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Manufacture of porous polymer nerve conduits by a novel low-pressure injection molding process
Biomaterials 24 (2003) 819–830
Manufacture of porous polymer nerve conduits by a novel low-pressure injection molding process Cathryn Sundbacka,*, Tessa Hadlockb, Mack Cheneyb, Joseph Vacantia b
a Department of Surgery, Massachusetts General Hospital, Harvard Medical School, Wellman 625 55 Fruit Street, Boston, MA 02114, USA Department of Otolaryngology, Massachusetts Eye and Ear Infirmary, Harvard Medical School, 243 Charles Street, Boston, MA 02114, USA
Received 1 May 2002; accepted 21 August 2002
Abstract A method to fabricate porous, biodegradable conduits using a combined injection molding, thermally induced phase transition technique was developed which produced conduits with dimensionally toleranced, longitudinally aligned channels. The geometry of the channels was designed to approximate the architecture of peripheral nerves and to support the monolayer adherence of physiologically relevant numbers of Schwann cells. The channel configuration could be varied from a single 1.35 mm diameter channel up to 100 0.08 mm diameter channels. A conduit with 100 channels has approximately 12.5 times the lumenal surface area of a single channel conduit and supports the adherence of five times the number of Schwann cells in the native peripheral nerve. In this study, poly(dl-lactide-co-glycolide) (dl-PLGA) was dissolved in acetic acid and injected into a cold mold which induced solid– liquid phase separation and, ultimately, solidification of the polymer solution. The acetic acid was removed by sublimation and the resulting foam had a macrostructure of high anisotropy. Semi-permeable skins formed on the outer and lumen diameters of the conduit as a consequence of rapid quenching. Macropores were organized into bundles of channels, up to 20 mm wide, in the dlPLGA matrix and represented remnants of acetic acid that crystallized during solidification. r 2002 Elsevier Science Ltd. All rights reserved. Keywords: Peripheral nerve regeneration; Polymer processing; Nerve conduit; Biodegradable polymer; Schwann cells; Tissue engineering
1. Introduction The capacity of axons to regenerate within injured peripheral nerves has long been recognized. Despite microsurgical technique advances and the attainment of cellular and molecular insight into the process of regeneration, functional recovery of the injured peripheral nerve is rarely complete; epineurial suture of autologous nerve grafts remains the most widely used method for repair of peripheral nerve defects [1]. Nerve autografts are not ideal repair conduits, as the function of donor nerves is sacrificed without achievement of full functional recovery of the repaired nerve. Research has focused on the creation of optimal scaffolds to serve as nerve guidance conduits between the proximal and distal nerve stumps ends. These guidance conduits must direct axons sprouting from *Corresponding author. Tel.: +1-617-726-4598; fax: +1-208-3615085. E-mail address: [email protected] (C. Sundback).
the proximal regenerating nerve end, provide a conduit for diffusion of neurotropic and neurotrophic factors secreted by the damaged nerve stump, minimize the infiltration of fibrous tissue, retain adequate mechanical strength and flexibility to support the regenerating nerve fibers, and be biocompatible and biodegradable, so as to be integrated into the surrounding tissue after complete regeneration [2]. Many studies have focused on devising nerve guidance conduits from natural, biological materials. Improved regeneration with respect to controls has been demonstrated from autogenous materials such as skeletal muscle basal lamina grafts, with complex basal lamina architecture, and vein grafts [3–7]. Regeneration has been modest through these conduits, and harvesting these materials has resulted in donor site morbidity. Several biodegradable synthetic materials, in the form of simple hollow conduits, have been shown to support nerve regeneration. Polyesters, such as polylactic acid (PLA), polyglycolic acid (PGA), and PLGA [8–10] have been used extensively because of their availability, ease
0142-9612/02/$ - see front matter r 2002 Elsevier Science Ltd. All rights reserved. PII: S 0 1 4 2 - 9 6 1 2 ( 0 2 ) 0 0 4 0 9 - X
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of processing, approval by the FDA, and low inflammatory response. Other biodegradable polyesters have also demonstrated promise for nerve regeneration applications, such as poly(lactide-e-caprolactone) [11–13], biodegradable polyurethanes [14], poly(organo)phosphazenes [15], and, most recently, trimethylene carbonate-co-e-caprolactone [16–18]. In addition to biodegradability and biocompatibility issues, the influence of a variety of physical parameters of the conduits has been examined. Nerve regeneration has been enhanced with respect to controls by proper choice of hollow-tube characteristics, such as conduit diameter and length [19], lumenal surface microgeometry [20,21], and wall porosity and permeability [22–25]. Incorporation of an oriented intraluminal framework into the conduit lumen, which qualitatively models biological autogenous materials, has been a significant advancement; these conduits have demonstrated superior regeneration because the intraluminal matrix supports cell attachment/migration and guides regenerating axons. A variety of aligned matrix strategies have been investigated, such as: autologous vein conduits or synthetic conduits with inserted acellular muscle [26,27], macroporous synthetic foams with oriented or interconnected pores [28], synthetic or collagen-based filaments within or without conduits [29–34], collagen sponges within synthetic conduits [34], magnetically aligned collagen gels [35], and silicon with micromachined holes of various diameters and spacings [36–38]. Permeable hollow wall conduits have been produced using a variety of processing technologies. Porous mesh or foam sheets have been rolled and sealed, either by welding [39], suturing [40–42], glueing [43,44], joining with solvent [45] or bonding with polymer solution [46]. Porous biodegradable hollow tube conduits have also been manufactured by dip-coating a mandrel into a polymer suspension containing water-soluble porogen particles, such as sugar or sodium chloride, which are leached out post-processing [47–49]. Conduits have been formed by extrusion of a polymer/salt composite [50] or by immersion precipitation, where a polymer nonsolvent effects phase separation of a polymer solution, followed by subsequent gelation to immobilize the microporous structure [51]. In addition to conduit modifications, manipulating the internal environment of nerve conduits has improved nerve regeneration by the supplementation of extracellular matrix molecules, growth factors, and Schwann cells [52,53]. Schwann cells serve as a living source of neurotropic and neurotrophic factors for the regenerating nerve, excrete extracellular matrix, and act as a substrate for elongating axons. Enhanced regeneration has been observed when Schwann cells were seeded into nerve conduits [54–62], largely as suspended Schwann cells. However, axonal growth is guided by Schwann cell columns in the Bungner . bands during
regeneration. To mimic this configuration, several studies have examined the impact of presenting Schwann cells as adherent monolayers, rather than as suspended cells. In most cases, the Schwann cells were attached only to the lumenal surface of the hollow tube nerve conduit, so that few Schwann cells were introduced [56,59,62], and the effect on regeneration was minimal. However, Shen et al. [32] recently described the design of a nerve conduit in which Schwann cells were presented adherent to matrigel coated Vicryl and polydioxanone filaments inserted into hollow tube conduits; Schwann cell chains were observed in structures similar to Bungner . bands. The objective of this study was to manufacture and characterize porous biodegradable polymer nerve conduits for guided tissue regeneration. These conduits contained longitudinally aligned channels and mimicked the geometry of autografts. These conduits were produced using a novel low pressure injection modeling process [63–65] which comprises a thermally induced phase transition (TIPS) process appropriate for the production of complex-shaped geometries. The mold design was flexible and allowed for the inexpensive production of variable numbers and diameters of channels, ranging from one 1.35 mm diameter channel to 100 0.08 mm channels. The channels provided an increased surface area for Schwann cell migration and adherence, as well as for axonal elongation. The chosen polymer material was a high molecular weight 85:15 poly(dl-lactide-co-glycolide) (dl-PLGA). This material was selected as a model polymer system for demonstrating molding process feasibility. The process could be easily translated into other materials such as poly(l-lactide) or poly(caprolactone) polymers or poly(lactide-e-caprolactone) and poly(trimethylene carbonate-co-e-caprolactone) copolymers.
