Design of Resorbable Porous Tubular Copolyester Scaffolds for Use

Publication Date (Web): March 30, 2009 ... nonload-bearing tissue engineering applications and were easy to handle without damaging the porous structu...
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Biomacromolecules 2009, 10, 1259–1264

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Design of Resorbable Porous Tubular Copolyester Scaffolds for Use in Nerve Regeneration Peter Plikk, Sofia Målberg, and Ann-Christine Albertsson* Department of Fibre and Polymer Technology, School of Chemical Science and Engineering, Royal Institute of Technology, SE-100 44, Stockholm, Sweden Received January 23, 2009; Revised Manuscript Received February 16, 2009

Copolymers of L,L-lactide (LLA), ε-caprolactone (CL), trimethylene carbonate (TMC), or 1,5-dioxepane-2-one (DXO) were used to design porous tubular scaffolds with various mechanical properties, porosities, and numbers of layers in the tube wall. The mechanical properties of the tubular scaffold types showed good suitability for nerve regeneration and other nonload-bearing tissue engineering applications and were easy to handle without damaging the porous structure. A low stannous 2-ethylhexanoate-to-monomer ratio of 1:10000 did not change the tensile properties of the copolymer tubes significantly compared to those of scaffolds made using a Sn(Oct)2to-monomer ratio of 1:600. The adaptability of the immersion coating and porogen leaching technique was demonstrated by creating tubes with different designs. Tubes with different wall layers were created by varying the immersion solutions, and the ease of altering the porosity, pore shape, and pore size was exemplified by using sodium chloride alone or mixed with poly(ethylene glycol) as porogen.

Introduction Tissue-engineered nerve regeneration of damaged or severed nerves using porous tubular scaffolds is a research area of great importance and potential. The ambition of tissue engineering is to facilitate the regeneration of new tissue in cases where the damage is too large or too complicated to heal naturally or when the standard treatments, for example, transplantation, surgical reconstruction, medical devices, or synthetic prostheses, are inadequate or there is a shortage of supply. Tissue engineering employs the use of biodegradable temporary guides that work as templates and provide support for new tissue formation and are totally resorbed when the need for support has ceased. The regeneration of a severed nerve using tubulization involves fitting and suturing the nerve ends into opposite sides of the tubular scaffold, which thus provides guidance for the nerve to heal and close the gap.1,2 Porous tubes are often fabricated using techniques similar to those used for other types of porous scaffolds. Leaching of porecreating additives (porogens) to create porosity is quite commonly used. Some examples of this technique are immersion coating and leaching of sodium fluoride and sugar,3 solvent casting and leaching of sodium chloride to create porous films, which are rolled into tubes by pressing the slightly dissolved edges together,4 and extrusion into a tubular shape followed by leaching of sodium chloride particles.5 Thermally induced phase separation (freeze-drying) is also used, for example, after immersion coating of a glass mandrel 6,7 or of a film before rolling it into a tube.8 Combinations of different methods are utilized as well, here exemplified by fabricating a two-layer porous tube by immersion coating and phase-separation plus salt leaching9 and a three-layer tube by coating a glass mandrel with a layer of nonwoven polymer and thereafter applying a nonporous layer followed by salt leaching to form the third layer.10 Relatively few commercial nerve regeneration tubes exist on the market today, but one example is a woven tube of * To whom correspondence should be addressed. Tel.: +46-8-790 82 74. Fax: +46-8-20 84 77. E-mail: [email protected].

polyglycolic acid (Neurotube) and another is a collagen-based tube (NeuroGen). Polyglycolic acid has a reasonable degradation time, but it is a stiff material with quite low flexibility. Materials of natural origin may suffer from variations in properties from one batch to another. The materials often do not have the required mechanical strength and need extensive purification and characterization. In addition to the need for sufficiently high porosity, interconnected pores, biodegradation, and biocompatibility, the materials used in nerve regeneration are required to have mechanical properties capable of withstanding not only handling and suturing during surgery but also the patients movements throughout the time of tissue formation. The degradation time has to be reasonably long as well, over 6 months, to provide support until the nerve is strong enough to function alone.11,12 Our group has a long experience of working with synthetic bioresorbable aliphatic polyesters,13-18 a type of material well suited for use in this kind of tissue engineering application due to the possibility of modifying both the properties and the degradation times. For example, copolymerizing 1,5-dioxepane2-one with L,L-lactide will alter the surface properties of the copolymer to make it more hydrophilic than a PLLA homopolymer.19 We have also previously shown how to tailor the properties of porous scaffolds for a range of different polyester and polyether-ester copolymers when using a solvent casting and salt leaching scaffold fabrication technique. The mechanical properties were for example varied from hard and brittle to soft and highly flexible.20-22 It was shown that when electron beam irradiation was used to sterilize the scaffolds, at a dose of 2.5 MRad, the changes in number average molecular weight were relatively limited, a decrease of approximately 10-30% was seen for the majority of the different copolymer scaffolds.21 The tensile properties also showed relatively small changes after irradiation.22 The aim is to design resorbable porous tubular scaffolds with the ductility and strength required for use in nerve regeneration applications. Using our previous knowledge of tissue engineering to develop the more complex tubular scaffold design and

