Cross-Linked Poly(trimethylene carbonate-co-l

Article pubs.acs.org/Biomac

Cross-Linked Poly(trimethylene carbonate-co-L-lactide) as a Biodegradable, Elastomeric Scaffold for Vascular Engineering Applications Bronwin L. Dargaville,†,‡,§ Cédryck Vaquette,†,§ Hui Peng,†,‡ Firas Rasoul,†,‡ Yu Qian Chau,† Justin J. Cooper-White,† Julie H. Campbell,† and Andrew K. Whittaker*,†,‡ †

Australian Institute for Bioengineering and Nanotechnology, University of Queensland, Cnr College and Cooper Rds, QLD 4072 St. Lucia, Australia ‡ Centre for Advanced Imaging, University of Queensland, QLD 4072 St. Lucia, Australia S Supporting Information *

ABSTRACT: A series of copolymers of trimethylene carbonate (TMC) and L-lactide (LLA) were synthesized and evaluated as scaffolds for the production of artificial blood vessels. The polymers were end-functionalized with acrylate, cast into films, and cross-linked using UV light. The mechanical, degradation, and biocompatibility properties were evaluated. High TMC polymers showed mechanical properties comparable to human arteries (Young’s moduli of 1.2−1.8 MPa and high elasticity with repeated cycling at 10% strain). Over 84 days degradation in PBS, the modulus and material strength decreased gradually. The polymers were nontoxic and showed good cell adhesion and proliferation over 7 days using human mesenchymal stem cells. When implanted into the rat peritoneal cavity, the polymers elicited formation of tissue capsules composed of myofibroblasts, resembling immature vascular smooth muscle cells. Thus, these polymers showed properties which were tunable and favorable for vascular tissue engineering, specifically, the growth of artificial blood vessels in vivo. required to subsequently degrade during tissue remodelling, as the new blood vessel gains compliance. The scaffold material for this application must be designed to have elastic properties similar to those of native blood vessels, to be biocompatible and have appropriate cell adhesion properties, to be processable into a porous architecture, and be degradable to nontoxic products over a predictable timecourse. The use of synthetic polymers is attractive for tissue engineering because of the high degree of control that can be exerted over their structure and properties. The most commonly studied materials for tissue scaffolds are based on aliphatic polyesters such as poly(L- or D,L-lactide) (PLLA, PDLLA),5 polyglycolide (PGA),6 poly(hydroxyl alkanoate)s,7 polycaprolactone (PCL),8 and their copolymers.9 In particular, PLLA and PGA have been frequently used to engineer smooth muscle-like tissue.10,11 However, many of these materials have inappropriate mechanical and degradation characteristics for soft tissue engineering. For example, PLA, PGA, and their copolymers are stiff, brittle materials. The elastic modulus of polylactide with moderate molecular weight (Mυ of 23K−67K) has been reported to be in the range 3500−3700 MPa and the

1. INTRODUCTION Vascular tissue engineering is an area of growing importance given the large and increasing number of deaths attributed to cardiovascular disease.1 In response to this, about a decade ago, Campbell et al. developed a novel method for generating autologous tissue tubes with the potential for use as vascular bypass grafts.2,3 This involves exploiting the foreign body response toward a polymeric tubing device implanted into the peritoneal cavity of an animal, with the view of expanding to human patients. Cellular infiltration initiated by the endogenous immune response allows formation of a tissue capsule, composed of layers of myofibroblasts and mesothelial cells, around a polyethylene tubular template. The tube of living tissue (polyethylene tube discarded) is harvested from the peritoneal cavity after 3 weeks and grafted into an artery of the same animal in which it was grown, thus avoiding rejection. Over the next months, pulsatile blood flow stimulates remodeling and differentiation of the tissue to resemble a native artery.4 However, in the freshly grafted tissue tube, elastin expression is low, causing irreversible vessel dilation under pressure and hence contributing to compliance mismatch and failure of the graft. In order to increase the elasticity of the grafts, it has been proposed to engineer a synthetic polymeric scaffold to coat the inner tube of the intraperitoneal device and to become incorporated into the tissue capsule. The scaffold is © 2011 American Chemical Society

Received: March 8, 2011 Revised: October 14, 2011 Published: October 14, 2011 3856

dx.doi.org/10.1021/bm201291e | Biomacromolecules 2011, 12, 3856−3869

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elongation at break in the range 1.5−7%.12 In addition, a PLLA−PGA composite scaffold under long-term mechanical loading showed significant permanent deformation within relatively short time periods (less than 2 weeks).13 Poly(ester carbonate)s based on poly(1,3-trimethylene carbonate) (PTMC) are low glass transition, synthetic polymers that have been widely studied in recent years as biomaterials for applications requiring flexible substrates for cell growth.14,15 A major advantage of using TMC polymers is that they degrade to nontoxic, nonacidic products, 16 although hydrolytic degradation is slow compared to polyesters. 17,18 Copolymerization of TMC with a rigid, higher Tg component such as L- or D,L-lactide produces a material in which the mechanical and degradation properties can be tuned by adjusting the molecular weight and ratio of the monomers. 19 Low molecular weight P(LLA−TMC) is soft with liquidlike properties. 20 High molecular weight (>100 000 g/mol) statistical P(DLLA−TMC) is amorphous and elastomeric for TMC contents above 20%.19,21 However, these materials show some permanent deformation under stress.22 The literature overwhelmingly suggests that in order to achieve good mechanical properties poly(ester−carbonate)s must be either of high molecular weight or in some way crosslinked (chemically or physically). The advantages of crosslinking include enhanced elasticity and mechanical strength, shape retention, and slower degradation. 20,22−24 Various authors have reported network formation by photo- 20,25−28 or thermal-cross-linking29 of acrylate or methacrylate-end-capped TMC copolymers. There are also reports of cross-linking TMC polymers functionalized with phenyl azide,30 fumaric acid monoethyl ester,23 and coumarin moieties31,32 as well as physical cross-linking via ureidopyrimidinone end-groups33 or stereocomplex formation.22 γ Radiation-initiated cross-linking of nonfunctionalized TMC polymers has also been explored.34 There is limited literature addressing chemically cross-linked statistical P(LLA-TMC) networks. In one study Matsuda and co-workers reported photocuring of acrylate end-capped oligomers of TMC and L-lactide to give solid polymer films which were subjected to hydrolytic degradation. The surface elastic modulus was found to decrease with increasing degradation time.20 In the present study we have investigated a series of photocured cross-linked networks from low molecular weight acrylate end-capped statistical copolymers of TMC and L-lactide. The mechanical properties, cell adhesion, in vitro degradation, and in vivo tissue response have been evaluated with the view of developing a biodegradable tissue scaffold suitable for production, using the intraperitoneal implantation method, of vascular grafts as well as other hollow smooth muscle organs.