2. Materials and methods 2.1. Materials Poly(d,l-lactide-co-glycolide) in an 85:15 monomer ratio was supplied by Birmingham Polymers (Birmingham, AL). The intrinsic viscosity (IV) and molecular weight of the polymer were measured to be 0.81 and 136,000, respectively. The polymer solvent was 99.99% glacial acetic acid (Aldrich, St. Louis, MO), with a freezing temperature of 16.21C. 2.2. Injection mold A 316L stainless-steel injection mold was fabricated based on the schematic depicted in Fig. 1. The mold consisted of top and bottom halves, aligned by two opposing dowel pins and held in place with two
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setscrews. The mold contained two independent mold cavities, which defined the outer dimension of a length of nerve conduit, each a 50 mm long, 2.3 mm diameter cylinder. A 2.3 mm diameter injection port perpendicularly intersected the nerve conduit cavity; the entrance of the injection port was conically shaped to rapidly align the injector in the mold at the start of an injection cycle. The injection port was designed so that polymer solution inside the conduit froze before the polymer solution in the injection port. The axial channels in the nerve conduit were defined by 316V stainless-steel wires (Small Parts, Inc., Miami Lakes, FL), with the diameter of the wires approximately corresponding to the post-lyophilized lumen channel diameter. The straightness and hardness of the wires was critical to eliminate the possibility that the
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wires touch the mold cavity wall or each other; the former case would allow regenerating axons to extend outside the protective environment of the nerve conduit while the latter case would allow axial channels to intersect. Consequently, 316V stainless wire was selected as it was spring tempered, had a high corrosion resistance, and was ‘‘gun barrel’’ straight. The wires were held in proper radial alignment by a perforated plate or wire mesh screen (Fig. 2). Stainless steel wire mesh (Small Parts, Inc., Miami Lakes, FL) was utilized in conduits with more than seven channels as the mesh and the wires could be tightly size toleranced. For example, for a 16 channel conduit, 305 mm wire was aligned in mesh with a square opening of 381 mm width, corresponding to a size difference of 76 mm. Brass-perforated plate, machined
Fig. 1. Injection mold for the manufacture of conduits with longitudinally aligned channels: (1) wire support circle slot, (2) injection port, (3) mold cavity, (4) dowel pin, and (5) thermocouple port. Each part is labeled only once although multiple numbers exist for each part. The wires, which define the channels, are not included on this drawing.
Fig. 2. End-view schematic depicting the stainless steel wires supported in perforated plate or wire mesh for the (A) 7-channel, (B) 16-channel, and (C) 45-channel configurations. Outer diameter of wire mesh or perforated plate was 6.35 mm, the diameter of the wire support circle slot in the mold. The dashed circle corresponds to a 2.3 mm diameter circle, the diameter of the mold cavity.
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to specification, was utilized in conduits with seven or fewer channels to maintain a tight tolerance between wire and support plate and to minimize the distance between hole openings. For example, for a seven channel conduit, 558 mm wire was aligned in a machine-perforated plate with round 609 mm holes, corresponding to a size difference of 51 mm. The outer diameter of the perforated plates was machined to size. However, the wire mesh circles were punched out with a hand punch (McMaster Carr Supply Company, New Brunswick, NJ).
During mold assembly, the wires which defined the axial channels were threaded under microscope guidance through two wire support circles, either wire mesh or perforated plate, which were aligned and touching. After threading all the wires, the wire support circles were separated and each press fit into slots in the mold bottom half, at either end of a cylindrical mold cavity; two assemblies were inserted into the mold as two conduit lengths were molded during each injection cycle. The wire support circles were held tightly in place when the mold halves were closed, thus ensuring proper
Fig. 3. Complete assembly of injection molding tool for the manufacture of conduits with longitudinally aligned channels: (1) machined brass halfcylinder clamps with Viton sheeting on clamping surfaces, (2) machined brass cylindrical spacer, (3) wire support circles, (4) stainless steel wires which define the conduit channels, (5) injection mold, and (6) compressed spring. Each part is labeled only once although multiple numbers exist for each part. For simplicity, two stainless steel wires are depicted to define the conduit channels. However the wire number could be varied from one to 100 in this mold assembly by varying the wire support circles and the diameter of the wires.
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radial alignment of the wires in the cylindrical mold cavities. The wires used to define the conduit axial channels were held taut and straight during the injection process, using a spring-loaded mechanism (Fig. 3). A machined brass cylindrical spacer was placed on each end of the mold so that the wires protruded through the middle of the spacer. For each set of wires, a spring was inserted into the lumen of one spacer and held in a compressed state by opposing machined brass halfcylinder clamps, with Viton sheeting on the clamping surfaces.