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materials with precise properties for that particular purpose, we take the next step toward an applicable product. Accurate modeling of the composition of L,L-lactide, ε-caprolactone, 1,5dioxepane-2-one, and trimethylene carbonate monomers will provide copolymers possessing the essential properties inherent in their backbones. Due to the cytotoxicity of stannous 2-ethylhexanoate (Sn(Oct)2),23 the use of a low amount in the polymerization, and thereby a low residual amount in the polymer, will reduce the risk of unwanted body responses when the scaffolds are implanted. Fabrication of tubular scaffolds using our immersion coating and porogen leaching technique also enables easy modification of the tubes, for example, forming different layers as well as altering the tube diameter, length, porosity, and pore size to suit different demands. The material and tubular scaffold properties were determined using nuclear magnetic resonance spectroscopy (NMR), size exclusion chromatography (SEC), differential scanning calorimetry (DSC), scanning electron microscopy (SEM), and tensile testing.

Experimental Section Materials. Toluene (Merck, Germany) was dried over a Na-wire, and diethyl ether (LabScan, Ireland) was dried over molecular sieves before use. Sodium chloride, NaCl (Fischer Chemicals, Germany), was used after agglomerates had been separated by grinding in a mortar. Stannous 2-ethylhexanoate, Sn(Oct)2 (Sigma-Aldrich, Germany), was dried before use. Poly(ethylene glycol) (Aldrich, Germany), poly(caprolactone) (Aldrich, Germany), ethylene glycol (Merck, Germany), chloroform (LabScan, Ireland), hexane (LabScan, Ireland), and methanol (BDH, United Kingdom) were used as received. Monomers. ε-Caprolactone, CL (Aldrich), was dried over calcium hydride for at least 24 h at room temperature and was then distilled under reduced pressure prior to polymerization. L,L-Lactide, LLA (Serva Feinbiochemica, Germany), was purified by recrystallization in dry toluene. The monomer was then dried for 24 h under reduced pressure at room temperature. Trimethylene carbonate (TMC) was supplied by Radi Medical SystemAB, Sweden, and was used as received. 1,5Dioxepane-2-one (DXO) was synthesized through a Bayer-Villiger oxidation as previously reported,24 and the product was then purified by recrystallization from dry ether followed by two distillations under reduced pressure. The monomer was dried over calcium hydride for 24 h prior to the final distillation. All monomers were stored under an inert atmosphere before use. Polymerization Technique. The polymerizations were performed in bulk using stannous 2-ethylhexanoate (Sn(Oct)2) and ethylene glycol as co-initiator. The monomer-to-Sn(Oct)2 ratio was 10000:1 and the monomer-to-ethylene glycol ratio was approximately 600:1. The amounts of monomer, co-initiator, and Sn(Oct)2 were weighed into silanized round-bottom flasks under a nitrogen atmosphere in a drybox (Mbraun MB 150B-G-I). The round-bottom flask was fitted with mechanical stirring and sealed. The polymerizations were started by immersing the flask in a thermostatted oil bath (110 °C for poly(LLAco-CL) and poly(LLA-co-DXO) and 140 °C for poly(LLA-co-TMC)) and allowed to proceed for 72 h. The polymer was precipitated in a mixture of cold hexane and methanol (95:5). The precipitation was performed three times until all monomer, detected with NMR, had been removed. Nuclear Magnetic Resonance (NMR). The chemical compositions of the copolymers and the degree of monomer conversion were determined by 1H NMR spectroscopy, comparing the relative intensities of the polymer peaks originating from the different monomers and the resonance peaks from the monomer and polymer. 1H NMR was obtained using a Bruker Avance DPX-400 nuclear magnetic resonance spectrometer operating at 400.13 MHz. The samples were prepared by dissolving 10 mg of sample in 1 mL of deuterochloroform (CDCl3) in a 5 mm diameter sample tube. Nondeuterated chloroform was used as an internal standard (δ ) 7.26 ppm).

