Biocompatibility of Resorbable Polymers: A Historical Perspective and

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Perspective

Biocompatibility of resorbable polymers: A historical perspective and framework for the future Daniela Pappalardo, Torbjörn Mathisen, and Anna Finne-Wistrand Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.9b00159 • Publication Date (Web): 11 Mar 2019 Downloaded from http://pubs.acs.org on March 12, 2019

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Biocompatibility of resorbable polymers: A historical perspective and framework for the future Daniela Pappalardo1, Torbjörn Mathisen2 and Anna Finne-Wistrand3*

1. Department of Science and Technology, University of Sannio, via dei Mulini, 82100 Benevento, Italy 2. Novus Scientific AB, Virdings allé 2, 754 50 Uppsala, Sweden 3. Department of Fibre and Polymer Technology, KTH Royal Institute of Technology, Stockholm, Sweden

KEYWORDS: resorbable polymers, poly(lactide), poly(caprolactone), poly(glycolide), poly(trimethylene carbonate), biocompatibility

ABSTRACT

The history of resorbable polymers containing glycolide, lactide, -caprolactone and trimethylene carbonate, with a special emphasis being placed on the timeframe of the 1960s1990s is described. Reviewing the history is valuable when looking into the future

ACS Paragon Plus Environment

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perspectives regarding how and where these monomers should be used. This story includes scientific evaluations indicating that these polymers are safe to use in medical devices, while the design of the medical device is not considered in this report. In particular, we wish to present the data regarding the tissue response to implanted polymers, as well as the toxicity and pharmacokinetics of their degradation products. In the translation of these polymers from “the bench to the bedside,” various challenges have been faced by surgeons, medical doctors, biologists, material engineers and polymer chemists. This perspective highlights the visionary role played by the pioneers, addressing the problems that occurred on a case by case basis in translational medicine.

1 Introduction Tissue engineering using resorbable polymers is an attractive area; it is as fascinating as it is complex, since it includes a number of competences. Different aspects of polymer technology are combined with advanced cell biology and cell-surface interactions.1 Generally, researchers are working with either advanced polymer chemistry, where they synthesize, functionalize resorbable polymers and design porous scaffolds, or they evaluate the cell response. This approach means that there is an area in between, which generates many questions, and there is a need for more details about cell-material interactions to advance in this area and develop


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new and improved commercial products in tissue engineering. It should also be remembered that the step from prototype to commercial product is large in the medical field compared to other sectors because of the strict regulatory requirements,2 but these aspects are not included in this perspective. Simple tests, such as cytotoxicity, are not sufficient to state that a material is biocompatible, and it can be applied in the medical field. With more knowledge from the biology side, it is obvious that there are many parameters that need to be evaluated and that the biocompatibility is, in many cases, dependent on the application and the design of the device – meaning that a polymer itself cannot be seen as biocompatible. The correct use of the term biocompatibility has been continuously under debate.3,4 In common usage, the term describes the ability of a material to exist in contact with the tissue of the human body without being harmful or toxic. However, materials and tissues can interact in numerous different ways. The definition of the term biocompatibility in one unique sense is thus very challenging, and it has evolved over time, developing in parallel with the knowledge in the field. The first implantable devices, developed during the years 1940 – 1980, were intended to be non-toxic, non-carcinogenic, and non-irritant. Consequently, this concept was translated to mean materials that were required to be non-chemically reactive and resistant to degradation. The methods to screen polymers for biocompatibility were described in the literature; however, it appeared increasingly evident that several factors influenced the success of the application.5 Subsequently, in light of the knowledge accumulated in both academic and clinical studies, a re-thinking of this concept occurred. It became obvious that the biocompatibility of the material / device would vary depending on the clinical use or indication; furthermore, in some applications, the material / device should actively interact with tissue, and in some cases, the materials should also degrade over time. The new paradigm for biocompatibility implied “the ability of a material to perform with an appropriate host response in a specific application”.6 This definition suggests that the


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materials are not passive, but they should have a function, and for achieving this, they should induce a response from the tissue that they are in contact with. As a result, the concept of biocompatibility is continuously in fieri, and as long as the knowledge in the field of biomaterials increases, novel visions and perspective are added and debated in the scientific arena.4,7 The document ISO 10993-1,2018 of the International Organization for Standardization describes the biological evaluation of medical devices within a risk management process.8 In addition to chemical, physical and mechanical tests, the biological testing is based upon in vitro and ex vivo test methods and upon animal models. Nevertheless, the results from the latter should always be considered with caution, as it cannot be unequivocally concluded that the same biological response will also occur when used in humans. Moreover, the biological responses are not only dependent on the application but may also be dependent also on the individual clinical history of the patient.9 There are indeed many factors influencing the interactions of a material with a biological environment and the end results. Several features, involving chemical, biochemical, physical and physiological mechanisms, as well as the dimension and shape of the device, are underpinning this result. The cellular cascades during wound healing depends on several factors, such as the extent of injury, loss of basement membrane structures, blood material interactions, provisional matrix formation, and extent of inflammatory response. Additionally, these factors in turn influence the extent or degree of granulation tissue formation, foreign body reaction, extent/degree of cellular necrosis, fibrosis or fibrous capsule development. This aspect has been largely described in previous literature.10-12 In the last decades, spectacular results in the field of materials for tissue engineering applications have been revealed. In this regard, synthetic degradable aliphatic polyesters emerged as leading materials. Their synthesis can be achieved by ring-opening polymerization (ROP) of lactones and lactides in the presence of a proper catalyst or initiator.


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Glycolide, lactide, -caprolactone and trimethylene carbonate are commonly used monomers for biomedical applications. The ROP approach can offer several advantageous features, such as a reproducible preparation and batch-to-batch uniformity, as well as a controlled microstructure and architecture of the polymeric chains, which, in turn, result in predictable properties of the final polymeric materials. Several reviews have focused on the synthesis and biomedical uses of aliphatic polyesters.13-17 The importance of the lifetime of degradable polymers has been summarized recently.18 Due to continuous and growing scientific interest since the first successful applications of synthetic degradable polymers dating back to the end of the 1960s with the development of the first commercial synthetic degradable suture, Dexon®, a plethora of degradable materials have been developed and used in several medical devices for implants, porous scaffolds or drug delivery systems.15,19 The use of degradable materials in different devices has been reviewed by several authors.17, 20-23 Current research efforts in the synthesis of polyesters strive to custom-design the material for the required final applications, and spectacular results have been achieved both from the academic and industrial milieu. In particular, materials based on polylactide (PLA) have been used to fabricate myriads of biomedical devices. A recent review by Tyler et al. described these applications, spanning from plastic surgery to orthopedic applications to the more advanced cardiac uses.20 Santoro et al. updated the research results related to the nanofibrous scaffolds made of PLA for tissue engineering applications. Although many of these materials have been tested in vitro, translations to humans will first require further studies in appropriate animal models.24 New advanced manufacturing technologies are further expanding the applications of these materials. Additive manufacturing, i.e., three-dimensional printing, emerged as an attractive and promising technology for tissue engineering and regenerative medicine, due to its ability