2.3. Polymer foam processing Polymer foam conduits were prepared using a novel injection molding technique. On the molding day, a 10% solution (wt/wt) of 85:15 dl-PLGA was dissolved in glacial acetic acid by stirring at room temperature. The mold was assembled, as described above, and inserted into a box of dry ice, at a temperature of 781C. The mold temperature was cooled to 401C, significantly below the freezing temperature of the polymer solution. At an initial mold temperature of 401C, polymer solution injected into the cold mold completely froze in the mold cavity prior to complete solidification of the polymer solution in the injection port. Room temperature polymer solution was injected into this prechilled mold. A glass syringe was filled with 10 ml of polymer solution. The mold was removed from the dry ice box and clamped into a vise so that injection ports were parallel to the floor. The pre-filled glass syringe was seated into one of the conically shaped injection ports and held tightly by hand against the mold; the injection port entrance had been machined to mate tightly to the syringe. The polymer solution was injected into the cold mold. During early injection, the air inside the mold cavity was expelled as some liquid polymer solution escaped from the mold in the gaps between the wires and the wire supports. Leaking of the polymer solution quickly abated as the polymer solution froze; the injection pressure was maintained for 12 min after leakage stopped so that liquid polymer solution could fill the central shrinkage void that was created when the polymer solution solidified. Shrinkage unavoidably occurs as result of volume contraction of the solvent during solidification. However, by maintaining the injection pressure throughout the injection process, shrinkage defect-free, tight size-tolerance nerve conduits could be produced with the optimal design of the mold injection port and selection of the injection temperature which targeted the polymer solution in the injection port to solidify last. After the first injection was completed, the process was quickly repeated in the second mold cavity. The mold with the frozen polymer solutions was
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returned to the dry ice box and cooled until the mold temperature returned to 401C. The frozen polymeric nerve conduits were demolded, prior to lyophilization. The mold was again clamped in the vise and the wire tension system removed. The wires were individually removed by pulling straight out with pliers. The mold halves were separated and each nerve conduit was lifted out of the mold by the frozen injection port, to avoid melting the part. The mold cavity surfaces were brushed with a Kimwipe tissue to remove frost and debris. New wire support assemblies were inserted into the mold; the mold was assembled and was returned to dry ice for the next cycle. The demolded frozen parts were stored on a 25 mm thick aluminum block held on dry ice until the molding batch was completed. The frozen conduits were lyophilized to sublimate and remove the solid acetic acid, locking the frozen structure in-place. The frozen nerve conduits, supported on the dry ice-chilled block, were placed inside a lyophilization flask. The flask was attached to a liquid nitrogen trap, which in turn was attached to the freeze drier (Labconco Corporation, Kansas City, MO); the liquid nitrogen trap confined all sublimated acetic acid inside the trap and prevented damage to the freeze drier. The nerve conduits were lyophilized under 30–40 mTorr vacuum for at least 72 h. After lyophilization, the nerve conduits were stored in a dry environment until use. A sharp scalpel blade was used to remove the injection port material and to section the conduit to the desired lengths. 2.4. Scanning electron microscopy Fracture surfaces were prepared by fracturing samples after immersion in liquid nitrogen. As molded surfaces and fracture surfaces were coated with gold/palladium ( using a high at an approximate thickness of 300 A, resolution ion beam coater (Gatan Model 681, Pleasanton, CA). Coated surfaces were viewed with a scanning electron microscope (FEI/Philips XL30 ESEM-FEG, Hillsboro, OR), operated at 2 kV accelerating voltage. 2.5. Image analysis Average dimensions were estimated based on measurements on five representative SEM micrographs made using the image analysis component of the Metamorph Imaging System (Universal Imaging Corporation, Downingtown PA). 2.6. NMR The lyophilized PLGA materials were dissolved in deuterated chloroform (CDCl3). 1 H NMR spectra were
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collected on a 400 MHz FT-NMR instrument (Bruker Advanced DPX 400 model, Billerica, MA).
3. Results and discussion Porous biodegradable nerve conduits were manufactured with a process where the number of axially aligned channels could easily be varied. These axial channels provided an oriented intraluminal framework which has been shown to enhance regeneration [22–34] and a configuration which mimicked the configuration of Bungner . bands, the Schwann cells columns which promote axonal growth during regeneration. The nerve conduits were produced using a modified TIPS. Traditionally, TIPS is practiced in the context of solvent casting in which the polymer solution is cast on a surface and solid–liquid phase separation is induced by lowering the solution temperature. Subsequent removal of the solidified solvent-rich phase by sublimation leaves a porous polymer scaffold. Our process modified traditional TIPS processing so that room temperature polymer solution was injected under low pressure into a cold mold. The low injection pressure of this manual process was applied and maintained from initial injection through complete solidification to allow for complete filling of complexshaped molds and to avoid the formation of a shrinkage void, related to the unavoidable density change of the solvent during solidification. As described above, the mold was designed so that polymer solution in the injection port was the last to solidify. Consequently, polymer solution injection could be maintained to fill any voids created during polymer solution solidification. Solvent choice was critical to successful processing. Many of the selection criteria were similar to those of other TIPS processes. Potential solvents must possess relatively high melting points to avoid melting during lyophilization, high vapor pressures to allow facile removal by sublimation, and significant polymer solubility [66]. In addition, the solvent must not be carcinogenic for in vivo applications [67] and the solvent must not decrease the activity of bioactive molecules incorporated into the polymeric matrices [68]. Finally, low volume change upon solidification is desired to minimize the driving force to form shrinkage voids upon solidification; however, little volume change data has been published. Other TIPS processes have employed solvents like benzene, naphthalene, and dioxane that have acceptable melting temperatures and vapor pressures. However, all these solvents are demonstrated carcinogens. Thus, we chose acetic acid, which is generated during physiological metabolic processes. Only a residual amount of acetic acid remained in the lyophilized PLGA foam, as estimated by 1H NMR. After 72 h of lyophilization, the
PLGA foam contained less than 0.01% acetic acid estimated by comparing the peak integration values of the methyl group from acetic acid with that of the lactide segments of PLGA. SEM micrographs of transverse fracture cross sections of 7-channel and single channel dl-PLGA (85:15) nerve conduits are shown in Fig. 4. The channels and the conduits are generally circular and the channels are wellaligned within the outer diameter of the conduit. In both cases, minimal distortion occurred at the mold sections interface line. From measurements on representative SEM micrographs, minimal shrinkage of the injection molded conduit occurred, as the measured conduit dimensions are at most 5% smaller than the original mold dimension. As little dimensional change occurred during processing, the conduit porosity is approximately 90%, which corresponds to the volume percent of solvent in the original polymer solution. Higher magnification SEM micrographs are shown in Fig. 5 of the transverse cross sections of 7-channel conduits within a section between two internal channels
Fig. 4. Scanning electron micrographs of transverse cross sections of (A) 7-channel and (B) single channel conduits. Bar is 500 mm.
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Fig. 5. Scanning electron micrographs of transverse cross sections of 7-channel conduits, depicting (A) (B) the bulk macrostructure including semipermeable lumen skins and (C) (D) the outer diameter including semi-permeable outer skin. Bars are 100 mm in (A) and (C) and 50 mm in (B) and (D).
Fig. 6. Scanning electron micrograph of transverse cross section of 7-channel conduit, depicting porosity which consisted of channels, 1–20 mm wide, between parallel, relatively nonporous, polymeric lamellae. A 3-mm thick semi-permeable skin on a lumen surface is shown in the upper left-hand corner. Bar is 10 mm.