to 100% conversion as determined by proton nuclear magnetic resonance spectroscopy (1H NMR) (Bruker Avance 300 or 500 highresolution spectrometer). The reaction mixtures were cooled to room temperature, and the solid polymers were dissolved in CH 2Cl2 and precipitated into a 20-fold volume excess of methanol. The polymers were dried in vacuo and characterized by the following methods. The polymer composition was determined by 1H NMR (CDCl3). Chemical shift values are given in ppm, relative to chloroform (δ 7.24). Number-average (Mn) and weight-average (Mw) molecular weight and molecular weight distribution (M w/Mn) were obtained by gel permeation chromatography (GPC) using a Waters Alliance 2690 separations module equipped with a Waters 410 differential refractive index detector. Three linear 7.8 × 300 mm columns (2 × Styragel, 1 × Ultrastyrogel) were used in series, and samples (1 mg mL −1 in THF) were eluted with THF at a rate of 1 mL min−1 at 30 °C. Values are reported relative to polystyrene standards (517−2 × 106 g mol−1). The thermal properties of the polymers were evaluated by differential scanning calorimetry (DSC) using a Mettler Toledo DSC1 Star system, with a heating rate of 10 °C/min. 2.2.2. Acrylate End-Capped Polymers. Polymers and triethylamine (20× molar excess) were dissolved in CH2Cl2 in oven-dried roundbottom flasks and stirred at 0 °C. A solution of acryloyl chloride (20× molar excess) in CH2Cl2 was added dropwise over 20−30 min after which the reaction was heated to 50 °C for 16 h. The cooled reaction mixture was washed with water (3×) and the volume reduced before precipitation into a 20-fold volume excess of methanol. The acrylated polymers were stored in CH2Cl2 solution in the dark at 4 °C to prevent premature cross-linking. The yield of acrylated product was calculated from the mass of a small aliquot after solvent removal. Acrylated polymers were characterized by 1H NMR spectroscopy and DSC. 2.3. Solvent Casting and UV/vis Light-Induced Photocuring. Solutions of acrylate end-capped polymers in propylene glycol methyl ether acetate (PGMEA) (50 wt %) containing camphorquinone (0.5 wt %) were cast onto glass plates and the solvent evaporated in vacuo for several days. Films were cross-linked by UV/vis light irradiation using a Sunray 600 UVA broad spectrum flood lamp (average dose 1.0 mW cm2 for 10 min). Films (thickness 100−300 μm) were removed from the glass plates and stored in vacuo for a further several days before use. The degree of cross-linking was characterized by the dichloromethane-insoluble fraction and by dispersive Raman spectroscopy, using a Nicolet Almega XR spectrometer and a 780 nm laser. 2.4. Mechanical Testing. Mechanical testing was performed using a 5848 Instron Microtester. Dog bone sample dimensions (overall length 35 mm, width of ends 6 mm, length of narrow portion 12 mm, width of narrow portion 2 mm) were in accordance with Australian standard AS1683.11-2001 (methods of test for elastomers; methods 11: tension testing of vulcanized or thermoplastic rubber). Prior to the tests, samples for wet testing were incubated overnight in phosphate buffered saline (PBS) at 37 °C. Uniaxial tests were performed with a crosshead speed of 50 mm/min. Specimens were tested dry at 18 °C and in water at 37 °C. Young’s elastic modulus was determined from the initial part of the stress/strain curve from 0 to 10% strain. For each polymer, five replicate samples were tested. The different polymers were compared using a one-way ANOVA analysis followed by a Tukey posthoc test (p < 0.05). Cyclic tests were also performed in which the sample was extended to a certain strain (10, 20, and 30%) and unloaded to its initial position, followed by a 5 min recovery period. Cycling was repeated eight times each strain. The mechanical fatigue behavior of the 56 and 70% TMC films was assessed using an Enduratec BioDMA (Bose) machine. The tests were performed in triplicate in PBS at 37 °C. A preload of 0.01 N was applied. 10 000 cycles were performed at a frequency of 1.0 Hz, in the range 0−20% strain. 2.5. In Vitro Degradation Study. Films were cut into squares (12−20 mg each) and then immersed in one of two media: (1) PBS (without calcium and magnesium, Dulbecco); (2) 10% FCS (fetal calf serum, Gibco, Australia) in PBS. Amphotericin B (5 μg/mL), penicillin (100 units/mL), and streptomycin (100 μg/mL) were

2. MATERIALS AND METHODS 2.1. Materials. All solvents were purchased from Sigma-Aldrich and were analytical or HPLC grade. 1,3-Trimethylene carbonate (TMC) was purchased from Richman Chemicals, recrystallized twice from ethyl acetate, and dried under vacuum before use. Acryloyl chloride was purchased from Tokyo Kasei Kogyo Co Ltd. and used as received. Camphorquinone was obtained from Sigma-Aldrich, Australia. All other reagents were purchased from Sigma-Aldrich and used as received, unless otherwise stated. 2.2. Polymer Synthesis and Characterization. 2.2.1. Poly(trimethylene carbonate-co-L-lactide) Polymers (1−5). Trimethylene carbonate and L-lactide (various mol %) and stannous octoate (1.0 mol %) were added to silanized, oven-dried Schlenk flasks which were then heated to 130 °C under a nitrogen atmosphere for 3 days, 3857