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to create custom-designed complex scaffolds. The use of PLA as a printable material for this purpose has been summarized recently.25 Moreover, recent research efforts have been devoted to selecting the synthetic pathways allowing the addition of biological motifs to elicit specific biological responses.26 The current and future generation of degradable polymers will indeed combine the ability of the materials to be resorbable and bioactive at the same time.7 It is evident, therefore, that the literature regarding novel and advanced synthetic approaches to resorbable polymers and/or novel manufacturing technologies for biomedical devices have been flourishing in the last few decades. At the same time, as summarized above, the review activities in these two fields can now count several notable examples. However, while we tend to develop new materials and manufacturing processes continuously, we forget the story behind the more important synthetic resorbable materials in terms of degradation, the local tissue response, toxicity of degradation products and pharmacokinetics. The history of the use of aliphatic polyesters in biomedical applications is full of examples of different issues that surgeons, medical doctors, biologists, material engineers and polymer chemists have faced in the translation of the polymers from “the bench to the bedside”. In particular, the timeframe including the years from the 1960s to 1990s is representative of the beginning of the use of resorbable polymers in the biomedical field. During that time, the pioneering researchers had excellent visions for the future of degradable materials, although their potential had not yet been unraveled. This perspective will concentrate on the tissue response to implanted polymers as well as on the toxicity and pharmacokinetics of the degradation products of the polymers containing glycolide, lactide, -caprolactone and trimethylene carbonate. We herein wish to review the historical evidence related to this forgotten 1960s-1990s time-frame that has shown that these polymers are safe for use in medical devices.


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2 Material chemistry and biological effects of breakdown products Clinical use of degradable polymers The numerous implantations in various animals and clinical use in humans have indicated the safe use of resorbable polymers made from glycolide, lactide, -caprolactone and trimethylene carbonate monomers, as well as their copolymers materials. Many of these studies have been accompanied by histological evaluations of the implant site (vide ultra). In addition, the use of radioactive labeled polymers to demonstrate the excretion of the degradation products from the body after complete degradation has occurred has been pioneered (vide ultra). The milestones relative to the advancement in clinical progress of resorbable polymers are briefly summarized in the timeline of Figure 1, while Figure 2 shows the product’s debut on the market strictly following the scientific progress.

Figure 1. Advancements in the clinical use of resorbable polymers.


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Figure 2. History of the development of the resorbable polymers: the scientific milestones (above) and the products on the market (below).

2.1 Polymers containing glycolide and lactide Medical devices based on polymers containing lactide and glycolide have a long history since their first use in 1970 as sutures.27 Notably, the majority of resorbable polymers used today for commercially available medical devices are still made from lactide and/or glycolide, except for those made from pure polydioxanone, e.g., PDS sutures.20,21 However, since the homopolymers are quite stiff and brittle, glycolide and lactide have been copolymerized with trimethylene carbonate28 or -caprolactone29 to fabricate softer and more pliable materials. Copolymerization with trimethylene carbonate and -caprolactone was shown to increase the degradation time of the materials because of their more hydrophobic character.30 The


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increased degradation time, in turn, could have a beneficial effect on, for instance, barriers and fixation devices, that should preserve their mechanical properties for a longer period of time than the timeframes offered by various compositions of lactide and glycolide alone. Notably, due to the bulk degradation by hydrolysis, the glycolide and lactide based homo- and co-polymers undergo mass losses within several months, depending on the polymer processing (e.g., crystallinity and specimen type), microstructure (e.g., tacticity), molar mass and distributions, while the loss of tensile strength can be manifested after a significantly shorter time, (i.e., only 1–2 months for PGA).17 The history of the use of aliphatic polyesters in biomedical applications is rich with examples of problems that have been faced and solved in the translation of the material in medical applications (vide ultra). The case of the use of poly(L-lactide) (PLLA) is emblematic of the kind of problems that may arise in transferring new material into biomedical use. The preparation of high-molecular weight PLLA was highly desired for biomedical applications where good mechanical properties were needed, such as surgical sutures or internal bone fixation devices. PLLA with a high molecular weight (Mw up to 106 g/mol, intrinsic viscosity [] up to 13 dl.g-1), a melting peak at 188.5 C and heat of fusion H = 64.7 J g-1 was thus obtained by ring-opening polymerization of L-lactide with tin 2-ethylhexanoate (SnOct2) at 100 C.31 This as-polymerized PLLA was used to develop bone plates and screws for the fixation of zygomatic fractures in maxillofacial surgery.32 After 3 years of implantation, four of nine patients manifested swelling, and six underwent reoperation. The explanted material showed remnants of degraded PLLA surrounded by a dense fibrous capsule. The swelling was classified as a nonspecific foreign body reaction to the degraded PLLA material. Notably, the presence of PLLA did not generate a detectable inflammatory reaction or foreign body reaction during a 2 years follow up. Instead, the tissue response may be caused by a change in the morphology, i.e., the fragmentation or the changes in the mechanical properties. After a


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continuous hydrolytic process, the PLLA disintegrated in small fragments, which could be observed by light or electron microscopy on the explanted tissue. An ultrastructural investigation of the degraded material showed an internalization of crystal-like PLLA material in the cytoplasm of various cells.32 The presence of these small particles, which may be engulfed by various cells such as macrophages, evoke the foreign body reaction. One way to overcome this problem is to decrease the crystallinity and especially the size of the crystallites. An example of this strategy is the copolymer of L-lactide with 4 % of D-lactide (PLA96). The biocompatibility and degradation of PLLA and PLA96 were investigated in vivo in rats by implanting in vitro pre-degraded particles. While the histological response was qualitatively similar for the PLLA and PLA96, for the PLA96, a lower crystallinity, higher degradation rate and smaller particle size were observed.33 It should be noted that the complications outlined above occurred several months (or years) after implantation.32 Further studies regarding the implantation of polymers containing glycolide and lactide showed a similar temporal behavior.34,35 The behavior of two amorphous and one crystalline PLA implanted in the paravertebral muscle of rats was studied during a period of 116 weeks. While the crystalline PLA remained almost unaltered after that period of time, the amorphous PLA materials were resorbed or metabolized (dependent on their molecular weight), and a mild to moderate foreign body reaction was observed.36 It should be mentioned that the degradation of polylactide (PLA) and polyglycolide (PGA) mainly occurs by bulk hydrolysis, and the hydrolysis rate is lowered by the crystallinity.18 The lactic acid and glycolic acid produced by the hydrolysis of PLA and PGA, respectively, are finally incorporated into the Krebs cycle (vide ultra, scheme 1) and are ultimately expired as CO2 by the lungs. Notably, the bulk degradation of both PLA and PGA does not immediately produce a decreasing of the mass of the implant, which is instead delayed to months or years, until the molecular weight of the polymeric chains is reduced to such extent that allows them to freely diffuse out from the