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and at the outer diameter. A semi-permeable skin existed in contact with the mold surfaces, with a very thin 3 mm skin on the inner diameter of the lumens and approximately a 25 mm skin on the outer diameter. An anisotropic sheet-like porous structure existed throughout the bulk macrostructure, outside the less permeable skin areas. This porosity consisted of channels, 1–20 mm wide, between parallel, relatively nonporous, polymeric lamellae (Fig. 6). Similar semi-permeable skins and bulk macrostructures are seen in single channel longitudinal sections (Fig. 7). The semi-permeable outer diameter and lumen skins of the single channel conduit are approximately similar in width at 25 mm. The bulk macrostructure again consisted of a lamellar structure with the porous channels perpendicularly oriented to the outer and inner mold surfaces and parallel to the cooling direction. Enhanced peripheral nerve regeneration has been demonstrated through highly permeable nerve conduits relative to regeneration through impermeable conduits [12]. The solid–liquid phase separation is responsible for the characteristic features of the foam morphology. Progress of the crystallization front of acetic acid dictates the main orientation of the pores, the long axes of which are parallel to the cooling direction. The polymer is rejected from the solvent crystal front with the formation of parallel radial sheets. High concentration solutions and fast quenching rates favor this morphology, because only limited polymer reorganization can occur before being ultimately frozen [67]. Because of the long aspect ratio of the nerve conduit and the mass of the mold, the cooling front moves radially from the outer diameter of the conduit toward the center of the conduit. This directional cooling is clearing seen in the bulk macrostructure alignment of the single channel conduit. However, the porosity alignment is perturbed in the 7-channel conduit because of the cooling perturbations effected by the wires. Additional microstructure perturbations are created by the injection process, which occurs simultaneously to the polymer solution cooling and liquid–solid phase separation. The semi-permeable skins form as a result of rapid quenching of the polymer solution as it initially contacts and solidifies on the mold surfaces. The skin thicknesses are a function of the solution cooling rate, and hence, the initial temperature and thermal mass of the mold surface adjacent to the polymer solution. The outer diameter skins of the single channel and 7-channel conduits are both approximately 25 mm thick, since the same mold and similar molding temperatures were used for both. However, the lumen skin of the single channel conduit is approximately 25 mm, significantly thicker than the 3 mm skin of the 7-channel conduit; the mass of the 1.35 mm diameter wire used for the single channel
Fig. 7. Scanning electron micrographs of longitudinal cross sections of single-channel conduit depicting (A) full cross section and lumen surface and (B) (C) increasing magnifications of bulk macrostructures and semi-permeable surface skins. Bars are 500, 200, and 100 mm in (A), (B), and (C), respectively.
conduit generates a significantly greater cooling rate relative to the 0.51 mm diameter wire of the 7-channel conduit.
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Fig. 8. Scanning electron micrographs of (A) (B) the lumen surface and (C) (D) the outer diameter surface of a single channel conduit. Bars are 100 mm for (A) and (C) and 5 mm for (B) and (D).
The inner lumenal and outer diameter surfaces of the single channel conduit are relatively smooth (Fig. 8). Some parallel shallow grooves existed on the lumenal surface, which were created as the wire was demolded and scraped the frozen polymer surface. However, the surface is generally smooth with submicron porosity. Lumenal surface roughness is known to affect peripheral nerve regeneration with smoother surfaces demonstrating enhanced regeneration [20]. The outer surface is rougher than the lumenal surfaces and appears relatively nonporous. The parallel aligned structures on the outer surface are mirror images of the machining marks on the mold surface. Tailoring the solvent system could easily modify the microstructures. Replacing the acetic acid with more typical TIPS solvents, such as 1,4-dioxane or benzene, should produce similar macrostructures [67,69] because a solid-liquid phase separation would occur with a crystallizable solvent. However, a microcellular architecture could be formed by the addition of a nonsolvent; i.e. the combined dioxane/water system. Instead of inducing a solid–liquid phase separation upon
cooling, a liquid–liquid phase separation would be induced which gives rise to an isotropic continuous cellular porosity [69], originating from the interconnected separated phases. Thus the solvent system choice could provide flexibility in the design of the porous matrix. Using this process, the number and diameters of the channels can easily and inexpensively be varied. These longitudinally aligned channels provide a framework which guides the regenerating axons but also provides a support for the adherence of physiologically relevant Schwann cells. Approximately 2.5 times more Schwann cells can be adhered in a monolayer to the lumen surfaces of a 7-channel conduit in comparison to a single channel conduit; the surface area of the 7 channel conduit would allow 3500 Schwann cells/mm to adhere in monolayer fashion, approximately the Schwann cell number in a healthy native peripheral nerve. In a 100channel conduit, the lumenal surface is 12.5 times that of a single channel conduit; approximately 14,000 Schwann cells/mm could be adhered, five times that of a healthy nerve.
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In vivo studies were conducted using the nerve conduits described above in which adherent Schwann cells were introduced or a model neurotrophin was encapsulated in the porous conduit walls and was slowly released over a period of weeks to months. In the first study, 7-channel conduits were seeded with Schwann cells and the adherent Schwann cell/conduit composites were implanted into 7 mm gap defects in rat sciatic nerves. Early regeneration results compared favorably with autografts [63]. In the second study, inosine, a purine analog thought to possess neurotrophic properties, was encapsulated in the polymer solution and single channel conduits were produced using the process described above. Inosine-loaded conduits were implanted into 7 mm gap defects in rat sciatic nerves. Again, early regeneration results suggest that enhanced regeneration occurred through the inosine-loaded conduits in comparison with inosine-free conduits [65].
4. Conclusions A method to fabricate porous, biodegradable conduits using a combined injection molding, thermally induced phase transition (TIPS) technique was developed which produced conduits with dimensionally toleranced, longitudinally aligned channels; the geometry of the channels was designed to approximate the architecture of native peripheral nerves and to promote the monolayer adherence of Schwann cells. As an example, the fabrication method was described as successfully applied using the biodegradable polymer dl-PLGA (85:15) and the solvent acetic acid to fabricate single-channel and 7-channel conduits. However, the process could easily be tailored to other polymer and solvent systems and has been used to produce conduits ranging from a single channel of 1.35 mm diameter to 100 channels of 0.08 mm diameter. As the polymer solution cooled following injection into a cold mold, a solid–liquid phase separation was induced which allowed for the production of macroporous foams of a high anisotropy. Semi-permeable skins formed on the outer diameter and lumen diameters of the conduit as a result of rapid quenching of the polymer solution upon contact with the cold mold surfaces. Macropores were organized into bundles of channels, up to 20 mm wide, in the dl-PLGA matrix and represented remnants of the acetic acid that crystallized during the polymer solution solidification.
