissue and showed no signs of degradation, which resulted in some clinical complications for the patient [44].

The degradation is governed by different factors such as the nature of the polymer, composition, molecular weight, crystallinity, structure, thickness, surface properties and environmental conditions. The mechanical strength of a medical device or implant is also a function of degradation rate. For instance, molecular weight has a direct correlation with the rate of degradation, the higher molecular weight leads to slower degradation due to lengthy polymer chains [45]. However, the degree of crystallinity of some polyesters such as PLLA can proportionally affect the direct relationship between molecular weight and the degradation rate [46]. The indirect effect of crystallinity on the degradation rate is controversial as a few groups show that crystallinity of polyesters increases the degradation rate due to an increase in hydrophilicity [47,48]. In contrast, some groups display a slower rate with an increase in sample crystallinity [49].

The rate of degradation depends on the intrinsic chemical properties of polymers as well as the physical properties and the shape of the implant or device. The physical properties are important because the water diffusion and, consequently the hydrolysis of the polymer structures are affected by the contact surface area of the implants with the body fluids. Therefore, the degradation rates of different polyesters are reported within a range. Most of the polyesters are stable in the body for at least 12 months except PGA and its copolymer PLGA. This polymer has been copolymerized from LA and GA to acquire a relatively fast degradable polymer for medical applications. The degradation rate of PLGA can also be altered by changing the molar ratios of LA to GA. For instance, increasing the weight ratio of the GA to LA from 25:75 to 50:50 can accelerate the degradation by two-fold from 100 to 50 days.

Hydrolytic and enzymatic degradation are the primary mechanisms of degradation of polyesters through bulk- or surface degradation of implants [50]. Hydrolytic degradation has an autocatalytic nature and it proceeds through the hydrolysis of carboxylic groups of hydroxy acids [51], whereas the enzymatic degradation significantly depends on the enzyme that is responsible for the degradation of a specific molecule [52]. PCL, for instance, undergoes lipase-type enzymatic degradation in the presence of Rhizopus delemer lipase [53], Rhizopus arrhizus lipase, and Pseudomonas lipase [54]. Among these enzymes, Pseudomonas lipase significantly accelerates the process to totally degrade the highly crystalline PCL within four days [55], in contrast with hydrolytic degradation, which lasts several years. The general mechanism of degradation of polyesters is by bulk hydrolysis [56]. The presence of some enzymes may expedite the degradation of some of the polyesters. As a result of bulk degradation, there is a risk of a sudden loss in the structural stability of a polymeric structure.

It is critical to examine the biocompatibility and toxicity of any degradation product of a polymer for the design of biomedical devices. By-products of a bulk degradation of a polymer are released in the surrounding environment such as the host tissue. For instance, the release of acidic by-product from the degradation of PLA or PLGA may drop the pH of surrounding tissues and lead to cell necrosis and inflammation at the site [57,58,59]. It is therefore imperative to quantify the biodegradation products of polymers in order to study the biological behavior of the host environment upon the degradation of polymers systematically. The average logarithmic acid dissociation constant, pK_a, of the intermediate degradation products of polyesters is used to quantify the acidity of the resulting products upon their degradation. The pK_a of the degradation products, the primary mechanisms of the degradation, and the in vivo degradation rate of the different polymers are summarized in Table 2 .

Table 2

The degradation behavior of the biodegradable polyesters.

Most of the polyesters, except PLA, PLGA, and PGA display a pK_a of 4–5, which is considered a relatively weak acidic environment, thus, the resulting biological inflammatory responses might not be severe. For instance, the haematoxylin and eosin staining results as displayed in Figure 1 shows that after eight weeks of PPC and PLA implantations in mice, there was no immune response to the PPC implant, whereas multi-layer fibrous tissues were noted around the PLA constructs due to the acidic degradation of this polymer. These results illustrate the favorable degradation properties of PPC [81]. Furthermore, it should be noted that the degradation byproducts of PHB can be useful for cell growth [82]. The average reported pK_a of the degradation products from PLA, PGA and PLGA are nearly 3.5, which can be considered as a semi-strong acidic environment. Therefore, upon clinical application of these polymers, care must be taken to ensure their long-term degradation.

The explanation site of PPC-ST50 (a) and polylactic acid (PLA) (b) eight weeks post-surgery, and haematoxylin and eosin staining of paraffin sections of the implantation site at eight weeks around PPC-ST50 composite (c) and PLA (d). After eight weeks, a prominent foreign body reaction could be observed in the PLA implantation zone. However, the inflammatory response to the PPC-ST50 composite resolved dramatically. The PPC-ST50 and PLA scaffolds are present in the H&E images may not adhere to the glass slides during histological staining. Figure reproduced with permission from [81]. Copyright (2015) American Chemical Society.