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polymer matrix. At this point, the surrounding tissue is not able to eliminate the acidic byproducts of a rapidly degrading implant, thus an inflammatory or adverse response may result. To overcome this problem, the incorporation of basic salts to prevent a drop in the pH during the degradation process was shown to be successful in vitro.34 Several studies concerned with the short and long term systemic and local effects following implantation of polymers containing glycolide and lactide have been reported in the scientific literature. Due to their excellent properties, they were quickly accepted by the medical community and, subsequently, rapidly tested in diverse biomedical applications. Several orthopedic devices that are approved for marketing by the FDA and are currently in use are made of the poly(lactide-co-glycolide) copolymer (PLGA).17 Apart from sutures for use in general surgery, various fixation devices for use in orthopedic surgery23 and barriers for the regeneration of soft and hard tissue in periodontal and oral surgery37 have dominated the clinical use of these materials up until today (vide ultra). An extensive summary discussing various pre-clinical and clinical results, analyzing the in vitro and in vivo studies regarding the toxicity and biocompatibility of polylactide (PLA) and polyglycolide (PGA) based materials, was written by Athanasiou et al.38 The pre-clinical studies with resorbable polymers in various orthopedic applications have been summarized by An et al.39 Fixation rods, tacks, suture anchors and screws are used to fixate small skeletal parts or fractures. Interference screws are used to fixate the anterior cruciate ligament in the knee. These implants are often quite large and some adverse events involving transient local fluid accumulation have been reported for a special type of fixation rod by Böstman et al.40 Although the reason for the local fluid accumulation is not entirely understood, the decrease in pH caused by degradation products from a fast degrading polymer such as PGA may easily override the buffering capacity of the surrounding tissue, especially in tissue with poor vascularization.


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Other uses of PLA and PGA homo- and copolymers are within periodontology. Within this field, Guided Tissue Regeneration (GTR) became an established modality for the regeneration of lost tissue in conjunction with periodontal defects.41,42 The first resorbable product made from synthetic resorbable polymer (GUIDOR Bioresorbable Matrix Barrier®) became available for clinical use in 1993.43,44 Resorbable devices for the closure of arterial punctures and guide tubes (Neurotube) for the reconstruction of peripheral nerve gaps also became available for clinical applications in 1990.45 The composition of the copolymers should be carefully tuned, since it can dramatically affect the final properties of the materials and, therefore, the final application. As an example of this, sintered microsphere matrix technology was used to fabricate tubular constructs mimicking the native ulnar bone of a rabbit by employing a semi-crystalline PLGA (composition 80:20 L-lactide:glycolide) and an amorphous PLGA (composition 85:15 D,L-lactide:glycolide). Measurements of the molecular weight, mechanical properties, and porosity were performed over a period of six months and provided a basis for theorization of the scaffold degradation process in vitro. The tubular scaffolds were implanted in a New Zealand white rabbit ulnar defect. The histological analysis and quantitative micro computed tomography (micro-CT) indicated that early solubilization of the semi-crystalline polymer created an acidic microenvironment that inhibited mineralized tissue formation. Amorphous PLGA tubular scaffolds, which contained D,L-lactide, demonstrated more mineralized tissue formation at the interior matrix compared to semi-crystalline PLGA tubular scaffolds, which contained Llactide.46 Thus, the PLGA composition influenced the biocompatibility of the polymer. 2.1.1. Toxicity and pharmacokinetics of the degradation products The evidence demonstrating that polymers based on lactide and glycolide are not dangerous


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or harmful goes back to several decades ago and are still valid. During the 1960s, Kulkarni et al. pioneered the use of aliphatic polyesters as surgical materials and biomedical devices. The degradation of poly(L-lactide) and poly(D,L-lactide) were studied both chemically, by hydrolysis promoted by NaOH, and biologically, by dispersing carbon-14 labeled poly(lactide) in 14-day-old chick embryo liver homogenates and determining the increase in radioactivity of the centrifugate filtrate as a function of time. The data showed that the poly(L-lactide) was less prone to degradation than the poly(D,L-lactide). The results of the medical evaluation of the polymer in the rod, suture and films were encouraging. Kulkarni envisaged that such degradable polymers could have a supportive role and that they could degrade by hydrolytic scission, forming non-toxic products that could be eliminated through normal excretory routes.47 Radioactive labeled PGA and copolymers with lactide have been implanted into rats to study the degradation kinetics and excretion of the degradation products. After complete degradation, the remaining radioactivity was described as negligible. In particular, carbon-14 and tritium labeled PGA and its copolymers with lactide were studied in Sprague- Dawely rats for documentation of the resorption rates in bone and soft tissue.48 The implanted materials were pellet-shaped with weights ranging from 5 to 6 mg and were implanted into the tibia and the abdominal wall. The animals were sacrificed after 1, 2, 3, 5, 7, 9 and 11 months, and the implantation sites were excised from each animal along with the liver, spleen, kidney, lung and a proportion of the muscle tissue. Depending on the “curing” time, which eventually affected the degree of crystallinity in the material, the half-life time of the implant was determined to vary from 0.85 to 5 months. More importantly, the hydrolytically degradation rate was dependent on the copolymer composition, with the 50:50 glycolide:lactide copolymer degrading faster (1 or 2 months), while the degradation time increased by increasing the amount of lactide (the 15:85 copolymer degraded in 5-6 months). Notably, a


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more recent description from the company declared that the DEXON sutures are cleared from the body in approximately 3 months. Although the samples were poorly characterized and, therefore, the results are not accurate, this observation represents an example of how the degree of crystallinity can influence the degradation. For the lactide/ glycolide copolymer 50:50 in composition, only 5.6 % of the original activity remained one month after implantation. Indeed, this 50:50 polymer is completely amorphous; as a result, the resorption process is much faster compared to the highly crystalline PGA. The distribution of the radioactivity in the tissues was negligible, with no significant tritium or carbon-14 activity observed in any of the tissues analyzed. The degradation behavior of microcapsules made of a carbon-14 labeled copolymer (50:50) of lactide and glycolide was further studied by Visscher et al. in vivo.49 The labeled PLGA microcapsules (average diameter of 30 m) were intramuscular injected in the gastrocnemius muscle of Sprague-Dawely rats to study the resorption of the polymer through histological examination as well as the remaining radioactivity at the implant site. At day 56 after implantation, only small remnants of the capsules could be detected and the SEM analyses showed that the microcapsules were well eroded and infiltrated with cells and the remaining radioactivity was negligible. These results encouraged the authors to draw the conclusion that correlations can be made between the actual breakdown and resorption of the polymer and their loss of radioactivity. The results are in good agreement with those obtained by Miller et al. in 1977, summarized above.48


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PGA hydrolysis Glycolic acid

Urine elimination

PLA

Glycine

hydrolysis Lactic acid

Serine

Pyruvic acid

Tricarboxylic acid - Krebs Cycle

CO2 + H2O

Scheme 1. Degradation pathways for polylactide and polyglycolide.