Acknowledgements We gratefully acknowledge Dr. Jinming Gao for conducting the NMR analysis.
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Manufacture of porous polymer nerve conduits by a novel low-pressure injection molding process Cathryn Sundbacka,*, Tessa Hadlockb, Mack Cheneyb, Joseph Vacantia b
a Department of Surgery, Massachusetts General Hospital, Harvard Medical School, Wellman 625 55 Fruit Street, Boston, MA 02114, USA Department of Otolaryngology, Massachusetts Eye and Ear Infirmary, Harvard Medical School, 243 Charles Street, Boston, MA 02114, USA
Received 1 May 2002; accepted 21 August 2002
Abstract A method to fabricate porous, biodegradable conduits using a combined injection molding, thermally induced phase transition technique was developed which produced conduits with dimensionally toleranced, longitudinally aligned channels. The geometry of the channels was designed to approximate the architecture of peripheral nerves and to support the monolayer adherence of physiologically relevant numbers of Schwann cells. The channel configuration could be varied from a single 1.35 mm diameter channel up to 100 0.08 mm diameter channels. A conduit with 100 channels has approximately 12.5 times the lumenal surface area of a single channel conduit and supports the adherence of five times the number of Schwann cells in the native peripheral nerve. In this study, poly(dl-lactide-co-glycolide) (dl-PLGA) was dissolved in acetic acid and injected into a cold mold which induced solid– liquid phase separation and, ultimately, solidification of the polymer solution. The acetic acid was removed by sublimation and the resulting foam had a macrostructure of high anisotropy. Semi-permeable skins formed on the outer and lumen diameters of the conduit as a consequence of rapid quenching. Macropores were organized into bundles of channels, up to 20 mm wide, in the dlPLGA matrix and represented remnants of acetic acid that crystallized during solidification. r 2002 Elsevier Science Ltd. All rights reserved. Keywords: Peripheral nerve regeneration; Polymer processing; Nerve conduit; Biodegradable polymer; Schwann cells; Tissue engineering
1. Introduction The capacity of axons to regenerate within injured peripheral nerves has long been recognized. Despite microsurgical technique advances and the attainment of cellular and molecular insight into the process of regeneration, functional recovery of the injured peripheral nerve is rarely complete; epineurial suture of autologous nerve grafts remains the most widely used method for repair of peripheral nerve defects [1]. Nerve autografts are not ideal repair conduits, as the function of donor nerves is sacrificed without achievement of full functional recovery of the repaired nerve. Research has focused on the creation of optimal scaffolds to serve as nerve guidance conduits between the proximal and distal nerve stumps ends. These guidance conduits must direct axons sprouting from *Corresponding author. Tel.: +1-617-726-4598; fax: +1-208-3615085. E-mail address: [email protected] (C. Sundback).
the proximal regenerating nerve end, provide a conduit for diffusion of neurotropic and neurotrophic factors secreted by the damaged nerve stump, minimize the infiltration of fibrous tissue, retain adequate mechanical strength and flexibility to support the regenerating nerve fibers, and be biocompatible and biodegradable, so as to be integrated into the surrounding tissue after complete regeneration [2]. Many studies have focused on devising nerve guidance conduits from natural, biological materials. Improved regeneration with respect to controls has been demonstrated from autogenous materials such as skeletal muscle basal lamina grafts, with complex basal lamina architecture, and vein grafts [3–7]. Regeneration has been modest through these conduits, and harvesting these materials has resulted in donor site morbidity. Several biodegradable synthetic materials, in the form of simple hollow conduits, have been shown to support nerve regeneration. Polyesters, such as polylactic acid (PLA), polyglycolic acid (PGA), and PLGA [8–10] have been used extensively because of their availability, ease
0142-9612/02/$ - see front matter r 2002 Elsevier Science Ltd. All rights reserved. PII: S 0 1 4 2 - 9 6 1 2 ( 0 2 ) 0 0 4 0 9 - X
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of processing, approval by the FDA, and low inflammatory response. Other biodegradable polyesters have also demonstrated promise for nerve regeneration applications, such as poly(lactide-e-caprolactone) [11–13], biodegradable polyurethanes [14], poly(organo)phosphazenes [15], and, most recently, trimethylene carbonate-co-e-caprolactone [16–18]. In addition to biodegradability and biocompatibility issues, the influence of a variety of physical parameters of the conduits has been examined. Nerve regeneration has been enhanced with respect to controls by proper choice of hollow-tube characteristics, such as conduit diameter and length [19], lumenal surface microgeometry [20,21], and wall porosity and permeability [22–25]. Incorporation of an oriented intraluminal framework into the conduit lumen, which qualitatively models biological autogenous materials, has been a significant advancement; these conduits have demonstrated superior regeneration because the intraluminal matrix supports cell attachment/migration and guides regenerating axons. A variety of aligned matrix strategies have been investigated, such as: autologous vein conduits or synthetic conduits with inserted acellular muscle [26,27], macroporous synthetic foams with oriented or interconnected pores [28], synthetic or collagen-based filaments within or without conduits [29–34], collagen sponges within synthetic conduits [34], magnetically aligned collagen gels [35], and silicon with micromachined holes of various diameters and spacings [36–38]. Permeable hollow wall conduits have been produced using a variety of processing technologies. Porous mesh or foam sheets have been rolled and sealed, either by welding [39], suturing [40–42], glueing [43,44], joining with solvent [45] or bonding with polymer solution [46]. Porous biodegradable hollow tube conduits have also been manufactured by dip-coating a mandrel into a polymer suspension containing water-soluble porogen particles, such as sugar or sodium chloride, which are leached out post-processing [47–49]. Conduits have been formed by extrusion of a polymer/salt composite [50] or by immersion precipitation, where a polymer nonsolvent effects phase separation of a polymer solution, followed by subsequent gelation to immobilize the microporous structure [51]. In addition to conduit modifications, manipulating the internal environment of nerve conduits has improved nerve regeneration by the supplementation of extracellular matrix molecules, growth factors, and Schwann cells [52,53]. Schwann cells serve as a living source of neurotropic and neurotrophic factors for the regenerating nerve, excrete extracellular matrix, and act as a substrate for elongating axons. Enhanced regeneration has been observed when Schwann cells were seeded into nerve conduits [54–62], largely as suspended Schwann cells. However, axonal growth is guided by Schwann cell columns in the Bungner . bands during
regeneration. To mimic this configuration, several studies have examined the impact of presenting Schwann cells as adherent monolayers, rather than as suspended cells. In most cases, the Schwann cells were attached only to the lumenal surface of the hollow tube nerve conduit, so that few Schwann cells were introduced [56,59,62], and the effect on regeneration was minimal. However, Shen et al. [32] recently described the design of a nerve conduit in which Schwann cells were presented adherent to matrigel coated Vicryl and polydioxanone filaments inserted into hollow tube conduits; Schwann cell chains were observed in structures similar to Bungner . bands. The objective of this study was to manufacture and characterize porous biodegradable polymer nerve conduits for guided tissue regeneration. These conduits contained longitudinally aligned channels and mimicked the geometry of autografts. These conduits were produced using a novel low pressure injection modeling process [63–65] which comprises a thermally induced phase transition (TIPS) process appropriate for the production of complex-shaped geometries. The mold design was flexible and allowed for the inexpensive production of variable numbers and diameters of channels, ranging from one 1.35 mm diameter channel to 100 0.08 mm channels. The channels provided an increased surface area for Schwann cell migration and adherence, as well as for axonal elongation. The chosen polymer material was a high molecular weight 85:15 poly(dl-lactide-co-glycolide) (dl-PLGA). This material was selected as a model polymer system for demonstrating molding process feasibility. The process could be easily translated into other materials such as poly(l-lactide) or poly(caprolactone) polymers or poly(lactide-e-caprolactone) and poly(trimethylene carbonate-co-e-caprolactone) copolymers.
