- Email: [email protected]
Design and evaluation of novel polyanhydride blends as nerve guidance conduits
Acta Biomaterialia 6 (2010) 1917–1924
Contents lists available at ScienceDirect
Acta Biomaterialia journal homepage: www.elsevier.com/locate/actabiomat
Design and evaluation of novel polyanhydride blends as nerve guidance conduits Jeremy Griffin a, Ashley Carbone b, Roberto Delgado-Rivera b, Sally Meiners c, Kathryn E. Uhrich a,b,* a
Department of Biomedical Engineering, Rutgers University, 599 Taylor Road, Piscataway, NJ 08854, USA Department of Chemistry and Chemical Biology, Rutgers University, 610 Taylor Road, Piscataway, NJ 08854, USA c Department of Pharmacology, Robert Wood Johnson Medical School, 675 Hoes Lane, Piscataway, NJ 08854, USA b
a r t i c l e
i n f o
Article history: Received 2 July 2009 Received in revised form 16 November 2009 Accepted 17 November 2009 Available online 23 November 2009 Keywords: Nerve regeneration Nerve conduit Polyanhydride Tissue engineering Xylene
a b s t r a c t Implantable biodegradable nerve guidance conduits (NGCs) have the potential to align and support regenerating cells, as well as prevent scar formation. In this study in vitro bioassays and in vivo material evaluations were performed using a nerve guidance conduit material made from a novel polyanhydride blend. In vitro cytotoxicity studies with both fibroblasts and primary chick neurons demonstrated that the proposed polyanhydride blend was non-cytotoxic. Subcutaneous implantation for 7 days in rats resulted in an initial fibrin matrix, minimal macrophage presence and angiogenesis in the surrounding tissues. Nerve guidance conduits fabricated from the proposed polyanhydride blend material may serve as favorable biocompatible tissue engineering devices. Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Peripheral nerves consist of parallel bundles of neuronal axons bound together by support tissue into long cables that extend from the brain and spinal cord for the purpose of innervating the numerous regions of the body. Damage to a peripheral nerve cable results in loss of sensation and motor control in the region that it innervates and is typically debilitating. Following laceration, natural nerve regeneration is an inefficient process involving obstructive scar formation, random axon regrowth and possible neuroma growth [1]. Recovery from nerve injury requires repair of the damaged nerve tissue, a challenge that has only been partially met by inserting autografts of functional nerve to bridge the injury. Success rates remain less than 80%, with the prognosis dropping off as the severity of the injury increases [2]. Autogenous nerve grafts may serve as ideal nerve conduits by providing a stimulating scaffold for regeneration; however, the number and length of potential donor nerves is limited [3]. Additionally, secondary surgical procedures to harvest the autogenous tissue results in permanent sensory deficiencies at the donor site [4–6]. Synthetic nerve guidance conduits (NGCs) are promising alternatives to autograft nerve for tissue regeneration because they have the potential to align and support regenerating cells, as well as prevent scar formation. An essential process that must occur
* Corresponding author. Address: Department of Chemistry and Chemical Biology, Rutgers University, 610 Taylor Road, Piscataway, NJ 08854, USA. Tel.: +1 732 445 0361; fax: +1 732 445 7036. E-mail address: [email protected] (K.E. Uhrich).