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during irradiation to allow maximum exposure, and then sterilized by γ-radiation (25 kGy, 60Co Gammacell source). The in vivo study was carried out according to the guidelines of and following ethical approval from the University of Queensland Animal Ethics Committee. Eight male adult Wistar rats (280−320 g) were used. The animals were anesthetized with 2% isoflurane, and the skin was sterilized with Betadine. A small incision was made in the shaved abdominal wall, and polymer rods were inserted into the peritoneal cavity along with 5 mL of 0.9% saline (to maintain hydration of the cavity). The abdominal wall was closed using 4.0 Vicryl, and the skin was sutured with 5.0 Softsilk. All rats received subcutaneously 10 mg of Perfalgan as pain medication. One rat was used for each polymer, and six replicate samples were implanted. Three weeks after surgery, the rats were euthanized in a chamber precharged with carbon dioxide. The polymer rods and tissue capsules were harvested and fixed in 4% paraformaldehyde (in 0.1 M PBS pH 7.4) at 4 °C overnight. 2.7.2. Fixation and Embedding. The fixed tissues were washed with PBS, dehydrated in ethanol, and then infiltrated with xylene. The samples were attached to embedding molds in wax which were then allowed to set overnight at 4 °C. Sections (5−8 μm) were cut using a microtome. 2.7.3. Hematoxylin and Eosin (HE) Staining Protocol. Paraffin sections were incubated in hematoxylin and eosin solutions and then dehydrated in absolute ethanol. The slides were washed with xylene before being mounted in DPX resin (Distrene80, tricresyl phosphate, xylene (BDH, UK)). 2.7.4. Collagen Staining. A general stain for collagen was used. Paraffin sections were first stained in Weigert’s iron hematoxylin working solution followed by Biebrich scarlet−acid fuchsin solution. After washing, they were then differentiated in phosphomolybdic− phosphotungstic acid solution, aniline blue solution, and acetic acid solution, successively. The sections were washed again, followed by dehydration in ethanol and clearing in xylene. The slides were mounted in DPX resin. 2.7.5. Elastin Staining. Paraffin sections were stained in Verhoeff’s solution and were allowed to differentiate in 2% ferric chloride solution. After washing, sections were treated with 5% sodium thiosulfate solution. They were counterstained in Van Gieson’s solution, followed by dehydration in ethanol and clearing in xylene. The slides were mounted in DPX resin. 2.7.6. Immunohistochemistry. Paraffin sections were permeabilized and blocked for nonspecific binding by incubation with blocking buffer. Primary antibody (diluted with 3% bovine serum albumen (BSA) in PBS) was added to the sections and incubated in a humidified chamber overnight at 4 °C. To identify myofibroblasts and macrophages, specific primary antibodies for these cell types were used: (1) anti-α-SM actin for myofibroblasts (mouse IgG2a; 1A4, 1/ 200, Sigma-Aldrich) and (2) anti-CD68 for macrophages (mouse IgG1; ED1, 1/25, Serotec). After washing, sections were incubated with an isoform-specific secondary antibody and conjugated with a fluorescent dye (diluted in 3% BSA/PBS). Secondary antibody used in this study was fluorescein isothiocyanate (FITC)-conjugated antimouse IgG1 (DAKO). The sections were washed and then incubated with the nuclei stain Hoechst 33528 (1/10000; Molecular Probes). The slides were washed again and mounted in a mounting medium containing an antifade agent. Mounting medium was allowed to set overnight at 4 °C before examination.

added to the FCS-containing media as antifungal and antibacterial agents, respectively. The pH of all media was initially 7.4 and remained between 7.4 and 8.2 for the duration of the experiment. Polymer samples were incubated in duplicate at 37 °C in a non-CO2 incubator. Samples were removed at weekly time points up to 84 days, and the media was changed weekly for the remaining samples. The swollen mass (after blotting with tissue paper) and dried mass (after drying in vacuo to constant mass) were recorded. Water uptake was calculated from the ratio of the mass increase to the dry mass according to eq 1.

(1) The mechanical properties of the polymers were also monitored over the course of the degradation at 0, 14, 28, 56, and 84 days. Cyclic testing, as described in section 2.4, was performed using three cycles for each strain. An unpaired two-tailed student t test was utilized to compare data to the previous time point of the same polymer (p < 0.05). 2.6. In Vitro Cell Adhesion and Proliferation. 2.6.1. Film Preparation and Sterilization. Polymer disks (diameter 6.0 mm and thickness 200−300 μm) were attached to the bottom of a 96-well plate by spotting the well with dichloromethane then placing the disk into the well. The solvent was thoroughly removed in vacuo overnight. The disks were sterilized successively with 25, 50, and 75% ethanol solution, each for 10 min, followed by UV irradiation for 30 min. The films were rinsed twice with PBS and kept sterile until use. 2.6.2. Human Mesenchymal Stem Cell (hMSC) Culture. Human bone-marrow MSCs (kindly supplied by Dr. Gary Brooke at the Mater Medical Research Institute, Brisbane, Australia) from two different donor patients were cultured in Dulbecco’s modified Eagle’s medium, low glucose (DMEM) supplemented with 10% FCS, penicillin (100 μg/mL), and streptomycin (100 μg/mL) (Gibco, Australia). Cells were cultured until 90% confluent and then passaged at 2500 cells/cm 2 using Triplex until P8. The films (in triplicate) were seeded with 5000 cells in 200 μL of culture medium. The cells were cultured up to 7 days, and the medium was changed every 2 days. 2.6.3. Crystal Violet Assay. Cell proliferation was assessed using a crystal violet assay. At 1 and 7 days after seeding, cells were rinsed with PBS and fixed for 10 min with a 4% formaldehyde solution. A 0.1% (w/v) crystal violet solution in 0.2 M MES at pH = 6 (50 μL) was added to the 96-well plate and incubated for 5 min at room temperature. The crystal violet solution was removed, and the wells were rinsed with PBS until a clear solution was obtained. The crystals formed by the reagent were dissolved using 50 μL of a 10% (v/v) glacial acetic acid solution and transferred to a new 96-well plate. Absorbance was read at 590 nm using a plate reader (Spectramax M5, Molecular devices, Australia). An unpaired two-tailed student t test was used to compare the different time points for each polymer (p < 0.05). 2.6.4. Confocal Imaging. Cell-seeded films (in duplicate) were stained using a 1/1000 Hoechst (nucleus staining) and a 1/50 Phalloidin Texas Red (actin staining) solution in PBS. The films were fixed in filtered formaldehyde for 10 min and rinsed with PBS before immersing in a 0.1% (wt/v) Triton 100X solution for 10 min. They were rinsed again in PBS and 70 μL of the staining solution was placed on each film for 60 min. The films were rinsed in PBS three times and mounted on a glass slide with a mixture of glycerol/PBS (90/100 (v/v)). The seeded films were imaged using a LSM Zeiss 710 confocal microscope with a 20× objective and a 63× oil immersion objective. 2.6.5. Scanning Electron Microscopy. hMSC-seeded films (in duplicate) were fixed in a 2.5% glutaraldehyde solution. They were dehydrated and sputter-coated with platinum for 5 min. SEM images were obtained using a JEOL-JSM-6300 scanning electron microscope. 2.7. In Vivo Implantation Study. 2.7.1. Sample Preparation and Surgery. Polymer samples for implantation were prepared by dip-coating solid polyethylene cylinders (diameter 3 mm, length 15 mm) with a solution of polymer in PGMEA (50 wt %), to an approximate thickness of 150 μm, and the solvent was removed in vacuo. Samples were cross-linked by UV as in section 2.3, turned