2. Materials and methods 2.1. Materials Poly(d,l-lactide-co-glycolide) in an 85:15 monomer ratio was supplied by Birmingham Polymers (Birmingham, AL). The intrinsic viscosity (IV) and molecular weight of the polymer were measured to be 0.81 and 136,000, respectively. The polymer solvent was 99.99% glacial acetic acid (Aldrich, St. Louis, MO), with a freezing temperature of 16.21C. 2.2. Injection mold A 316L stainless-steel injection mold was fabricated based on the schematic depicted in Fig. 1. The mold consisted of top and bottom halves, aligned by two opposing dowel pins and held in place with two
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setscrews. The mold contained two independent mold cavities, which defined the outer dimension of a length of nerve conduit, each a 50 mm long, 2.3 mm diameter cylinder. A 2.3 mm diameter injection port perpendicularly intersected the nerve conduit cavity; the entrance of the injection port was conically shaped to rapidly align the injector in the mold at the start of an injection cycle. The injection port was designed so that polymer solution inside the conduit froze before the polymer solution in the injection port. The axial channels in the nerve conduit were defined by 316V stainless-steel wires (Small Parts, Inc., Miami Lakes, FL), with the diameter of the wires approximately corresponding to the post-lyophilized lumen channel diameter. The straightness and hardness of the wires was critical to eliminate the possibility that the
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wires touch the mold cavity wall or each other; the former case would allow regenerating axons to extend outside the protective environment of the nerve conduit while the latter case would allow axial channels to intersect. Consequently, 316V stainless wire was selected as it was spring tempered, had a high corrosion resistance, and was ‘‘gun barrel’’ straight. The wires were held in proper radial alignment by a perforated plate or wire mesh screen (Fig. 2). Stainless steel wire mesh (Small Parts, Inc., Miami Lakes, FL) was utilized in conduits with more than seven channels as the mesh and the wires could be tightly size toleranced. For example, for a 16 channel conduit, 305 mm wire was aligned in mesh with a square opening of 381 mm width, corresponding to a size difference of 76 mm. Brass-perforated plate, machined
Fig. 1. Injection mold for the manufacture of conduits with longitudinally aligned channels: (1) wire support circle slot, (2) injection port, (3) mold cavity, (4) dowel pin, and (5) thermocouple port. Each part is labeled only once although multiple numbers exist for each part. The wires, which define the channels, are not included on this drawing.
Fig. 2. End-view schematic depicting the stainless steel wires supported in perforated plate or wire mesh for the (A) 7-channel, (B) 16-channel, and (C) 45-channel configurations. Outer diameter of wire mesh or perforated plate was 6.35 mm, the diameter of the wire support circle slot in the mold. The dashed circle corresponds to a 2.3 mm diameter circle, the diameter of the mold cavity.
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to specification, was utilized in conduits with seven or fewer channels to maintain a tight tolerance between wire and support plate and to minimize the distance between hole openings. For example, for a seven channel conduit, 558 mm wire was aligned in a machine-perforated plate with round 609 mm holes, corresponding to a size difference of 51 mm. The outer diameter of the perforated plates was machined to size. However, the wire mesh circles were punched out with a hand punch (McMaster Carr Supply Company, New Brunswick, NJ).
During mold assembly, the wires which defined the axial channels were threaded under microscope guidance through two wire support circles, either wire mesh or perforated plate, which were aligned and touching. After threading all the wires, the wire support circles were separated and each press fit into slots in the mold bottom half, at either end of a cylindrical mold cavity; two assemblies were inserted into the mold as two conduit lengths were molded during each injection cycle. The wire support circles were held tightly in place when the mold halves were closed, thus ensuring proper
Fig. 3. Complete assembly of injection molding tool for the manufacture of conduits with longitudinally aligned channels: (1) machined brass halfcylinder clamps with Viton sheeting on clamping surfaces, (2) machined brass cylindrical spacer, (3) wire support circles, (4) stainless steel wires which define the conduit channels, (5) injection mold, and (6) compressed spring. Each part is labeled only once although multiple numbers exist for each part. For simplicity, two stainless steel wires are depicted to define the conduit channels. However the wire number could be varied from one to 100 in this mold assembly by varying the wire support circles and the diameter of the wires.
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radial alignment of the wires in the cylindrical mold cavities. The wires used to define the conduit axial channels were held taut and straight during the injection process, using a spring-loaded mechanism (Fig. 3). A machined brass cylindrical spacer was placed on each end of the mold so that the wires protruded through the middle of the spacer. For each set of wires, a spring was inserted into the lumen of one spacer and held in a compressed state by opposing machined brass halfcylinder clamps, with Viton sheeting on the clamping surfaces.