The degradation of the PLA and PGA occurs by hydrolytic random scission of the ester bonds (Scheme 1). Polyglycolide is degraded into glycolic acid, which is cleared from the human body through urine or may be further oxidized into glyoxylate and oxalic acid, which is also cleared via urine. Glycolic acid is also an intermediate in the photorespiratory carbon oxidation cycle; it is converted into glycine, which is further degraded into carbon dioxide and expired via the lungs. PLA is degraded to lactic acid, which is a human metabolic byproduct. Within the body, lactic acid is derived from glycogen breakdown, from amino acids, and from dicarboxylic acid. The sources of production include muscular activity and liver and blood metabolism. Normal human blood contains lactic acid, while it is present in the human skin as a lactate. Its metabolism has been studied in rabbits after injection of

14

C-labeled L-lactic acid. The

main part is oxidized in the tissues to carbonic acid and subsequently expired via the lungs.


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Only a small part was found as glucose and glycogen. The D-lactic acid is first converted to the L isomer in the liver before oxidation to carbon dioxide. Quantitative data with regard to the acute toxicity of glycolic acid and lactic acid are respectively summarized in Table 1. The table has been reproduced from a toxicological evaluation of polymers and copolymers based on glycolic and lactic acid.50

Table 1. Toxicological evaluation of glycolic acid and of lactic acid. Compound

Species

Glycolic acid rat Glycolic acid rat Glycolic acid guinea pig D,L-lactic acid rat D,L-lactic acid rat D,L-lactic acid mouse D,L-lactic acid mouse D,L-lactic acid rabbit D,L-lactic acid guinea pig s.c. = subcutaneous; i.p. = intraperitoneal

Application oral intravenous oral oral, s.c. i.p. oral s.c. oral oral

LD mg/kg LD50 LD50 LD50 LD50 LD50 LD50 LD50 LDlow LD50

1950 1000 1920 3730 2000 4875 4875 500 1810

The acute toxicity of glycolic acid is low and it is present as a metabolite in the human body. The average daily excretion of glycolic acid in urine from children (3.5 to 13.5 years of age) has been found to be 42 mg.51 The low toxicity of glycolic acid is further reflected by the low toxicity of ethylene glycol since glycolic acid is an important metabolite of ethylene glycol degradation in man. The lethal oral dose of ethylene glycol in humans is estimated to be 1.6 g/kg body weight, which correspond to 112 g for a person weighing 70 kg.52 Large doses > 500 mg/kg/day of ethylene glycol can lead to systemic and developmental toxicity in mice. It is thought that glycolic acid is the main reason, when the levels of glycolic acid exceed the capacity of maternal blood buffers to neutralize the acid, that a maternal metabolic acidosis


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may result, which is the major acute toxicity symptom of accidental ethylene glycol poisoning.53 In accordance with its role in the intermediary metabolism of cells, the acute toxicity of lactic acid is low. Most mammals, including humans, tolerate oral doses of more than 1500 mg/kg. Massive oral doses of D,L-lactic acid caused weight loss and anemia in rats and lowered the level of carbon dioxide in the blood. Like other acids of moderate strength, the free acid acts as an irritant in particular on the skin, eyes or mucous membranes. The few studies carried out with subchronic and chronic applications of D,L-lactate showed no accumulation or cumulative effects. When babies aged 2-4 weeks were fed with a diet containing 0.4 % D,L-lactic acid, no effect on their growth was observed. After a two-month period of feeding premature babies with an average daily dose of 800 mg per kg body weight of D,L-lactic acid or D-lactic acid, a metabolic acidosis was noted. L-lactic acid did not have such an effect.54 Both glycolic acid and lactic acid have also found safe uses as cosmetics components.55 Glycolic acid has been approved by FDA for use as an indirect food additive.56 Lactic acid was approved for use as direct food additives and is recognized as safe for use beyond infancy at concentrations that do not exceed good manufacturing practices. An acceptable daily intake (ADI) was not be specified for L-lactic acid, while for D-lactic acid, the ADI of 0 to 0.1 g/kg was established.57 PLGA has been largely used also in drug delivery systems for controlled release. Microspheres containing bioactive agents did not show undesirable reactions locally or systemically. Additionally, in this case, the degradation occurs via hydrolytic chain cleavage.58 Notably, the reaction to the degradable polymer is dependent on the different tissues; therefore, the toxicological evaluation it is strictly in reference to the application. Recently, a method to detect the degradation behavior of PLGA nanoparticles intravenously


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injected into different tissues of mice was developed. The authors underlined the difference between the degradation observed in vitro and in vivo. The degradation rate was not only dependent on the nanoparticle´s sizes, it also changed depending on the tissue.59

2.2. Polymers containing trimethylene carbonate Trimethylene carbonate has been in clinical use since early 1980. The use of trimethylene carbonate in resorbable polymers has several advantages as long as the device is not intended for use in high load applications, e.g., orthopedic fixation devices. The homopolymer of trimethylene carbonate, poly(trimethylene carbonate) (PTMC), is an amorphous material with a soft rubber-like character.60 PTMC is stable in phosphate buffer solution. The hydrolysis behavior of specimens made of PTMC was monitored at 37 C in a phosphate buffered solution at pH 7.4 for a period of over two years; in these conditions, the PTMC does not degrade.61 The degradation behavior of PTMC polymers was studied further both in vivo and in vitro in various conditions. PTMC rods (3 mm in diameter and 4 mm in length) implanted in the femur and tibia of rabbits degraded by surface erosion. The mass loss of high molecular weight PTMC (number average molecular weight = 457x103 g/mol) specimens was 60 wt% in 8 weeks, whereas the mass loss of the lower molecular weight PTMC (number average molecular weight = 89 x103 g/mol) specimens in the same period was 3 times lower. The molecular weights decreased with time, but the decrease was slower than the mass loss. It could be concluded that the enzymatic degradation plays an important role in the surface erosion of PTMC in vivo. Degradation experiments in vitro carried out with lipase solutions from Thermomycens lanuginosus confirmed the degradation by surface erosion. Again, in this latter case, it was observed that higher molecular weight specimens had higher enzymatic erosion rates than the lower molecular weight ones. This behavior was related to the