3. RESULTS AND DISCUSSION 3.1. Synthesis and Characterization of Statistical Copolymers of Trimethylene Carbonate and L-Lactide. Copolymers were synthesized in high yield by bulk ringopening polymerization using stannous octoate as a catalyst, according to Scheme 1. No initiating species was added, as initiation was expected to occur adventitiously via trace amounts of water present.35,36 TMC feed contents of 30, 40, 45, 55, and 70 mol % were used. The properties of the 3858


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Scheme 1. Ring-Opening Polymerization of Trimethylene Carbonate (TMC) and L-Lactide and End-Group Functionalization with Acryloyl Chloride

Table 1. Composition, LLA Block Length, and Molar Mass of PLLA−TMC Polymers 1−5 polymer

feed ratio TMC:LLA

obsd polymer ratio TMC:LLA

yield (%)

Mn (g/mol) × 103

Mw/Mn

av LLA block length (monomer units)

theor statistical LLA block length (monomer units)

1 2 3 4 5

30:70 40:60 45:55 55:45 70:30

30:70 42:58 48:52 56:44 70:30

94 99 95 97 95

13.7 14.0 14.3 12.6 12.9

1.92 1.54 1.57 1.55 1.55

6.4 3.6 1.3 1.5 1.9

3.3 2.5 2.2 1.8 1.4

polymers are presented in Table 1. Polymers 1−5 had numberaverage molecular weights (Mn) between 12 600 and 14 300 g mol−1 as determined by GPC. 1H NMR analysis of the TMC and LLA peak intensities showed final polymer compositions which were either identical to the feed ratio or which differed by only 1−3%. The materials varied from a soft, pliable solid for the 70% TMC sample to a hard solid for the 30% TMC sample. The block structure of the polymers was determined by 1H NMR analysis using a previously published method.19 Briefly, the average block length of lactide was calculated from the relative peak intensities of the lactide methine resonances of LA−LA (δ 5.13) and LA−TMC (δ 5.00) diads. Block lengths ranged between 1.3 and 6.4, with the block length increasing with lactide content. The observed block length was generally slightly greater than the theoretical block length, calculated using standard copolymer equations, for completely random statistical copolymers (rTMC = rLLA = 1.0) (Table 1), indicating that the reactivity ratio of lactide slightly exceeds that of TMC under the experimental conditions. The thermal properties of polymers 1−5 are listed in Table 2. All polymers were completely amorphous, with no evidence of

3.2. Solvent Casting and Photo-Cross-Linking. Films were cast from PGMEA solution containing a photoinitiator. Several photoinitiators were trialed, including Irgacure 2959, Irgacure 819, Darocur TPO, diethoxyacetophenone (DEAP), and camphorquinone. After curing by UV irradiation, the crosslink density was characterized by dispersive Raman spectroscopy and by measuring the dichloromethane-insoluble gel fraction. Similar results were obtained with each photoinitiator at concentrations spanning the range 0.1−5.0 wt %. The films characterized in this paper were made using camphorquinone. For each of the five polymers, the dichloromethane-insoluble gel fraction was in the range 0.74−0.79, and the samples completely retained their shape during and after solvent immersion. Figure 1 shows an expanded region of the dispersive Raman spectra of the 56% TMC polymer (1) before (a) and after

Table 2. Thermal Properties of PLLA−TMC Polymers 1−5 and Cross-Linked Filmsa polymer

TMC:LLA

Tg (°C) polymers

Tg (°C) cross-linked films

1 2 3 4 5

30:70 42:58 48:52 56:44 70:30

6.4 0.8 −6.5 −5.8 −10.8

17.9 15.7 16.3 14.9 4.95

a

Figure 1. Dispersive Raman spectra of (a) a 56% TMC polymer film before cross-linking and (b) after cross-linking. 1765.0 cm −1 is the carbonyl stretching band of the LLA−TMC backbone; 1632.5 cm −1 is the double-bond stretching band of the acrylate end-group; 1451.6 cm−1 is the C−H bending band. The spectra are normalized to the C−H band.