2.3. Polymer foam processing Polymer foam conduits were prepared using a novel injection molding technique. On the molding day, a 10% solution (wt/wt) of 85:15 dl-PLGA was dissolved in glacial acetic acid by stirring at room temperature. The mold was assembled, as described above, and inserted into a box of dry ice, at a temperature of 781C. The mold temperature was cooled to 401C, significantly below the freezing temperature of the polymer solution. At an initial mold temperature of 401C, polymer solution injected into the cold mold completely froze in the mold cavity prior to complete solidification of the polymer solution in the injection port. Room temperature polymer solution was injected into this prechilled mold. A glass syringe was filled with 10 ml of polymer solution. The mold was removed from the dry ice box and clamped into a vise so that injection ports were parallel to the floor. The pre-filled glass syringe was seated into one of the conically shaped injection ports and held tightly by hand against the mold; the injection port entrance had been machined to mate tightly to the syringe. The polymer solution was injected into the cold mold. During early injection, the air inside the mold cavity was expelled as some liquid polymer solution escaped from the mold in the gaps between the wires and the wire supports. Leaking of the polymer solution quickly abated as the polymer solution froze; the injection pressure was maintained for 12 min after leakage stopped so that liquid polymer solution could fill the central shrinkage void that was created when the polymer solution solidified. Shrinkage unavoidably occurs as result of volume contraction of the solvent during solidification. However, by maintaining the injection pressure throughout the injection process, shrinkage defect-free, tight size-tolerance nerve conduits could be produced with the optimal design of the mold injection port and selection of the injection temperature which targeted the polymer solution in the injection port to solidify last. After the first injection was completed, the process was quickly repeated in the second mold cavity. The mold with the frozen polymer solutions was
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returned to the dry ice box and cooled until the mold temperature returned to 401C. The frozen polymeric nerve conduits were demolded, prior to lyophilization. The mold was again clamped in the vise and the wire tension system removed. The wires were individually removed by pulling straight out with pliers. The mold halves were separated and each nerve conduit was lifted out of the mold by the frozen injection port, to avoid melting the part. The mold cavity surfaces were brushed with a Kimwipe tissue to remove frost and debris. New wire support assemblies were inserted into the mold; the mold was assembled and was returned to dry ice for the next cycle. The demolded frozen parts were stored on a 25 mm thick aluminum block held on dry ice until the molding batch was completed. The frozen conduits were lyophilized to sublimate and remove the solid acetic acid, locking the frozen structure in-place. The frozen nerve conduits, supported on the dry ice-chilled block, were placed inside a lyophilization flask. The flask was attached to a liquid nitrogen trap, which in turn was attached to the freeze drier (Labconco Corporation, Kansas City, MO); the liquid nitrogen trap confined all sublimated acetic acid inside the trap and prevented damage to the freeze drier. The nerve conduits were lyophilized under 30–40 mTorr vacuum for at least 72 h. After lyophilization, the nerve conduits were stored in a dry environment until use. A sharp scalpel blade was used to remove the injection port material and to section the conduit to the desired lengths. 2.4. Scanning electron microscopy Fracture surfaces were prepared by fracturing samples after immersion in liquid nitrogen. As molded surfaces and fracture surfaces were coated with gold/palladium ( using a high at an approximate thickness of 300 A, resolution ion beam coater (Gatan Model 681, Pleasanton, CA). Coated surfaces were viewed with a scanning electron microscope (FEI/Philips XL30 ESEM-FEG, Hillsboro, OR), operated at 2 kV accelerating voltage. 2.5. Image analysis Average dimensions were estimated based on measurements on five representative SEM micrographs made using the image analysis component of the Metamorph Imaging System (Universal Imaging Corporation, Downingtown PA). 2.6. NMR The lyophilized PLGA materials were dissolved in deuterated chloroform (CDCl3). 1 H NMR spectra were
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collected on a 400 MHz FT-NMR instrument (Bruker Advanced DPX 400 model, Billerica, MA).
3. Results and discussion Porous biodegradable nerve conduits were manufactured with a process where the number of axially aligned channels could easily be varied. These axial channels provided an oriented intraluminal framework which has been shown to enhance regeneration [22–34] and a configuration which mimicked the configuration of Bungner . bands, the Schwann cells columns which promote axonal growth during regeneration. The nerve conduits were produced using a modified TIPS. Traditionally, TIPS is practiced in the context of solvent casting in which the polymer solution is cast on a surface and solid–liquid phase separation is induced by lowering the solution temperature. Subsequent removal of the solidified solvent-rich phase by sublimation leaves a porous polymer scaffold. Our process modified traditional TIPS processing so that room temperature polymer solution was injected under low pressure into a cold mold. The low injection pressure of this manual process was applied and maintained from initial injection through complete solidification to allow for complete filling of complexshaped molds and to avoid the formation of a shrinkage void, related to the unavoidable density change of the solvent during solidification. As described above, the mold was designed so that polymer solution in the injection port was the last to solidify. Consequently, polymer solution injection could be maintained to fill any voids created during polymer solution solidification. Solvent choice was critical to successful processing. Many of the selection criteria were similar to those of other TIPS processes. Potential solvents must possess relatively high melting points to avoid melting during lyophilization, high vapor pressures to allow facile removal by sublimation, and significant polymer solubility [66]. In addition, the solvent must not be carcinogenic for in vivo applications [67] and the solvent must not decrease the activity of bioactive molecules incorporated into the polymeric matrices [68]. Finally, low volume change upon solidification is desired to minimize the driving force to form shrinkage voids upon solidification; however, little volume change data has been published. Other TIPS processes have employed solvents like benzene, naphthalene, and dioxane that have acceptable melting temperatures and vapor pressures. However, all these solvents are demonstrated carcinogens. Thus, we chose acetic acid, which is generated during physiological metabolic processes. Only a residual amount of acetic acid remained in the lyophilized PLGA foam, as estimated by 1H NMR. After 72 h of lyophilization, the
PLGA foam contained less than 0.01% acetic acid estimated by comparing the peak integration values of the methyl group from acetic acid with that of the lactide segments of PLGA. SEM micrographs of transverse fracture cross sections of 7-channel and single channel dl-PLGA (85:15) nerve conduits are shown in Fig. 4. The channels and the conduits are generally circular and the channels are wellaligned within the outer diameter of the conduit. In both cases, minimal distortion occurred at the mold sections interface line. From measurements on representative SEM micrographs, minimal shrinkage of the injection molded conduit occurred, as the measured conduit dimensions are at most 5% smaller than the original mold dimension. As little dimensional change occurred during processing, the conduit porosity is approximately 90%, which corresponds to the volume percent of solvent in the original polymer solution. Higher magnification SEM micrographs are shown in Fig. 5 of the transverse cross sections of 7-channel conduits within a section between two internal channels
Fig. 4. Scanning electron micrographs of transverse cross sections of (A) 7-channel and (B) single channel conduits. Bar is 500 mm.
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Fig. 5. Scanning electron micrographs of transverse cross sections of 7-channel conduits, depicting (A) (B) the bulk macrostructure including semipermeable lumen skins and (C) (D) the outer diameter including semi-permeable outer skin. Bars are 100 mm in (A) and (C) and 50 mm in (B) and (D).
Fig. 6. Scanning electron micrograph of transverse cross section of 7-channel conduit, depicting porosity which consisted of channels, 1–20 mm wide, between parallel, relatively nonporous, polymeric lamellae. A 3-mm thick semi-permeable skin on a lumen surface is shown in the upper left-hand corner. Bar is 10 mm.