Plikk et al. Size Exclusion Chromatography (SEC). SEC was used to monitor the molecular weights of the polymers after polymerization. The polymers were analyzed with a Waters 717 plus autosampler and a Waters model 515 apparatus equipped with three PLgel 10 µm mixed B columns, 300 × 7.5 mm (Polymer Laboratories, U.K.). Spectra were recorded with a PL-ELS 1000 evaporative light scattering detector (Polymer Laboratories, U.K.). Millenium32 version 3.05.01 software was used to process the data. Chloroform was used as an eluent, at a flow rate of 1.0 mL/min. Polystyrene standards with a narrow molecular weight distribution in the range of 4000-900000 g/mol were used for calibration. Differential Scanning Calorimetry (DSC). The thermal properties of the synthesized polymers were investigated using a DSC (Mettler Toledo DSC 820 module) under a nitrogen atmosphere. To erase the thermal history, the specimens were heated above the melting temperature and then cooled at a rate of 10 °C/min. The second scan was used to record the heat of fusion at a heating rate of 10 °C/min. The melting temperatures, Tm, were noted as the maximum values of the melting peaks and the midpoint temperature of the glass transition was determined as the glass transition temperature, Tg. When evaluating the crystallinity of the copolymers, it was assumed that the only contribution to the heat of fusion in the poly(DXO-co-LLA) was from the poly(L-lactide). Poly(DXO) has earlier been shown to be a fully amorphous polymer.25 The crystallinity, Xc, was calculated according to26

Xc )

∆Hf ∆Hf0

(1)

where ∆Hf (J/g of the crystalline polymer) is the enthalpy of fusion of the specimen and ∆Hf0 is the enthalpy of fusion of a 100% crystalline polymer. For PCL and PLLA ∆H0f are 139.527 and 93 J/g,28 respectively. Scanning Electron Microscopy (SEM). Two samples were randomly selected for each porous tube. The surface and cross section characteristics were evaluated by the means of a JEOL JSM-5400 SEM. The samples were mounted on metal studs and sputter-coated with gold-palladium (60/40%) using a Denton Vacuum Desk II cold sputter etch unit. Samples were also analyzed using a Hitachi S-4300 FE-SEM. The samples were mounted on metal studs and sputter-coated with gold-palladium using an Agar HR Sputter Coater. Tensile Testing. The tensile testing of the tubes was performed with a computer controlled Instron 5566 equipped with pneumatic grip. The computer program used was Bluehill with the settings for tubular testing. The tensile measurements were performed with a load cell with a maximum of 0.1 kN at a crosshead speed of 50 mm/min and a gauge length of 10 mm. Five different samples of the same tubes were tested for each copolymer. The average outer diameter of each tube was calculated from five different measurements with a micrometer. The inner diameter was constant at 1.5 mm. The samples were preconditioned before testing according to ASTM D618-96 (40 h at 50 ( 5% relative humidity and 23 ( 1 °C). All samples kept their cylindrical shape in the segment between the grips and all samples broke in the cylindrical area (not close to the grips).

Results and Discussion Copolymers of poly(LLA-co-CL), poly(LLA-co-DXO), and poly(LLA-co-TMC), consisting of approximately 80 mol % of LLA and 20 mol % of the second monomer, were synthesized for use in the fabrication of porous tubular scaffolds. Sodium chloride as well as poly(ethylene glycol) (PEG) was used as porogens in an immersion coating and porogen leaching technique. Polymer Composition and Molecular Weight. The bulk polymerizations were performed using a minute amount of stannous 2-ethylhexanoate to minimize the residual tin content

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Table 1. Copolymer Composition and Molecular Weight monomer molar ratio in feed [mol %] copolymer

reaction temp. [°C]

LLA

CL

poly(LLA-co-DXO) poly(LLA-co-CL) poly(LLA-co-TMC)