Data are reported for the second heating scan.

crystalline melting transitions. The glass transition temperature (Tg) of the homopolymers of L-lactide and trimethylene carbonate have been reported as 57 and −17 °C, respectively.37,38 As expected, the Tg values of 1−5 were intermediate between the homopolymer values, decreasing with increasing TMC content. Acrylate end-capped polymers were prepared by reaction of the terminal hydroxyl groups with excess acryloyl chloride (Scheme 1). The degree of acrylation was determined by endgroup analysis using 1H NMR and ranged from 93 to 100% conversion of hydroxyl groups to acrylates.

photocuring (b). The peak intensities were normalized to the C−H bend at 1452 cm−1, which remained unchanged after photocuring.39 The carbonyl stretching band of the TMC and LLA backbone appeared at 1765 cm−1 and the carbon−carbon double bond stretch of the acrylate end-group at 1632 cm −1.39 After cross-linking the latter peak was reduced in intensity. The extent of cross-linking was quantified by comparing the 3859


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Figure 2. Quasi-static tensile properties of the PLLA-TMC films: (a) representative curves obtained in dry and wet states using an extension rate of 50 mm/min; (b, c, d) Young’s modulus, ultimate stress, and ultimate strain. Bars, +, ∗, #, †, and § show significant differences (one-way ANOVA) (p < 0.05).

mechanical properties of non-cross-linked LA-TMC have been reported.19,21 It was found that TMC was more ductile and lower in strength than the hard PLA component, and therefore it was expected that the higher the PLA content, the stiffer the polymer and consequently the higher the Young’s modulus.19 However, such a trend was not observed for our cross-linked polymers, indicating that the cross-link density may have more influence on the modulus than the monomer ratio, for the range of compositions used in this study. Pêgo et al. reported the Young’s modulus of a high molecular weight (Mn > 170 000) linear statistical PDLLA-TMC (50/50) material to be around 16 MPa, under dry conditions.19 For a similar composition (48% TMC) we obtained a modulus of 6.5 MPa for our cross-linked material. The ultimate stress of the dry polymers ranged from 5 to 11.7 MPa, consistent with values reported in several other studies,19,21 although again no clear effect of the copolymer composition could be observed. The ultimate strain was above 300% for all polymers, with the highest value of 400% found for the 48% and 56% TMC polymers. Grijpma et al. have reported similar lower ultimate strain values for a photo-cross-linked TMC-based copolymers.23 Pêgo et al. and Buchholz have reported much higher ultimate strain values, up to 900%, for non-cross-linked polymers.19,21 When the tests were performed in water at 37 °C after incubation overnight in PBS, the shape of the tensile curves remained similar, but the materials showed a 2-fold lower Young’s modulus (Figure 2b). The “wet” moduli of these polymers ranged from 1.5 to 1.8 MPa without any significant

CC/CH peak ratio before and after cross-linking and found to be in the range 33−51% for all of the polymers. Since, on average, only one acrylate group of each bifunctional polymer chain needs to undergo cross-linking to render the chain insoluble, the Raman data are consistent with the obtained dichloromethane gel fraction data. Cross-linking resulted in increased glass transition temperatures of the materials (Table 2) as a result of more restricted chain movement. The values remained below both room temperature and body temperature, and thus the materials were flexible for handling and for use in vivo. 3.3. Mechanical Testing. 3.3.1. Static Tensile Testing. Films of polymer 1, with the lowest TMC content (30%), were too brittle to cut into dog-bone samples and thus could not be tested. Figure 2a presents a typical stress−strain curve for the remaining four polymers. All polymers showed behavior typical of elastomeric materials. The Young’s modulus of human arteries, usually reported in the circumferential direction, is in the range 0.1−1.0 MPa for healthy subjects.40 These values tend to increase, that is, the arteries become stiffer, with age and with various disease states, such as atherosclerosis. The modulus of a vascular graft material must mimic as closely as possible that of the native tissue in order to avoid graft failure due to mechanical mismatch. Figure 2b−d shows the quasi-static tensile properties of the polymers as a function of polymer composition. No general trend was observed with TMC content. The moduli for the series 42, 48, 56, and 70% TMC were 3.8, 4.7, 6.5, and 2.9 MPa, respectively, in air at 18 °C. Previously, several studies on the 3860


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Figure 3. Stress/strain cyclic tensile behavior of the PLLA−TMC films. The test consisted of successively performing eight cycles at each of 10, 20, and 30% strain with a 5 min break between cycles. (a) 42% TMC, (b) 48% TMC, (c) 56% TMC, (d) 70%, and (e) three cycles at 30% strain for 70% TMC.

3.3.2. Periodic Cyclic Mechanical Testing. Cyclic tests at three strains (10, 20, and 30%) were carried out in water at 37 °C, with a 5 min recovery period between cycles. Figure 3a−d shows the typical cyclic stress/strain curves obtained for the four polymers. The Young’s modulus was measured for each cycle. The “wet” modulus was unaffected by both the strain applied and by the successive cycles, indicating that no strain hardening or softening occurred. Values of 1.6, 1.8, 1.1, and 1.5 MPa were obtained, in order of increasing TMC content, which are consistent with the static testing. The residual strain after unloading the sample and the plastic deformation after the 5 min recovery period were monitored, giving a measure of elasticity and resilience, respectively. It was found that all residual strain values decreased with increasing TMC content. The polymers with the lowest TMC content (42 and 48%) showed lower resilience than the higher TMC polymers (56 and 70%), as shown by a shift toward higher strain after each cycle in Figure 3a,b. This plastic deformation gradually increased when higher strain was applied; however, for both polymers the recovery was greater than 90% even at 30% strain. As can be seen from the slight residual strain immediately after unloading for the 42 and 48% TMC polymers at all three strain values (x-axis, Figure 3a,b), these two materials were not completely elastic. This residual strain ranged from

differences between the polymers. These values are approaching the modulus of human arteries.40 The decrease in the modulus under wet conditions has been previously reported for non-cross-linked DLLA−TMC copolymers, albeit to a smaller extent.19 This effect is most likely due to both water uptake and the higher temperature. After incubation in PBS overnight, the previously semitransparent dog-bone samples became opaque and had increased in thickness by about 30% due to swelling. The swollen dimensions were used for calculation of the modulus, and this may, in part, explain the lower values obtained for the swollen samples. The ultimate stress was significantly lower for the same reason, as shown in Figure 2c. Water uptake was expected to also increase the ultimate strain; however, the opposite trend was observed, as shown in Figure 2d. This may be explained by the plasticizing effect of low molecular weight oligomer or residual solvent in the dry samples which may have leached out of the wet samples. Pêgo et al. have reported this phenomenon and consequently used compression molding instead of solvent casting.19 However, those authors reported thermal degradation of the polymers at the high temperatures required by this technique. Compression molding was not used in our study for this reason and also due to difficulty in handling the low molecular weight semisolid prepolymers. 3861