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and at the outer diameter. A semi-permeable skin existed in contact with the mold surfaces, with a very thin 3 mm skin on the inner diameter of the lumens and approximately a 25 mm skin on the outer diameter. An anisotropic sheet-like porous structure existed throughout the bulk macrostructure, outside the less permeable skin areas. This porosity consisted of channels, 1–20 mm wide, between parallel, relatively nonporous, polymeric lamellae (Fig. 6). Similar semi-permeable skins and bulk macrostructures are seen in single channel longitudinal sections (Fig. 7). The semi-permeable outer diameter and lumen skins of the single channel conduit are approximately similar in width at 25 mm. The bulk macrostructure again consisted of a lamellar structure with the porous channels perpendicularly oriented to the outer and inner mold surfaces and parallel to the cooling direction. Enhanced peripheral nerve regeneration has been demonstrated through highly permeable nerve conduits relative to regeneration through impermeable conduits [12]. The solid–liquid phase separation is responsible for the characteristic features of the foam morphology. Progress of the crystallization front of acetic acid dictates the main orientation of the pores, the long axes of which are parallel to the cooling direction. The polymer is rejected from the solvent crystal front with the formation of parallel radial sheets. High concentration solutions and fast quenching rates favor this morphology, because only limited polymer reorganization can occur before being ultimately frozen [67]. Because of the long aspect ratio of the nerve conduit and the mass of the mold, the cooling front moves radially from the outer diameter of the conduit toward the center of the conduit. This directional cooling is clearing seen in the bulk macrostructure alignment of the single channel conduit. However, the porosity alignment is perturbed in the 7-channel conduit because of the cooling perturbations effected by the wires. Additional microstructure perturbations are created by the injection process, which occurs simultaneously to the polymer solution cooling and liquid–solid phase separation. The semi-permeable skins form as a result of rapid quenching of the polymer solution as it initially contacts and solidifies on the mold surfaces. The skin thicknesses are a function of the solution cooling rate, and hence, the initial temperature and thermal mass of the mold surface adjacent to the polymer solution. The outer diameter skins of the single channel and 7-channel conduits are both approximately 25 mm thick, since the same mold and similar molding temperatures were used for both. However, the lumen skin of the single channel conduit is approximately 25 mm, significantly thicker than the 3 mm skin of the 7-channel conduit; the mass of the 1.35 mm diameter wire used for the single channel
Fig. 7. Scanning electron micrographs of longitudinal cross sections of single-channel conduit depicting (A) full cross section and lumen surface and (B) (C) increasing magnifications of bulk macrostructures and semi-permeable surface skins. Bars are 500, 200, and 100 mm in (A), (B), and (C), respectively.
conduit generates a significantly greater cooling rate relative to the 0.51 mm diameter wire of the 7-channel conduit.
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Fig. 8. Scanning electron micrographs of (A) (B) the lumen surface and (C) (D) the outer diameter surface of a single channel conduit. Bars are 100 mm for (A) and (C) and 5 mm for (B) and (D).
The inner lumenal and outer diameter surfaces of the single channel conduit are relatively smooth (Fig. 8). Some parallel shallow grooves existed on the lumenal surface, which were created as the wire was demolded and scraped the frozen polymer surface. However, the surface is generally smooth with submicron porosity. Lumenal surface roughness is known to affect peripheral nerve regeneration with smoother surfaces demonstrating enhanced regeneration [20]. The outer surface is rougher than the lumenal surfaces and appears relatively nonporous. The parallel aligned structures on the outer surface are mirror images of the machining marks on the mold surface. Tailoring the solvent system could easily modify the microstructures. Replacing the acetic acid with more typical TIPS solvents, such as 1,4-dioxane or benzene, should produce similar macrostructures [67,69] because a solid-liquid phase separation would occur with a crystallizable solvent. However, a microcellular architecture could be formed by the addition of a nonsolvent; i.e. the combined dioxane/water system. Instead of inducing a solid–liquid phase separation upon
cooling, a liquid–liquid phase separation would be induced which gives rise to an isotropic continuous cellular porosity [69], originating from the interconnected separated phases. Thus the solvent system choice could provide flexibility in the design of the porous matrix. Using this process, the number and diameters of the channels can easily and inexpensively be varied. These longitudinally aligned channels provide a framework which guides the regenerating axons but also provides a support for the adherence of physiologically relevant Schwann cells. Approximately 2.5 times more Schwann cells can be adhered in a monolayer to the lumen surfaces of a 7-channel conduit in comparison to a single channel conduit; the surface area of the 7 channel conduit would allow 3500 Schwann cells/mm to adhere in monolayer fashion, approximately the Schwann cell number in a healthy native peripheral nerve. In a 100channel conduit, the lumenal surface is 12.5 times that of a single channel conduit; approximately 14,000 Schwann cells/mm could be adhered, five times that of a healthy nerve.
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In vivo studies were conducted using the nerve conduits described above in which adherent Schwann cells were introduced or a model neurotrophin was encapsulated in the porous conduit walls and was slowly released over a period of weeks to months. In the first study, 7-channel conduits were seeded with Schwann cells and the adherent Schwann cell/conduit composites were implanted into 7 mm gap defects in rat sciatic nerves. Early regeneration results compared favorably with autografts [63]. In the second study, inosine, a purine analog thought to possess neurotrophic properties, was encapsulated in the polymer solution and single channel conduits were produced using the process described above. Inosine-loaded conduits were implanted into 7 mm gap defects in rat sciatic nerves. Again, early regeneration results suggest that enhanced regeneration occurred through the inosine-loaded conduits in comparison with inosine-free conduits [65].
4. Conclusions A method to fabricate porous, biodegradable conduits using a combined injection molding, thermally induced phase transition (TIPS) technique was developed which produced conduits with dimensionally toleranced, longitudinally aligned channels; the geometry of the channels was designed to approximate the architecture of native peripheral nerves and to promote the monolayer adherence of Schwann cells. As an example, the fabrication method was described as successfully applied using the biodegradable polymer dl-PLGA (85:15) and the solvent acetic acid to fabricate single-channel and 7-channel conduits. However, the process could easily be tailored to other polymer and solvent systems and has been used to produce conduits ranging from a single channel of 1.35 mm diameter to 100 channels of 0.08 mm diameter. As the polymer solution cooled following injection into a cold mold, a solid–liquid phase separation was induced which allowed for the production of macroporous foams of a high anisotropy. Semi-permeable skins formed on the outer diameter and lumen diameters of the conduit as a result of rapid quenching of the polymer solution upon contact with the cold mold surfaces. Macropores were organized into bundles of channels, up to 20 mm wide, in the dl-PLGA matrix and represented remnants of the acetic acid that crystallized during the polymer solution solidification.
Acknowledgements We gratefully acknowledge Dr. Jinming Gao for conducting the NMR analysis.
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