110 110 140

60 50 70

50

DXO

polymer composition [mol %]a

TMC

LLA

CL 25

30

82 75 79

40

DXO

TMC

Mnb

PDIb

21

137600 84600 85800

1.29 1.22 1.37

18

Determined by 1H NMR in CDCl3 at δLLA ) 5.17 ppm, δDXO ) 3.65 ppm, δCL ) 2.30 ppm, and δTMC ) 2.16 ppm. b Determined by SEC using polystyrene standards and CHCl3 as eluent. a

in the copolymers, according to work previously performed by our group.29 This approach was successful in all the copolymerizations, yielding copolymers of the sought-after composition and molecular weight (see Table 1). Compared to previously performed polymerizations containing higher concentrations of Sn(Oct)220 the polymerization time had to be increased considerably, from 10 to 72 h, due to the considerably slower conversion. The initial monomer feed ratio was 60 mol % LLA and 40 mol % DXO for the poly(LLA-co-DXO) copolymer and 50 mol % of each monomer for poly(LLA-co-CL), the polymerization temperature was 110 °C. To achieve an efficient copolymerization of LLA and TMC in a reasonable reaction time, the temperature was set to 140 °C and the initial monomer feed ratio was set to 70 mol % LLA and 30 mol % TMC. At a temperature of 110 °C and similar initial monomer feed ratios, TMC was incorporated into the chain to a low extent and after 72 h the copolymer contained only a small percentage of TMC. Increasing the feed ratio of TMC to 40% did not increase the content of TMC significantly. Ruckstein et al.30 showed that a polymerization time of 90 h was needed to reach a composition of approximately 20 mol % TMC in the copolymer when using a 500:1 molar ratio of monomer to Sn(Oct)2. Here, a reaction temperature of 140 °C was chosen because the ratio of Sn(Oct)2 was 20 times smaller, thereby slowing down the conversion and consequently yielding a considerable increase in reaction time. Thermal Properties. The glass transition temperatures, Tg, were 19, 10, and 34 °C for poly(LLA-co-DXO), poly(LLA-coCL), and poly(LLA-co-TMC), respectively. The melting temperatures of both poly(LLA-co-DXO), 148 °C, and poly(LLAco-CL), 147 °C, were closer to the melting temperatures of the PLLA homopolymer, indicating that the crystals are formed mainly from LLA segments. A simplified calculation of the crystallinity was therefore performed based only on the ∆Hf0 of PLLA. The poly(LLA-co-DXO) copolymer had a weight crystallinity (Wc) of approximately 25% and the poly(LLA-coCL) copolymer showed a low Wc of about 15%. The poly(LLAco-TMC) copolymer was shown to be amorphous after the repeated heating and cooling. However, in the first heating scan, poly(LLA-co-TMC) showed a slight degree of crystallinity, probably due to the result of crystal formation during storage, but when cooled at a cooling rate of 10 °C/min, the polymer was not able to crystallize. Tubular Scaffold Preparation. The porous tubular scaffolds were prepared by a solvent casting and salt leaching technique as described earlier.20 The technique was, however, modified to suit the fabrication of tubes. A solution containing chloroform, polymer, and sodium chloride (weight ratio of 1:10) was kept under stirring, and a glass mandrel was immersed repeatedly into the solution until the desired thickness/diameter was reached. The polymer/salt composite formed was allowed to dry before it was immersed in methanol to loosen it from the mandrel. The tubular scaffold was thereafter cut to the desired length and the salt was leached out in deionized water for four days. The water was changed according to the previously described technique,20 and the tubes were subsequently dried

in a vacuum oven. To create tubes containing different layers, the glass mandrels were immersed successively in different polymer solutions. Tubes with a nonporous inner film were first immersed in a salt-free polymer solution, dried, and subsequently immersed in a solution containing salt particles. The inner diameter of the tubes was set to 1.5 mm and the outer diameter was chosen to be 3 mm. The tubes were cut to a length of 20 mm. Two different types of tubes were prepared from each of the copolymers, one containing a nonporous inner film and another without the inner film layer. Figure 1 shows that the tubes had pores that were homogeneously distributed within the scaffolds. The surfaces also had an open pore structure, see Figure 1. The porosity of the tubes (without a homogeneous inner layer) was 93-94% for all copolymers, determined as

(

porosity ) 1 -

)