3.3. Commercial Application of Polyesters

PLGA, PLA, and PCL are amongst the most widely used polyesters for the fabrication of sutures, drug delivery and implants as summarized in Table 3 . PLGA has been used in commercial sutures since the 1970s (e.g., Vicryl ® with the latest and most widely used PGA-sutures on the market as Vicryl Rapide ® and Panacryl ® , manufactured by Ethicon Inc., Edinburgh, United Kingdom) [83]. In addition, PLGA has been used for drug delivery applications, e.g., Lupron Depot ® , Sandostatin ® Depot, and Risperdal ® Consta ® [83]. PCL is used for the fabrication of tissue repair patches (i.e., Ethicon Inc., Edinburgh, United Kingdom) and as a filling agent to fill non-load bearing cavities in bone. PHB based biomaterials are mainly sutures (i.e., Phantom Fiber™ (Tornier Co., Amsterdam, The Netherlands), MonoMax ® (Braun Surgical Co., Melsungen, Germany)) and surgical mesh such as TephaFlex ® mesh (Tepha Inc., Lexington, MA, USA), GalaFLEX mesh (Galatea Corp., Lexington, MA, USA) and Tornier ® surgical mesh (Tornier Co., Amsterdam, The Netherlands). Furthermore, a few medical disposable products are available in the market made of PBS such as Bionolle ® 1000 and 3000 (Showa Highpolymer Co. Ltd., Tokyo, Japan).

Table 3

Commercial products made from biodegradable polyesters and their applications.

For load bearing applications, PLA is the most used polyester due to its intrinsic high mechanical strength (56.96 MPa compression and 3500 MPa tensile modulus) [33]. PLA is used in internal fixation devices, such as screws, plates, pins, and rods to support the repair of broken bones and hold them together [84]. However, in vivo studies show that PLA interferes with the bone remodeling process by imbalancing the number of osteoblast and osteoclasts during the bone remodeling [85,86]. Considering the commercially available polyester-based products as shown in Table 3 , it can be observed that such products are mainly used as non-load bearing biomedical applications due to some unmet drawbacks. It is well-acknowledged that chemical and physical alterations of current-biodegradable polyesters are promising for enhancing their applications in the biomedical field. These approaches can be exploited to further extend the medical use of polyesters.

4. Modification of Polyesters

Polyesters are broadly used for biomedical applications. However, different approaches are undertaken to address their shortcomings. Polyesters are commonly hydrophobic with a low number of cell-motif sites within their structures which results in inferior cell interaction behavior. Different physical and chemical modification techniques have been used to enhance their biological activities that are briefly described in this section.

In the physical modification, the molecular structure of polymers is not changed and an additional component(s) is mixed with the polymer; either by solvent casting or melt blending techniques. In the chemical modification, the molecular structure of the polymer is changed. There are two pathways; (a) copolymerization of the building blocks of polyesters to form a new class of polymers; and (b) modification of the polymer chain of the polyesters post-synthesis. In the following sections, the physical and chemical modification methods of the most used biodegradable polyesters for biomedical applications are discussed.

4.1. PLA

According to the European Bioplastics Association, more than 142,000 tons of PLA was consumed in 2013 which is more than 11.4% of the global bioplastic production capacity [91]. In biomedical applications, this polymer is also the most commonly used, and, thus, has been extensively modified by incorporating different organic and inorganic components. Additionally, PLA is the only member of the polyester family that has been used for load bearing applications such as orthopedic screws and plates, owing to the high mechanical strength of this polymer [92,93]. The properties of PLA depend on its molecular characteristics, crystallinity, morphology and degree of chain orientation.

Lactic acid, the building monomer of PLA, provides chiral configuration for PLA including D and l -polylactic acid. For load bearing applications, l -PLA is preferable because of the high strength and toughness of the resulting polymer; however, d -PLA is used in drug delivery systems due to its faster degradation rate. Three different crystallinity of the PLA including α, β, and γ forms are available. These three crystalline structures of PLA (α, β, and γ forms) display melting points of 185, 175 and 235 °C, respectively [94]. Regardless of the crystalline structure, and chiral configurations, PLA exhibits a very hydrophobic nature and a low ultimate elongation strain of nearly 10% [95]. In addition, PLA degradation in the body decreases the pH of surrounding tissues substantially, which may cause clinical complications such as necrosis and delayed healing. Similar to all other polyesters, the lack of cell motif sites within the structure of this polymer has also been a significant driving force to modify PLA. Therefore, PLA has been changed (a) to enhance its hydrophilic properties; (b) to increase the ultimate elongation strain; (c) to address the formation of acidic biodegradation products; (d) to improve the bioactivity; (e) and to increase the number of cell motif sites within its structure. Table 4 summarizes some of these physical and chemical modification approaches.