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difference in surface properties of the two polymers, which strongly affected the lipase activity. In the absence of lipase, moreover, the hydrolysis almost did not occur, and it was also independent on the pH. The degradation products are non-acidic, and this is a positive aspect for the in vivo applications. The mechanical properties of the polymers are mostly preserved with surface erosion.62 The response of the surrounding tissues to properly prepared PTMC membranes was histologically evaluated for Guided Bone Regeneration (GBR) of mandibular angle defects in Sprague-Dawley rats. The PTMC membranes (Mw = 443000 g/mol) were compared to collagen and expanded polytetrafluoroethylene (PTFE) membranes. Similar bone healing was observed in defects covered with all barrier membranes, with no statistically significant differences between them. A qualitative histological evaluation by light microscopy showed that after 12 weeks, most of the membranes made of PTMC or of collagen were degraded. The quantitative evaluation of bone regeneration was assessed by micro-radiography and micro-computed tomography. After 12 weeks, a comparable amount of bone formation was observed in defects treated with the PTMC membrane and with collagen or the PTFE membrane used in the same study. The in vivo resorption occurred, the homopolymer degraded via surface erosion involving cellular-mediated processes, and, after 12 weeks, the polymer was phagocytosed. The polymer induced a mild and transient tissue reaction when subcutaneously implanted. It was concluded that the PTMC membrane appeared suitable for GBR, and the lack of the requirement of a second operation for removal is a valuable features compared to the other non-degradable barrier membranes.63,64 An increase in the amount of trimethylene carbonate in the copolymer will delay the resorption profile and the tensile strength retention in time. For example, the TMC/D,Llactide copolymer degraded faster than the correspondent homopolymers in vitro at 37 C in phosphate buffered solution at pH 7.4 and were resorbed in less than one year; the -


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caprolactone copolymer degraded more slowly.61 Furthermore, the incorporation of trimethylene carbonate reduces the risk for a local pH decrease around the implant, because the degradation products have a neutral pH. Resorbable polymers that contain trimethylene carbonate will primarily degrade in vivo by simple hydrolysis. In the later stage of the degradation process, phagocytosis play an important role in the final clearance of the material from the body. The degradation products from the trimethylene carbonate moiety are carbon dioxide and 1,3-propanediol, and the latter will mainly be excreted in the urine.65 The degradation and the tissue response after subcutaneous implantation in rats of polymer films made of PTMC (Mn = 3.16 x 10-5 g/mol) and its copolymers with D,L-lactide (52 mol %) and -caprolactone (89 mol %) were evaluated for a period up to one year. As above, for the homopolymer PTMC, a decrease in mass without changing the molecular weight was observed with time. The PTMC was totally resorbed in less than one year. The copolymers, instead, degraded much slower than PTMC; moreover, they degraded by bulk hydrolysis of the esters bonds. When using PTMC, at the tissue-implant interface, giant cells as well as polymer phagocytosis could already be observed after 5 days. The tissue was very active, with newly formed blood vessels and infiltrating macrophages observed at the implant site. After 12 weeks, the polymer was phagocytosed, and after 1 year, no evidence of the polymer was found.63,66 A more recent evaluation of crosslinked copolymers of TMC and -caprolactone, prepared by gamma irradiation, evidenced the tissue response and degradation after subcutaneous implantation in the back of rats and underlined the importance of sample manufacturing. The crosslinked copolymers of TMC and -caprolactone were prepared starting from puriﬁed high molecular weight polymers, which were first compression moulded into ﬁlms and then crosslinked by gamma irradiation under vacuum. It was shown that by increasing the gamma irradiation dose, denser networks with higher gel percentages were obtained, the erosion rate of the copolymers in vivo was reduced, and the number of


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lymphocytes in the tissue around films were also reduced. In contrast, the amount of caprolactone in the copolymers neither substantially affected the tissue response nor the in vivo erosion rate. Moreover, by increasing the gamma irradiation dose, the number of macrophages and giant cells at the tissue-polymer interphase, evaluated after 5 days, decreased.67 The increase in gamma irradiation, indeed, yielded a flexible copolymer network having a lower crystallinity with respect to PTMC; at the same time, TMC is not inflammatory upon degradation. These aspects could affect the tissue response and could be responsible for the observed milder tissue response in these cases. The MAXON suture was the first material containing trimethylene carbonate that was used clinically. Implantation in both rats and dogs followed by histopathological examination revealed a mild foreign body reaction towards the material.68 A mild foreign body reaction toward copolymers containing trimethylene carbonate and D,L-lactide was also described in an application for bone regeneration on rabbit skulls. The pilot study aimed to evaluate for the first time the suitability of degradable polymers for guided tissue regeneration (GTR). Bone defects were created in the frontal and parietal skull bones of New Zealand white rabbits. One test defect was covered with a polylactic acid membrane, while other tests were performed with the copolymer of lactic acid and trimethylene carbonate. The histological analysis after 6 weeks postsurgical showed that, in the tests covered by the membrane, a continuous bridge of regenerated bone was formed, with almost the same thickness as the surrounding bone in the case of the TMC based copolymer. Inflammatory cell infiltrations or necrotic tissues were not observed.37 In a study concerning bone augmentation around oral implants in monkeys, a copolymer made of D,L lactic acid with trimethylene carbonate in a 70/30 ratio was implanted as a barrier to hinder soft tissue from invading the bone defect.69 The control implant was made of a non-resorbable polytetrafluoroethylene (PTFE) barrier. Although the study discovered a superior performance of the non-resorbable polymer with respect to the TMC


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based copolymer, the tissue response around the barrier was described as a mild foreign body reaction. The author concludes that the test barrier probably failed due to its inability to maintain space and to the lack of selective tissue exclusion properties until the initial osseous fill of the defect area was complete.69 Other uses of copolymers of TMC are in the field of nerve regeneration. Peripheral nerve regeneration has been studied with tubes made from copolymers of glycolide and trimethylene carbonate. The incorporation of trimethylene carbonate into the PGA polymer showed several advantages, such as a longer resorption time, which will shield and protect the regeneration of the nerve for a longer period of time. Moreover, the copolymers are preferred with respect to PGA because the softer materials facilitated the use of more pliable tubes. Another argument is that the material in general will generate less acidic degradation products, which may cause local inflammatory reactions. Polyglycolide tubes resorb too fast to allow the nerve to regenerate over distances longer than 5-10 mm and the incorporation of trimethylene carbonate has also been used by other researchers to increase the possibility of regenerating nerve tissue over a longer distance.70,71 In a pioneering evaluation, the Maxon® copolymer made of glycolide and trimethylene carbonate has been evaluated in vivo. After resection of the radial and ulnar nerve in the forearm of a primate, tubes made of collagen and tubes made from Maxon® were used to bridge the gap. The histological analysis after 14 months showed no remains of the material. For the 2 cm nerve gap in the radial nerve, it was difficult to determine the actual site of the implanted tube since the gap had been replaced with grossly normal nerve tissue, which is desirable. For the 5 cm gap in the ulnar nerve, the regeneration was poor for both materials, and only a very thin white structure of nerve tissue spanned the 5 cm long gap.72 2.2.1 Toxicity and pharmacokinetics of the degradation products The degradation products from the trimethylene carbonate moiety in PTMC and copolymers


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are carbon dioxide and 1,3-propanediol (Scheme 2).65

O O

H 2O

O n

n HO

OH

+

n CO2

Scheme 2. Degradation of poly(trimethylene carbonate)