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Figure 4. Representative curves of the fatigue behavior of the PLLA−TMC. The test consisted of cyclic extension to 20% strain at a frequency of 1.0 Hz for 10 000 cycles.

commercial polyurethane (Pellethane-80AE), while remaining elastic, displayed dynamic stress relaxation of 15% after 12 000 cycles, at 10% strain and a frequency of 1.0 Hz.43 Song et al. investigated γ-cross-linked PTMC and reported low levels of plastic deformation after 20 cycles at 50% strain but after 2 h recovery.44 For 20% strain, both our films did not show any plastic deformation at the end of the test. 3.4. In Vitro Degradation Study. The rate of in vitro hydrolytic degradation of PTMC is known to be molecular weight dependent and is slow compared to polyesters.17,18 In particular, acrylate-cross-linked PTMC has been reported to undergo negligible hydrolytic degradation.20 Several studies have shown higher degradation rates in vivo compared to in vitro, suggesting that an enzymatic or other cell-mediated process may be involved.17,45,46 However, little is known about the enzymatic or other reactive species that mediate this degradation. For this reason a degradation experiment was undertaken using a nonspecific biological source, 10% fetal calf serum in PBS. 3.4.1. Mass Study. The loss of dry mass over 12 weeks (84 days) for the five polymers is shown in Figure 5. The time course of degradation was very similar for each of the polymers, with no significant differences in mass loss between polymers or between the different fluids. In all cases, the greatest rate of mass loss was observed during the first week (4−9%), indicating initial leaching of low molecular weight, non-crosslinked components. The average weekly mass loss for each material over the 84 days was around 1%. Figure 6 shows the uptake of wet mass for the same samples. In most cases this can be interpreted as water uptake. For all samples, water absorption increased steadily with time (Figure 6), suggesting a gradual increase in hydrophilicity due to degradation. Water uptake decreased with increasing TMC content, as expected, since TMC is more hydrophobic, and hence less prone to hydrolytic cleavage, than lactide. Interestingly, the samples incubated in 10% FCS showed significantly greater mass uptake than those incubated in PBS, especially from 56 days onward. This high uptake could be due to protein adsorption onto the polymer surface.47 However, this appeared to have no effect on the degradative mass loss. Matsuda et al. studied the hydrolytic degradation of photocross-linked acrylate end-capped polymers of TMC and L-lactide in phosphate buffer up to 4 weeks and found that mass loss was positively correlated with water uptake.20 PTMC and PLLA−TMC (50/50) containing a poly(ethylene glycol) (PEG) 1000 block showed 7 and 12% weight loss and 40 and 65% water uptake, respectively, whereas PTMC containing PEG 200 or TMP (trimethylolpropane) showed only 4.3 and

0.7 to 5.4% for the 10 and 30% cycles. However, the 56 and 70% TMC polymers were significantly more elastic, showing complete recovery at 10% strain and between 0.3 and 1.7% residual strain for the 20 and 30% cycles, respectively (Figure 3c,d). This elasticity is demonstrated in Figure 3e which shows complete overlap of three nonsuccessive cycles for the 70% TMC polymer at 30% strain. Additionally, the two higher TMC polymers were completely resilient, showing full recovery after the 5 min rest period for all cycles and hence no shift to higher strain. A full quantitative analysis of the residual strain values after unloading the sample and after the 5 min recovery period for each cycle appears in the Supporting Information (Table S1). The literature shows that in order to obtain a flexible but strong material, high molecular weight appears to be essential for PLA−TMC polymers. Pêgo et al. have extensively studied these polymers and shown that statistical linear copolymers with TMC content above 50% were weak and irreversibly deformed even at low strain.19 Pospiech et al. synthesized a PLLA−TMC (50/50) block copolymer of intermediate molecular weight and reported that it behaved more as a thermoplastic rather than an elastomer.41 Cyclic testing on this material showed plastic deformation at less than 10% strain. Andronova et al. performed cyclic testing of a LA−TMC (34/66) block copolymer and reported a constant decrease in the area covered by the hysteresis loops, indicating a strain softening in this material, most likely leading to permanent deformation.42 Our cross-linking strategy was effective in enhancing elasticity and resilience compared to similar non-cross-linked materials, particularly below 10% strain. This is advantageous for cardiovascular and more generally for soft tissue engineering. 3.3.3. Fatigue Testing. Assessing the material response under continuous cyclic loading is important for materials that are to be used in a mechanically dynamic environment such as the circulatory system. The two most promising polymers from this study (4 and 5; 56 and 70% TMC) were subjected to cyclic elongation to 20% strain for 10 000 continuous cycles at a frequency of 1 Hz, without any recovery period between cycles. This simulates the conditions in vivo at a resting heart rate of 60 beats per minute. Throughout this experiment both polymers consistently showed elasticity as no buckling of the sample was observed. The Young’s modulus decreased by 13% for the 56% TMC films, whereas it stayed constant for the 70% TMC films, as shown in Figure 4a. The stress at peak gradually decreased by about 20% for both polymers (Figure 4b), indicating some dynamic relaxation due to internal chain rearrangement. Such behavior has also been observed for polymers known to be highly elastic. For example, Puskas et al. reported that 3862


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Figure 5. Degradative dry mass loss of the LLA−TMC films, treated with 10% fetal calf serum (FCS) or phosphate buffer (PBS) at 37 °C for 84 days. Error bars represent duplicate samples.

versus 9000 g mol−1 for polymer 2), allowing water to more easily penetrate the network. 3.4.2. Mechanical Properties. The tensile properties of the four polymers during degradation in PBS are shown in Figure 7. Figure 7a shows representative stress/strain curves. The tensile properties gradually decreased with incubation time. All polymers showed the same profile, that is, a decrease in slope of the initial linear part of the tensile curve, although the shape of the curve was retained throughout the degradation, indicating that they retained typical elastomeric behavior. As shown in Figure 7b,c, a decrease in the mechanical properties occurred after as little as 2 weeks. The decrease in Young’s