F ·100 Fh

(2)

where F is the density of the porous tube and Fh is the density of a homogeneous tube fabricated in a similar manner to the porous tube. The measurements of weight and volume were performed after salt leaching. The tubes with a homogeneous inner film showed a decrease in porosity of a few percent when using this calculation method. The porosity of the porous outer layer was, however, unchanged; the decrease in porosity was due to an increase in density originating from the homogeneous film. The homogeneous films on the inner surface of the tubes were shown to be very thin, only a few micrometers thick, and practically without pores, as expected (Figure 2). Tubes with a length of 5 mm, an inner diameter of 6 mm, and an outer diameter of 10 mm were also successfully prepared from PCL (Aldrich, Mn 80000 g/mol) to show that the tubular dimension can be altered freely to suit the application, see Figure 3. The versatility of the fabrication technique was also exemplified here by adding PEG (30 wt %, molecular weight of 1000 g/mol) as a porogen, in addition to NaCl at a weight ratio of 10:1 to PCL, to further increase the porosity. PEG forms separated microphases in the continuous PCL-phase, is water soluble, and dissolves in the same process as the salt leaching. By phase separation, the PEG thus forms pores inside the PCL phase, complementing the larger pores created by the salt particles. The pores created by phase separation of PCL and PEG are generally smaller than the pores created by the larger salt particles. Figure 4 is presented as an example to show the pore size from pores formed in a PCL scaffold created using only PEG as porogen. When the immersion coating and salt leaching technique is used, it is also possible to enhance the surface porosity by coating the semisolidified salt/polymer tube with a layer of sodium chloride powder, as shown in Figure 5. The total

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Figure 1. SEM micrographs of (a) poly(LLA-co-CL) cross-section; (b) poly(LLA-co-CL) surface; (c) poly(LLA-co-DXO) cross-section; (d) poly(LLAco-DXO) surface; (e) poly(LLA-co-TMC) cross-section; and (f) poly(LLA-co-TMC) surface.

Figure 3. Digital camera photo of a PCL tube (PEG and NaCl as porogens) with an inner-diameter of 1 cm: (a) cross-section and (b) surface. Figure 2. SEM micrograph of a poly(LLA-co-DXO) porous tube crosssection containing a homogeneous inner film.

porosity (determined using eq 2) of the coated tubes was shown not to increase significantly compared to that of noncoated tubes.

Mechanical Properties. The mechanical properties of tubular scaffolds are important in tissue engineering applications. Even though this type of scaffold does not have mechanical properties suitable for load-bearing applications, the ductility and strength determines how easy it is for the surgeon to handle and suture

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Figure 4. SEM micrographs of porous scaffolds of PCL using PEG as porogen: (a) Mn,PEG ) 1000 g/mol, 40 wt % PCL; (b) Mn,PEG ) 1000 g/mol, 30 wt % PCL.

Figure 5. SEM micrograph of porous poly(LLA-co-CL) tube, with an inner diameter of 1.5 mm, after coating with NaCl powder: (a) surface; (b) surface and cross-section. Table 2. Tensile Properties of Porous Copolymer Tubes copolymer

stress at max load [MPa]

strain at max load [%]

modulus [MPa]

poly(LLA-co-DXO)a poly(LLA-co-DXO)b poly(LLA-co-CL)a poly(LLA-co-CL)b poly(LLA-co-TMC)a poly(LLA-co-TMC)b

0.14 ( 0.02 1.0 ( 0.10 0.14 ( 0.04 0.29 ( 0.08 0.52 ( 0.09 0.51 ( 0.02

33 ( 6.1 55 ( 7.2 72 ( 20 62 ( 26 13 ( 2.2 16 ( 1.4

1.1 ( 0.33 10 ( 2.6 0.44 ( 0.19 1.4 ( 0.35 11 ( 2.9 8.9 ( 2.0

a

Without an inner film layer (single-layer).

b

With an inner film layer.