The pharmacokinetics of carbon-14 labeled Maxon sutures were measured in rats and dogs by Katz et al. in 1985.68 Approximately 70 mg of sutures radiolabeled in the glycolide part were implanted in rats, while the sutures labeled in the trimethylene carbonate part were implanted into rats and Beagle dogs (723 mg was used in the Beagle dogs). The total radioactivity in the whole blood and plasma and the residual radioactivity were measured at the suture site, as well as in the urine, feces and expired air, for both animals. Of the total implanted radioactivity for the trimethylene carbonate labeled suture, 43.8 % was excreted in the urine, 2.3 % in the feces and 54.7 % in expired CO2 of the rats. Approximately 100 % of the implanted radioactivity was excreted after 28 weeks. Very little information regarding the toxicity of 1,3-propanediol is found in the literature. A single oral dose of LD50 was found to be 10 ml/kg in rats.73 Unpublished studies performed by Shell Chemical Company in 1977 disclosed that LD50 in rats was found to be 10 g/kg. In the same studies, the single dose dermal LD50 in rabbits exceeded 4.2 g/kg and 1,3propanediol was found to be practically nonirritant to rabbit eyes. It was also found not to be a skin sensitizer in guinea pigs when tested by the Magnusson-Klingman maximization assay. In mice, the single dose oral LD50 has been reported to be 4.7 g/kg.74 In one report where rats were fed a diet containing 500 ppm 1,3-propandiol for up to 15 weeks, it was claimed that the metabolites caused crosslinking of DNA in the liver and, to a limited extent, in the testis.75 Through in vitro tests with liver homogenates, they showed that 1,3-propanediol could be


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metabolized into malondialdehyde, which is a powerful crosslinking agent. Malondialdehyde, if at all formed from 1,3-propandiol, is expected to be a very short-lived metabolic intermediate, which is further metabolized into 3-hydroxypropionic acid and malonic acid. A subchronic toxicity study aimed to specifically look for DNA cross-linking, as proposed in the article by Summerfield cited above, and if this caused any toxic effects.76 Three groups of twenty rats, 10 female and 10 male, were nourished with food consisting of a vehicle and 1,3propanediol/kg/day. The study was ongoing for 90 days, after which the rats were sacrificed and macroscopic and microscopic examinations of the organs and tissue samples where performed. A spermatogenic analysis was also performed. The conclusion of the study was that oral administration of 1,3-propanediol for 90 days to rats at doses up to 1000 mg/kg/day did not cause any systemic toxicity. There was no effect on the sperm production, either in terms of numbers or morphology. Likewise, there were no pathological or functional effects on the liver, as determined by histopathology and clinical chemistry. The 1,3-propandiol is not normally a metabolite in the human body. Although not fully studied, is expected to be excreted in the urine. The presence of 1,3-propan-diol was reported in conjunction with abnormal bacterial metabolism in the gut.77 1,3-propandiol may be oxidized into hydroxypropionic acid or malonic acid according to the following pathway:


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HO

HO

OH

HO

O

1,3-propandiol

H

H

H

H

O O malondialdeide

HO

OH

O hydroxypropionic acid

OH

O O malonic semialdehyde

OH O

O

malonic acid

Scheme 3. Oxidation routes of 1,3-propandiol

Malonic acid, as well as its CoA derivative, hydroxypropionic acid and malonic semialdehyde are all normal metabolites in humans. It has been speculated that malondialdehyde is formed from 1,3-propandiol,75 but no studies so far have proven this to be true,76 probably because the compounds have extremely short lives. Injections of

14C-labeled

malonedialdehyde in

mice resulted in rapid metabolism,78 probably via oxidation by mitochondrial aldehyde dehydrogenase and decarboxylation, to finally produce CO2 and acetate.79

2.3 Polymers containing -caprolactone Products containing -caprolactone have been in clinical use since early 1990. Poly-caprolactone (PCL) is currently not used as a homopolymer in medical devices because of its extremely slow degradation rate due to the hydrophobic character of the polymer, as for polytrimethylene carbonate, when the molecular weight is above approximately 40 000 g/mol80, the hydrolytic degradation rate was assessed on PCL electrospun nanofibers, which were obtained in a controlled way by an hydrolytic degradation-assisted method from a


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solvent mixture of acetic acid and formic acid.81 The in vitro hydrolytic degradation of PCL nanofibrous scaffolds, analogously prepared by electrospinning of 1:1 formic/acetic acid solutions, were evaluated in pure water and a phosphate buffer solution according to the standard method ISO 10993-13:2010.82 After 100 days of immersion in water or the phosphate buffer solution, the original nanofiber structures were preserved, erosion was not observed, and the molar mass scarcely changed. However, an increase in the crystallinity was noticed, probably due to the temperature of the used method (37 °C), which is close to the crystallization temperature of PCL. In was concluded that the PCL structure was unaffected after 100 days of immersion both in water and in the buffer solution.83 While the PCL is highly stable in water or the phosphate buffer solution, the PCL degradation rate could be increased in the presence of lipases. As other aliphatic polyesters, PCL can indeed be hydrolyzed by lipases from various microorganisms. In an early report by Tokiwa and Suzuki, the hydrolysis rate with different lipases was determined by measuring the water-soluble total organic carbon (TOC) concentration. After 16 h of incubation at 30 C of a PCL (Mn = 6740 g/mol), the Rhizopus delemar lipase had the strongest activity, with a TOC of 310 ppm.84 The degradation behavior of solution cast films of PCL having different molecular weights has also been evaluated; it was found that in the presence of pseudomonas lipases, the weight loss was reaching 100 % after 80 h.85[83] Jenkins et al. analyzed the degradation rate of different PCL samples and recognized that the polymers with higher molecular weight degraded more slowly. Moreover, during the degradation, an increase in the degree of crystallinities was noticed. This observation was explained by hypothesizing that a degradation mechanism by chain scission operating in the amorphous phase could be responsible for the formation of shorter chains that are able to crystallize in new lamellae. The pure homopolymer of -caprolactone is a semicrystalline material, such as polyglycolide


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and poly(L-lactide); however, the very low glass transition temperature (Tg = - 60 °C) makes the polymer much more flexible than these latter two. Copolymerization with more hydrophilic monomers, such as glycolide or lactide, has been used to overcome the slow degradation rate of the homopolymer PCL and to speed up the resorption of devices containing -caprolactone. The use of -caprolactone in medical devices is indeed highly desired. It has been used to moderate the stiffness characteristics of polymers such as polyglycolide and poly-L-lactide, making the materials softer and more flexible. Resorbable polymers that contain -caprolactone will primarily degrade in vivo by simple hydrolysis. An analysis of the degradation products of films of PCL (Mn = 50000 g/mol) implanted in New Zealand white rabbits showed that the degradation is due to random scission with hydrolytic cleavage of ester groups. A kinetic study indicated that an autocatalytic process is operative, where the carboxylic acids end groups formed by hydrolysis participated in the transition state. The final stage of the degradation process was studied by implanting tritium labelled PCL, in form of capsule and powder, pre-degraded at a lower molecular weight ((Mn = 3000 g/mol). The process was monitored by measuring the radioactivity in the urine, feces, expired water, and the residual activity at the implant site. While the powders resorbed rapidly, with the 50 % of the sample radioactivity eliminated in 60 days, the capsules were resorbed slowly and after 180 days showed a 22 % weight loss.86,87 The degradation product from PCL is hydroxycaproic acid and will mainly be excreted in the urine. These early implantation studies with pure poly(-caprolactone) showed that degradation of this polymer in vivo was very slow. However, when the material had degraded to a certain stage where the polymer started to fragment, the particles generated were found to almost totally absorb within 60 to 120 days.87 Implanted cylinders of -caprolactone with smooth surfaces provoked only a mild transient inflammatory response, and in approximately two weeks, the encapsulation of the implant was