2.9% weight loss and water uptake of less than 5%. This suggests that an increased hydrophilicity, in Matsuda and colleague’s case imparted by PEG 1000, most likely leads to an increase in hydrolytic degradation. Our results are consistent with the behavior they observed for the more hydrophobic polymers. Storey et al. studied the degradation in PBS of methacrylatecross-linked DLLA−TMC polymers and reported 6.2% swelling after 69 days for a 40/60 TMC/lactide polymer. 29 This compares to 22% observed for our 42% TMC polymer. This could be explained by a lower cross-link density in our case (Mn of prepolymer 2400 g mol−1 for the literature polymer 3863


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Figure 6. Water swelling of the LLA−TMC films, treated with 10% fetal calf serum (FCS) or PBS at 37 °C for 84 days. Error bars represent duplicate samples.

A systematic quantitative analysis was carried out on the Young’s modulus, elasticity, and resilience over the time course of degradation, and the full tabulated data can be found in the Supporting Information (Table S2). For all polymers, the modulus was constant with repeated cycling at a given time point for the three strains. The TMC content had a strong impact on the recovery properties of the films, as it did for the nondegraded samples, with the higher TMC contents showing more complete recovery. For the 42 and 48% TMC polymers the residual strain after each cycle and the final plastic deformation gradually increased. The 56% TMC polymer was fully resilient up to 56 days. The 70% TMC polymer displayed

modulus and ultimate stress followed a linear path, which is consistent with the mass loss profile for these polymers. At the end of the experiment, all polymer films had retained only ∼10% of their initial tensile modulus. The ultimate strain remained relatively constant up to 56 days and then tended to decrease (Figure 7d). Pêgo et al. showed that a high molecular weight 50/50 DLLA-TMC polymer retained only 20% of its initial molecular weight after 10 weeks of degradation and that after 12 weeks no mechanical testing could be performed.48 In contrast, each of our polymers retained enough mechanical strength to be easily handled and tested up to 12 weeks, confirming the greater resistance of the cross-linked polymers to degradation. 3864


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Figure 7. Evolution of the mechanical properties during degradation for the PLLA−TMC films. (a) Stress/strain behavior for the 70% TMC polymer. (b−d) Mechanical properties up to 84 days in PBS. + shows significant differences (p < 0.05), determined by an unpaired two tailed t test between different measurements of the same polymer. (e) 70% TMC; three cycles successively applied at 10, 20, and 30% strain (5 min between cycles). (f) 70% TMC; three cycles at 30% strain (5 min between cycles).

the most promising properties. After 84 days the curves at the three different strain values followed the same initial path (Figure 7e), and the cycles at 30% strain were still overlapping (Figure 7f), indicating that this polymer was still able to fully recover even after this time. In addition, this polymer showed elastic behavior until 28 days, regardless of the strain. After this, it still retained reasonable elasticity for 10% strain, but not for the higher strains. Most of the previous studies of the degradation of TMCbased copolymers have been limited to mechanical characterization under quasi-static conditions. However, the cyclic performance of these materials during degradation is of major concern for vascular applications as the materials should retain their elasticity and resilience until the cells have developed an extracellular matrix which can take over the dynamic load bearing. The 70% TMC polymer 5 is a promising candidate for such applications. Polyesters are well-known to undergo bulk degradation, 49,50 whereas PTMC undergoes surface erosion.16,45 For LA−TMC copolymers it has been reported that a bulk mechanism predominates when the lactide content exceeded 30 mol %.51 In accordance with this, for our study, the onset of the loss of mechanical properties after only 2 weeks, steady increase in

water uptake, increase in opacity of the samples, retention of the sample shape, and minimal mass loss all suggest the operation of a bulk mechanism. For vascular applications the gradual degradation demonstrated for these materials will ensure that functional tissue can be generated before any significant scaffold failure occurs. 3.5. In Vitro Cell Adhesion and Proliferation. The biocompatibility and cytotoxicity of PLA−TMC polymers have long been investigated using different cell types, ranging from smooth muscle cells52 to cardiomyocytes.48 In our study, cell adhesion and proliferation were assessed with human mesenchymal stem cells (hMSC) from two different donors using a crystal violet assay. These cells are commonly used as a model in tissue engineering studies since they have the ability to differentiate into a range of different tissues. Figure 8a,b shows the cell proliferation profile for the five polymers. The cells adhered and proliferated on the polymers as indicated by the significant increase in optical density. The donor 1 cells gave more variation between polymers than donor 2; however, statistical analysis (one-way ANOVA with Games-Howel post hoc for unequal variance) showed that the only significant difference between materials at day 7 was between the control and 56% TMC. 3865


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Figure 8. Human mesenchymal stem cell adhesion and proliferation. (a, b) Crystal violet assay for two different human donors. An increase in the optical density indicates cell proliferation. + shows significant difference compared to the previous time point for each polymer. (c, d) Representative SEM images after 1 and 7 days of culture for the 42% TMC film. (e, f) Confocal images after 1 and 7 days of culture for the 42% TMC film (blue is the nucleus; green is the α-actin).