Table 4

Polylactic acid (PLA)-based structures applied in biomedical and tissue engineering applications.

Y-M: Young’s modulus; T-S: Tensile strength; C-S: compressive strength; R: resistance; E-M: Elastic modulus; S: stiffness; T-M: Tensile modulus; C-M: Compressive modulus.

The primary motivation to chemically modify PLA and to copolymerize lactic acid with glycolic acid to form PLGA was to develop a polymer with a more hydrophilic nature that degrades into less acidic products. This concept was initially hypothesized as glycolic acid has higher (more neutral) pK_a compared with lactic acid. However, the degradation products of PLGA are lactic acid and glycolic acid, and both of them still lower the pH of the surrounding tissue. In addition, PLGA displays a faster degradation rate, which is favorable for biomedical applications such as bioabsorbable sutures or drug delivery devices. Therefore, in parallel with PLA, the medical use of PLGA has also been expanded and, thus, a wide range of physical and chemical modifications have been made to both PLA and PLGA to enhance their properties.

The mechanical properties of PLA are favorable for load bearing applications, and the only mechanical shortcoming of PLA is its low ultimate tensile strain (e.g., around 10%). To enhance this property of PLA, thermoplastic polyurethane (TPU) and PCL have been physically added to this polymer [96,97]. TPU can tune its tensile modulus within the range of 7–1007 MPa at the strain of above 15% for neat PLA and a blend with 1:1 weight ratio, respectively. While, the addition of 50 wt %, PCL increases the elongation at break by nearly 10 fold (107% ± 4.7%). PLGA intrinsically displays very stretchable behavior with high ultimate tensile strain. However, the elongation and compression moduli of this polymer are lower than PLA, which drives the use of PLA for load bearing applications. In few cases, PLGA is blended with other polymers such as PHBV, which is a brittle but stiff polymer (high tensile modulus), to enhance the compression modulus and tensile moduli by two to three fold [106].

For tissue regeneration applications, the cell interaction behavior of PLA and PLGA-based composites needs to be improved, and the first material of choice to address this challenge is natural polymers, such as polysaccharides, polypeptides, and proteins. Tanase et al. introduced a polyester blend modified with chitosan and keratin to enhance cell interactions of the polyester [100]. An in vitro cell study using human osteosarcoma cell line shows a good cell viability and proliferation. Furthermore, the incorporation of polyethylene glycol (PEG) into the PLA matrix is used to enhance the surface hydrophilicity, and therefore, its biological behavior [98]. However, the addition of PEG results in a decrease in mechanical performance.

The cell interaction of PLGA also needs to be improved. Similar to PLA, natural polymers have been widely used to enhance the cell interaction capability of PLGA. Accordingly, PLGA knitted mesh is modified with collagen type I to develop a supporting biomaterial for cartilage and bone regeneration applications [111,112]. For chondrocyte growth and proliferation to help cartilage repair, 3D biodegradable scaffolds were formed with a different configuration of collagen inside the PLGA matrix and led to homogeneous cell distribution, natural chondrocyte morphology, and abundant cartilaginous ECM deposition. However, the mechanical strength of the most promising scaffold was at least half of the requirement for cartilage regeneration [111]. In another study, laminated mesh of PLGA and collagen was modified this time for bone-cartilage interface reconstruction. In this study, the collagen microsponge was crosslinked by treatment with 25% glutaraldehyde saturated vapor to cover the surface of the PLGA knitted mesh. The tissue engineered scaffold possessed the same behavior as a native osteochondral plug nine weeks after post-implantation regarding DNA expression of collagen type I and II. Another research group modified the surface of PLGA with poly- l -lysine using a water-in-oil-in-water emulsion or solvent evaporation technique [113]. Surface modification promoted the cell differentiation; however, it showed an adverse effect on the mechanical properties of PLGA. Gelatin was also used to modify a biodegradable polyester microfiber using electrospinning [107]. These examples demonstrate that various strategies can be used to enhance the biological properties of PLA and PLGA by incorporating natural polymers. The addition of natural proteins and polysaccharides, however, cannot potentially address the acidic degradation products and low bioactivity of PLA. To tackle this problem and to enhance the bioactivity of the PLA and PLGA based constructs, bioactive ceramics can be added to PLA, as the degradation products of ceramics are mostly basic and can promote the proliferation of native bones in the load bearing applications of these polymers.