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noted, with macrophages and some giant cells present.88 A first stage of the degradation involves the non-enzymatic bulk hydrolysis of the ester bonds of the polymeric chains. A secondary foreign body reaction could be observed when the cylinders started to fragment after 9 months. When the molecular weight is decreased to approximately 5000 g/mol, the rate of chain scission slows down, and the cylinders start to fragment, thus generating a moderate foreign body reaction. This secondary stage is characterized by an increasing number of macrophages. The same reaction could be observed much earlier by subcutaneous implantation of PCL in powder form into rats. The macrophages were observed to engulf the polymer particles, and it is presumed that phagocytosis is an important part of the final degradation and absorption of the polymer.88 Although Capronor®, a slow release vehicle, was developed and used in some markets, the 

first device on the market containing -caprolactone in the western world was Monocryl , a monofilament suture made of -caprolactone and glycolide, having a multi-blocks structure composed of hard and soft segments. The tissue reaction was studied in rats and described as slight or minimal. The tissue response was characterized primarily by the presence of macrophages and fibroblasts, the typical foreign body response, with few lymphocytes, some white blood cells (PMN cells) and occasional giant cells. The cell response gradually diminished in concentration during the 119-day study period.28 Several articles dealing with nerve regeneration have described the use of copolymers of caprolactone with lactide. A first report described the use of a copolymer of -caprolactone with L-lactide for the inner layer of a two-ply nerve guide. The outer layer was made from a polyesterurethane/poly(L-lactide) mixture. The two plies of the guide were engineered with different porosities, and the guide was successfully used for the sciatic nerve regeneration of rats, and it was found to prevent neuroma formation.89 To avoid the use of polyurethanes, which may contain diphenylmethane diisocyanate and, upon processing and degradation,


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could release toxic and carcinogenic products, a two-ply novel nerve regeneration guide was subsequently developed by exclusively using a 50:50 copolymer of -caprolactone with Llactide.90 The cytotoxicity studies demonstrated no significant difference in cell death or growth inhibition when compared with the negative controls. According to ISO/EN standards, the material was not toxic.91 This novel guide was evaluated for the sciatic nerve regeneration of the rat.92 A two-year observation from 6 implanted nerve guides still showed traces of the implanted material. The foreign body reaction was mild and no foreign body giant cells were observed at this stage. Only small numbers of macrophages and fibroblasts were observed.93[ To overcome the long degradation time, a new rapidly degrading nerve guide composed of lactide and -caprolactone (50:50) was prepared by using a D,L-lactide composed of 85 % Llactide and 15 % D-lactide. The healing of a resected segment of the sciatic nerve in rats was studied by placing the distal and proximal portion of the nerve into a 12 mm long nerve guide. Histological analysis were performed after 1, 2 and 3 months and the foreign body reaction was described as mild, characterized by fibroblasts and macrophages and foreign body giant cells.92 In another study, bars (15x3x3 mm) of the same copolymer were implanted subcutaneously in rats and histological analyses of the implant area were carried out in the interval 3 to 12 months after implantation. The above picture of a mild foreign body reaction towards polymers containing -caprolactone was confirmed. Phagocytosis was observed after 4 months; at 6 months, the first ingrowths of cells into the fragmented material were observed. At that time, only 5 % of the initially implanted material volume was left. After 12 months, macrophages were still present at the area of implantation; however, no visual signs of the implanted material were observed in the histology sections.94 

PCL has also been extensively tested as a vehicle material in the product Capronor for slow release of levonorgestrel as a long acting contraceptive. Several animal and human studies


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have been conducted.95 Eight ovulatory women participated in a pilot test to evaluate the capsule release of levonorgestrel. All of them, except one, experienced suppression of ovulation. The study included serum assays of luteinizing hormone, follicle-stimulating hormone, estradiol, progesterone, and levonorgestrel on days 5, 8 to 18, and 22.96

2.3.1 Toxicity and pharmacokinetics of the degradation products Pitt studied the resorption of pre-degraded tritium labeled PCL, Mn = 3000 g/mol, by the subdermal implantation of both particles (approximately 10 m in size) and capsules (1 cm x 2.4 mm x 2.0 mm) into female New Zealand rabbits.86 O O

hydrolysis

O PCL

hydroxycaproic acid

n Urine elimination

Krebs Cycle

Scheme 4. Degradation and elimination routes of polycaprolactone

The resorption was studied by measuring the radioactivity in the urine, feces, expired water and the residual activity at the implant site. The powder showed a rapid resorption and approximately 50 % of the radioactivity had been excreted within 60 days after implantation. After 120 days, the remaining radioactivity was 9  4 % (n = 3). The study shows that once caprolactone has been degraded into fragments containing less than 25 monomeric units, it is rapidly absorbed by the surrounding tissue and excreted. The resorption of the capsules was instead slower; the measurement of the residual radioactivity at the implantation site after 180 days resulted in a weight loss of approximately 22 %. The pharmacokinetics of the Monocryl suture were also studied by analyzing a filament


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prepared by using carbon-14 labeled -caprolactone. The suture was implanted into 64 rats, which were killed after 1, 2, 3, 4, 8, 9, 10, 14 and 16 weeks. The radioactivity levels in the blood and in various tissues were measured by liquid scintillation counting.