in the scaffold. A highly brittle material may for example crack or crumble, whereas a ductile material with sufficient strength is possible to implant and suture without destroying the porous structure or tearing the material. Electron beam sterilized tubes of poly(DXO-co-LLA), with a homogeneous inner film, having controlled and smaller pore sizes (>90 µm, created by sodium chloride leaching), have been used in a study on sciatic nerves in rats at the department of surgical sciences section for plastic surgery, Haukelands University Hospital, Bergen, Norway, and the department of experimental plastic surgery, Linko¨ping University, Linko¨ping, Sweden. The tubes were deemed very easy to handle and suture in by the surgeons. The ductility and strength were high enough for the tubes to retain their structure without damage during implantation, and the suturing to anchor the nerve ends was easily performed without tearing or damaging the tubes. The tensile properties of tubular copolymer scaffolds created with sodium chloride as porogen are presented in Table 2. The scaffolds with an inner film are also compared to those without the inner film to show the effect of the nonporous layer on the tensile properties. The stress and strain at maximum load are reported because no yield point was detected. In many cases, the porous tubes were too weak to induce total stop of the load cell. The porous structure probably weakens the device sufficiently for it to break before yielding occurs. Without a homogeneous film layer, the poly(LLA-co-TMC) tubes were shown to be the strongest, although all the copolymer scaffolds have values in the same region. However, the ductilities of the copolymer tubes differ to a greater extent, the

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tubes of poly(LLA-co-CL) and poly(LLA-co-TMC) being the most and the least ductile, respectively. The low extensibility of poly(TMC-co-LLA) might limit its applicability in parts of the body where high flexibility is required. The tensile properties of the tubes were compared to those of scaffolds, with the same weight ratio of NaCl-to-polymer, fabricated from materials synthesized using a Sn(Oct)2-to-monomer ratio of 1:600 with a composition of 77 mol % LLA and 23 mol % DXO or CL, respectively.20 The single-layered poly(LLA-co-DXO) tubes showed slightly lower tensile properties whereas the properties were relatively similar for the single layered poly(LLA-co-CL) tubes. For both copolymers, the tubes with a homogeneous inner film show higher tensile properties than the previously made scaffolds. Compared to scaffolds made of pure PLLA,20 a commonly used material (here chosen for comparison reasons), all the tubes were much more ductile and in many cases also similar in strength or even stronger. The exception was the most ductile poly(LLA-co-CL) tube without a nonporous inner layer. The stiffness of the PLLA scaffolds was in the same range as that of the poly(DXO-co-LLA) tubes with an inner film and of both the poly(LLA-co-TMC) tube types, whereas the poly(LLAco-CL) tubes had much lower stiffness. The strength of both poly(LLA-co-CL) and poly(LLA-co-DXO) tubes was increased by the nonporous layer, while it was practically unchanged for the poly(TMC-co-LLA) tubes, implying that the inner film in the latter case was too thin to have a noticeable impact on the tensile properties. The ductility of the poly(LLA-co-DXO) tubes was also considerably increased by the addition of an inner film.

Conclusions Porous tubular scaffolds having a ductility and strength suitable for nerve-regeneration applications were successfully created from poly(LLA-co-DXO), poly(LLA-co-CL), and poly(LLA-co-TMC) copolymers, synthesized using a ratio of monomer to Sn(Oct)2 of 10000:1. The tubes made from poly(LLA-coCL) and poly(LLA-co-DXO) were more ductile, whereas tubes from poly(LLA-co-TMC) were somewhat stronger and showed a higher stiffness. However, none of the tubes were brittle and all were easy to handle without fracturing. The presence of a homogeneous inner film layer generally increased the tensile properties of the tubes. The tensile properties of the singlelayer poly(LLA-co-DXO) and poly(LLA-co-CL) copolymer tubes were only slightly different from those of scaffolds of similar materials created using a monomer to Sn(Oct)2 ratio of 600:1, which showed that the reduction in Sn(Oct)2 has only a very limited effect on the mechanical properties. The polymerization procedure readily allows copolymers with the required composition, molecular weight and narrow molecular weight distribution to be obtained. All the copolymers had glass-transition temperatures lower than the human body temperature, giving a pliable behavior when implanted. Porosities of around 93% were attained in tubes composed of a single porous copolymer layer, created using an immersion coating and porogen leaching technique with sodium chloride as the porogen. A surface coating of a thin layer of sodium chloride particles was shown to enhance the porous structure of the surface, eliminating possible problems with a homogeneous outer layer formation. The ability to easily alter the scaffold properties and designs was exemplified by producing tubes containing a homogeneous inner layer and adding poly(ethylene glycol) as porogen to alter the pore size and porosity. Acknowledgment. The authors gratefully acknowledge the Swedish Foundation for Strategic Research (Grant No. A302139) for financial support of this work.

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