SEM observation confirmed cell proliferation, since the cells were initially sparsely distributed on the films but started to form a cell sheet by day 7. Figure 8c,d shows representative SEM images of the cells on the 42% TMC polymer. The cells, as observed in these images, were well spread and showed no evidence of cytotoxicity (for example, changes in cell morphology) after 7 days of culture. As discussed in section 3.2, there was a residual noncross-linked fraction in the films (21−26%), but this did not appear to impart cytotoxicity. These results are consistent with the findings of Declerq et al., who have developed a photopolymerizable polymer based on D,L-lactide, ε-caprolactone, and TMC and have shown that the residual sol content (around 10%) had little effect on the biocompatibility.53

We have also investigated the morphology of the cells using confocal microscopy. Figure 8e,f shows some representative images at days 1 and 7. Cells were well spread, and the α-actin staining revealed a well-developed network of stress fibers in the cytoskeleton. This morphology was consistently observed and was independent of the polymer composition. It has been shown that the stiffness of a substrate can affect the differentiation of stem cells.54,55 However, in our study, since the four polymers had similar stiffness to each other, no differences in morphology were observed. 3.6. In Vivo Rat Implantation Study. A rat implantation study was carried out in order to further evaluate the suitability of polymers 1−5 as scaffold materials for the growth of artificial vascular tissue using the peritoneal model described in section 1. The aim of this experiment was not to generate a functional 3866


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Figure 9. In vivo tissue capsule growth on the surface of polymer 5 (70% TMC). (a) Hematoxylin and Eosin Y stain for cell morphology; ×20; all cells stained pink. (b) Masson’s Trichrome stain; ×20; collagen stained blue; nuclei stained red. (c) 1A4 immunohistochemical stain; ×20; α-smooth muscle actin/myofibroblasts stained green; nuclei stained blue. (d) CD68 immunohistochemical stain; ×20; macrophages if present would be stained green. In all cases, the polymer layer is in the upper left section of the image (visible in part a; depicted by arrow).

muscle actin expression, characteristic of myofibroblast cells (Figure 9c). A similar result has been reported for peritoneal implants of various modified PLA scaffolds.56 When tubules of such tissue have been transplanted into arteries as autologous grafts, they develop over 1−2 months many of the characteristics of mature blood vessels, with the constituent myofibroblasts “differentiating”, complete with contractile myofilaments and associated proteins.3,57 There was little evidence for the presence of macrophages after the 3 weeks of the study, determined by CD68 immunohistochemical staining (Figure 9d), and this along with the α-smooth muscle actin staining is further evidence for the differentiation of the inflammatory cells. An abundance of collagen, particularly in the innermost cell layers (Figure 9b), indicated the secretion of extracellular matrix components. Elastin expression was not detected in any of the tissue samples, as expected for tissue grown under static conditions,3,58 providing further validation of the need for an elastic polymeric scaffold in this application. We have not addressed the in vivo degradation of our materials in the present study since doing so would require removal of the capsule tissue by enzymatic or chemical

tissue graft, since this requires a porous scaffold incorporated into an implantation device, but simply to assess biocompatibility and the inflammatory response to the materials. Samples of the polymers were dip-coated onto polyethylene rods and implanted into the peritoneal cavity of rats for a period of 3 weeks. After this time, tissue capsules were observed on the surface of all samples, which remained free-floating in the peritoneal cavity with no adhesions. Tissue thickness ranged from 0.25 to 1.0 mm, with no apparent trend in relation to TMC content. Figure 9 shows histological and immunohistochemical staining of cross sections of the tissue capsules growing on polymer 5, which showed the thickest growth. These images are representative of the cellular composition for all samples. The staining of samples for the other polymers is shown in the Supporting Information (Figures S3−S5). In Figure 9a, the polymer layer can be seen as a transparent section in the upper left corner of the image. Capsule tissue consisting of tightly packed, elongated cells can be seen adhered to the polymer surface, whereas more diffuse and more rounded cells are visible in the outer tissue layers. Cells throughout the matrix, particularly in the inner layers, were positive for α-smooth 3867


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digestion which may accelerate polymer degradation, thus masking true in vivo degradation effects. The in vivo degradation of PTMC and PDLLA networks has been recently studied by Jansen et al., by implantation into the rabbit eye vitreous.59 No significant mass loss was observed, although the vitreous environment is virtually acellular in comparison to the peritoneal cavity.

work was also supported by the Australian Research Council (ARC) through ARC Linkage Infrastructure Grants LE0775684 and LE0668517 and Discovery Grant DP0987407. The authors thank Dr Joris Rottmans for help in the design of the in vivo study.

4. CONCLUSIONS We have evaluated a series of cross-linked PLLA−TMC copolymers as elastic scaffold materials for an application involving the growth of autologous vascular graft tissue in the peritoneal cavity. It is the first time that synthetic elastomers have been assessed as supports for this method of tissue generation. This study represents an extension over previous work on TMC-based copolymers as we have carried out comprehensive mechanical analysis under periodic and continuous cyclic conditions and during degradation. The polymer films showed “wet” Young’s moduli in the range 1.5−1.8 MPa at 37 °C, approaching the moduli of human arteries, which is of crucial importance for a vascular graft. These values are expected to be reduced further when the polymers are made into porous structures. The TMC content had a large effect on the elasticity and resilience of the materials, with the higher TMC materials showing superior properties. The 56 and 70% TMC polymers showed constant tensile modulus and no plastic deformation after 10 000 continuous mechanical cycles, and this is a significant improvement over other similar materials. The 56 and 70% TMC polymers, while losing around 10% of their mass, retained much of their elasticity and resilience during degradation, but the modulus and ultimate strength decreased slowly over 84 days. This degradation profile will enable remodeling vascular tissue to gradually take on the load-bearing role of the polymer. All the polymer films showed good cell adhesion and proliferation properties using human MSCs. When implanted into the peritoneal cavity of rats, the polymers induced a foreign body response, resulting in myofibroblast tissue capsule growth, showing that they will support growth of artificial vascular tissue using the peritoneal bioreactor model. Given the results, the polymers with TMC content of 56 and 70% in particular are promising candidates for further study as scaffolds for many types of soft tissue engineering.

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ASSOCIATED CONTENT

S Supporting Information *

Detailed quantitative mechanical analysis; confocal and SEM images of all time points for all polymers for the in vitro cell study; and immunohistochemistry conjugate controls and tissue staining for polymers not shown in the text. This material is available free of charge via the Internet at http://pubs.acs.org.

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AUTHOR INFORMATION Corresponding Author *E-mail: [email protected]. Author Contributions § These authors have contributed equally to this work.

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ACKNOWLEDGMENTS Funding for the International Biomaterials Research Alliance was provided by the Queensland State Government, under the National and International Research Alliance Scheme. This 3868

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