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Urine, feces and expired air were also collected from these animals, and the weekly output of radioactivity up to the time of killing was determined. After 14 weeks, the radioactivity left in the animal corresponded to 0.15 % and 0.24 % of the implanted dose in males and females, respectively. Approximately 49 % of the radioactivity was excreted in the urine, 41 % was recovered in the expired air and 5 % was eliminated in the feces. These data clearly indicated that the main routes of resorption of the Monocryl suture involve the urinary excretion and pulmonary elimination as CO2. This behavior was confirmed in a more recent evaluation, where a tritium labeled low molecular weight -caprolactone polymer was implanted subcutaneously in rats to follow the absorption and excretion of degradation products.97[ After 135 days, approximately 92 % of the implanted radioactive dose was excreted from the animals and found in the feces and urine. No information is given regarding the analysis of expired air. The radioactivity of organs (heart, liver, spleen, lung, kidney, stomach, intestine, brain, ovary and uterus) showed that all organs were close to the background levels, indicating that degradation products did not accumulate in the body tissues. The main degradation product following the hydrolytic breakdown of PCL has been suggested to be hydroxycaproic acid (Scheme 4).88 Copolymers containing -caprolactone have been degraded in a cell culture medium containing fibroblasts; the degradation product was indeed determined to be hydroxycaproic acid.98 No data have been found in the literature on acute or chronic toxicity of this hydroxyacid. No literature has been found regarding the metabolism and excretion of 6-hydroxycaproic acid, the main degradation product of PCL. The most likely elimination routes of this acid would be the urinary excretion and the pulmonary elimination as carbon dioxide expired from the lungs, after a currently unidentified metabolic pathway in the body.28


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3 Conclusions and future perspective Several examples of products approved for marketing in both the US and Europe were included in this report and illustrate the history of the debut of the resorbable polymers in the biomedical field. Polymers containing glycolide and lactide have been in clinical use since 1970. Trimethylene carbonate has been in clinical use since early 1980 and -caprolactone since early 1990. All of these products have been extensively tested pre-clinically since the beginning of 1970 and are still in use today. However, in relation to the amount of research in this area, the number of new products on the market are extremely low. The reason for this is the difficulty of optimizing the design for a specific application, remembering that the design of the medical device is critical to the final results, including the mechanical properties, degradation profile, etc. We should also keep in mind that the next step, registration of the medical device and to prove its effectiveness, is also time consuming and very costly and might be an important threshold. The history of the clinical uses of polymers based of glycolide, lactide, -caprolactone and trimethylene carbonate suggests that these materials are safe, and several findings by toxicity and pharmacokinetics evaluations have been collected over several decades. In the translation of these polymers to clinical use, different problems have occurred and have been faced by surgeons, medical doctors, biologists, material engineers and polymer chemists, which have been the real “heroes” of this story. Thanks to them, we are not only able to trust these materials, but can also use them to develop further devices properly intended to face specific issues. The take-home message from this long story is the following “warning”: we should never forget that the design and development of devices for actual and future biomedical applications always require a deep understanding of the behavior of the device in vivo as well as of the complex phenomena


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determining the tissue response of the material, which are underpinning the specific biocompatibility. Before translating the materials to medical applications, several evaluations and testing of the materials must be done in vitro and in vivo. In tissue engineering, it appeared increasingly evident that the performances of the materials can be dramatically different if evaluated in vitro or in vivo. The history of the resorbable polymers in tissue engineering applications teaches us that the behavior of the material cannot be simply predicted from the in vitro study. The degradation in vivo is always more intricate than in vitro. Indeed, the complex in vivo environment around the material will affect the erosion, and in turn, the material and the degradation products may affect the environment. Moreover, before implanted in vivo and used for biomedical applications, these materials need to undergo various treatments required by the in vivo applications, for instance, the sterilization procedure, which may also dramatically influence the properties of the materials, as well as their degradation rate.99 To be successful with these polymers also in the future, we have to focus more on the design and obtain a better understanding of how these resorbable polymers can stimulate tissue regeneration using, for example, variations in mechanical properties and degradation profiles. The next step does not have to be advanced functional resorbable polymers, the old non-functionalized monomers and polymers have still much to give. Nondestructive methods for the evaluation of the behavior of the material interaction with the tissue are also required and we can foresee that they will be an important tool in the future. Some efforts have been made in this direction. The in vivo monitoring of the erosion of materials by using non-invasive and nondestructive fluorescence imaging has been pioneered by Artzi et.al.100 This concept was applied to monitor the PLGA degradation ex vivo in a chick femur model, and it may be used further for other imaging methods.101


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It appears increasingly evident that several aspects must be carefully considered and taken into account before a device can be turned from the “the bench to the bedside”. A feature that needs to be considered is the importance of mechanical forces and of mechanotransduction.102 Mechanical forces have long been implicated in regulating many physiologic and pathologic processes. The human body has dynamic complex biomechanical processes, and therefore, the mechanical properties of the material implanted play a significant role. The selection of materials for human use must consider the mechanical properties. Indeed, it is now evident that not only can the cells sense mechanical forces in different tissues but also that these mechanical forces are transduced into biochemical signals. Mechanotransduction may therefore have a pervasive role in regulating the cellular function at the tissue/material interface. Moreover, from the surgeon’s and clinical point of view, several questions arise,103 such as the following: How can the mechanotransduction of the materials be controlled? What is the role of the mechanical properties of the material in controlling cell growth and differentiation? The history summarized in this perspective indicated that the “biocompatibility” is tissue/application dependent. How can the response of the tissue to the surgical implant be controlled? Additionally, as the biocompatibility will rely on the patient, can we adjust the devices to the individual biology? While the 3D printing technique seems to have great potential for tailoring the shape and size of the devices, much still needs to be understood prior to customizing the devices to individuals or group of individuals. Starting from long known, well-established data -, i.e., “resorbable polymers are safe to be used in a biomedical device, and we have the proof”- we may envisage that the knowledge will further increase, and new techniques and methods will allow a deeper understanding of the mechanisms and behavior of interactions between the materials and tissue. The research is continuously


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ongoing, and we may foresee further notable advances in this field.

Acknowledgements AW and DP acknowledge the financial support from the Swedish Foundation for Strategic Research (RMA15-0010). DP acknowledges the financial support from VINNOVA, Mobility for Growth, (Grant Number 2013-04323) and the Marie Curie Actions FP7-PEOPLE-2011COFUND (GROWTH 291795).

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Martello, F.; Gerges, I.; Tocchio, A.; Tamplenizza, M.; Bellezza, G. Bringing Bioresorbable Polymers to Market; Elsevier Ltd, 2016; Vol. 2013.

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Williams, D. F. On the Mechanisms of Biocompatibility. Biomaterials 2008, 29 (20), 2941–2953.

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Ratner, B. D. The Biocompatibility Manifesto: Biocompatibility for the Twenty-First Century. Journal of Cardiovascular Translational Research 2011, 4 (5), 523–527.

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Gourlay, S. J.; Rice, R. M.; Hegyeli, A. F.; Wade, C. W. R.; Dillon, J. G.; Jaffe, H.; Kulkarni, R. K. Biocompatibility Testing of Polymers: In Vivo Implantation Studies. Journal of Biomedical Materials Research 1978, 12 (2), 219–232.

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within

a

Risk

Management

Process,

available

at

https://www.iso.org/obp/ui/#iso:std:iso:10993:-1:ed-5:v2:en. (9)

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Title : Biocompatibility of resorbable polymers: A historical perspective and framework for the future Authors: Daniela Pappalardo, Torbjörn Mathisen, Anna Finne-Wistrand


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Biocompatibility of Resorbable Polymers: A Historical Perspective and

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