There are numerous studies as summarized in Table 4 that investigates the effect of adding bioactive ceramics such as hydroxyapatite (HA) and β-tricalcium phosphate (β-TCP) to neutralize the acidic degradation media of polyesters and to evoke bioactive properties to these polymers [57,114]. The results of these studies demonstrate that the basic degradation of ceramic particles can neutralize the acidic environment. In a more clinical-based study, a method is developed for the treatment of skull defects by using PLA plates supplemented with carbonated apatite bone cement [115]. In these implantable plates, carbonated apatite cement particles are dispersed into the PLA sheets and are fixed to skull fractures. After 3–60 months’ follow-up, no complications concerning dislodgement or structural failure of the cranioplasty construct were observed. Several studies reported the positive impact of adding bone cement particles within the structure of PLA to enhance the cell interaction and bioactivity of PLA based structures [116,117]. Care must be taken to prepare a homogeneous composite of ceramic-polymer to achieve suitable mechanical properties and also predictable degradation behavior.

Hydrolysis by an alkali is the first step of chemical modification to provide an active site on the surface of a polymer [118]. In this procedure, the ester bond of biodegradable polymer is activated to bond with the hydrophilic –COOH and –OH or reactive –NH₂ groups in components such as an arginine-glycine-aspartic acid (RGD)-containing peptides, chitosan (CS), arginine and lysine, PEG, collagen, etc. Enhancement of wettability of the surface and biocompatibility of the scaffold are the main aims of these surface modifications. For instance, a PLA modified with RGD results in improvement in the cell densities and proliferation mediated through RGD–integrin interactions [119]. In spite of all the mentioned advantageous features for the polymers driven by post-polymerization, the possibility of side reactions, such as chain scission and racemization along with the complexity of this process, are the main disadvantages of this method. Therefore, post-polymerization functionalization is not the preferred route to obtain functional polyesters, and, also, these methods are not practical for the formation of 3D structures [21].

Advanced chemical modification methods are carried out to improve the physical and biological characteristics of both PLA and PLGA for the fabrication of 3D structures [21]. A general synthetic route for functionalization of PLA is copolymerization with 3-(S)-[(benzyloxycarbonyl)methyl]-1,4- dioxane-2,5-dione protected with benzyl alcohol followed by diazotization with sodium nitrite [120]. The deprotection process performed via catalytic hydrogenolysis of the benzyl groups using both PtO₂ and Pd/C catalysts results in an enhanced in vitro hydrolysis rate compared to PLA. The monomer functionalization has been extensively studied; however, few types of research evaluated the monomer functionalized polyesters for tissue engineering applications due to unknown biological properties that may lead to clinical complications [121,122,123,124].

The ring opening copolymerization of lactic acid through its carboxyl and hydroxyl groups is a possible way to chemically modify PLA and can produce high molecular weight polymers in combination with glycolide, δ-valerolactone, and trimethylene carbonate, as well as with monomers like ethylene oxide [125]. For instance, for drug delivery application, a range of PLA-PEG copolymers have been synthesized by using PEG block with a certain molecular weight and varying PLA segment lengths (e.g., M_n = 2000–110,000) using ring-opening polymerization of d , l -lactide catalyzed by stannous octoate [126]. Furthermore, PLA copolymerized with polyurethanes by copolymerization of l -LA and 1,4-butanediol to acquire mechanical properties for soft tissue engineering [127]. In addition to these general approaches to enhancing the physical and biological properties of PLA-based materials, more advanced polymer synthesis methods have been employed to make more clinically appropriate PLA-based materials. For instance, to eradicate the need for using organic solvents, there are numerous studies that attempt to generate water-soluble forms of PLA by grafting different molecules to this polyester.

Polymer grafting such as chitosan-grafted-PLA can be prepared by attaching PLA to the chitosan main chain, and these materials can be dissolved in low pH aqueous based solution [128,129]. PLA and PEG were also functionalized with FuCl to form a water soluble and crosslinkable form of PLA. This polymer has been extensively studied and analyzed by Jabbari’s research group [130,131,132,133,134]. In yet another study, a green approach was developed to synthesize this polymer under high-pressure CO₂ to eradicate even the use of organic solvent during its synthesis [135]. Conducting the synthesis in CO₂ gas expanded solution remarkably increased the fumarate crosslinking active site in the backbone of poly(lactide-ethylene oxide fumarate) (PLEOF) copolymer, hence, enhancing the mechanical properties and osteoblast cell adhesion and proliferation [135,136]. Interpenetrated polymer networks of PLEOF reinforced with gelatin and methacrylated gelatin were also synthesized with enhanced primary human osteoblast cell adhesion and proliferation [137,138]. As shown in Figure 2 , these interpenetrating polymer network structures were composed of micro (~20 μm), and macropores (540 μm) pores that promote the nutrient mass transfer and cell growth, respectively.