for functional recovery from nerve injury is the regeneration of damaged axons across the injury site. Schwann cells, the support cells of the peripheral nervous system, aid in this process by clearing debris and producing factors that promote regeneration [3]. The most fundamental function that a nerve graft serves is to provide structural support at the lesion site [1], acting as a physical bridge across the damaged region to direct migrating Schwann cells and regenerating axons to the distal nerve stump. For this purpose a synthetic nerve graft material must be able to interact favorably with the cellular components of nerve tissue and provide appropriate physical strength. Ideally the NGC would biodegrade following regeneration, leaving the biological structure intact and avoiding a chronic foreign body reaction or possible second surgery [7,8]. Most current nerve guidance products on the market are hollow tube structures that serve as a protective environment for nerve regeneration. Examples include the Synovis GEM NeurotubeÒ bioabsorbable poly(glycolic acid) (PGA) mesh and the Integra NeuraWrap™ collagen nerve protector. In this study we fabricated and evaluated an alternative NGC with appropriate handling and chemical properties: a,a0 -bis(o-carboxyphenoxy)-p-xylene-based polyanhydrides [poly(o-CPX)] [9] were admixed with poly(lactide anhydride) (PLAA). NGCs comprised of polyanhydrides are advantageous when compared with the biologically derived NCGs commercially available because the adverse potential for batch to batch variability is eliminated by a well-defined chemical composition and degradation products. The molecules chemically incorporated into the polymer have previously been investigated as components of pharmacological and medicinal compounds [9–
1742-7061/$ - see front matter Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.11.023
1918
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
13]. The enhanced solubility and thermal properties [i.e., glass transition temperature (Tg) and melting temperature (Tm)] of the surface eroding aromatic polyanhydrides allow easy fabrication into NGCs. Furthermore, polymer synthetic approaches potentially allow drugs to be incorporated into the NGC that can be released in a controlled and predictable manner [10–12,14]. 2. Materials and methods 2.1. Polymer preparation and visualization Poly(o-carboxyphenoxy)-p-xylene) [poly(o-CPX)] (Fig. 1A) was prepared as previously described [9]. Briefly, the synthetic approach involved coupling an o-substituted phenol to dibromoxylene to yield a benzylic ether. The methyl esters were hydrolyzed to give a free diacid, which was then acetylated to yield the monomer. The polyanhydrides (Fig. 1A) were then prepared by melt condensation of the monomer using previously established polymerization conditions [13]. Poly(lactide anhydride) (PLAA) (Fig. 1B) was obtained from Bioabsorbable Therapeutics (Menlo Park, CA). Polymers were blended and extruded using a HAAKE MiniLab II Micro Compounder (Thermo Scientific, Waltham, MA) into hollow conduits with an outside diameter of 2.5 mm and inside diameter of 2.3 mm and were cut to a length of 12 mm. Following melt extrusion, the surface morphology was observed on an AMRAY 1830 I scanning electron microscope. Samples were mounted on aluminum studs and sputter coated with gold–palladium using a SCD 004 Sputter Coater (Bal-Tec, Liechtenstein). 2.2. In vitro cytotoxicity Cell compatibility of the polymer blend of poly(o-CPX)–PLAA was performed by culturing NCTC clone 929 (strain L) mouse areolar fibroblast cells (ATCC, Manassas, VA) in medium containing the dissolved polymer. These L929 fibroblast cells are a standard cell type for cytocompatibility testing as recommended by the ASTM [15]. The polymer was dissolved in 10 mg ml1 dimethyl sulfoxide (DMSO) (Sigma, St Louis, MO) as a stock solution and serially diluted with cell culture medium to two concentrations (0.01 and 0.10 mg ml1), based on standard cytotoxicity protocols [16–19]. The cell culture medium consisted of Dulbecco’s modified Eagle’s medium (DMEM) (Sigma-Aldrich, St Louis, MO), 10 vol.% fetal bovine serum (Atlanta Biologicals, Lawrenceville, GA), 1% L-glutamate (Sigma) and 1% penicillin/streptomycin (Sigma). The polymer-containing medium was distributed in a 96-well plate (Fisher, Fair Lawn, NJ) and seeded at an initial concentration of 2000 cells per well (0.33 cm2). The medium with dissolved polymer was compared with three controls: (i) poly(lactic–co-glycolic acid) (PLGA)
(50:50) (BPI, Birmingham, AL) dissolved in medium (0.10 and 0.01 mg ml1); (ii) medium containing DMSO (0.10 and 0.01 mg ml1); (iii) medium alone. Cellular morphology was observed and documented at 10 original magnification (IX81, Olympus, Center Valley, PA) at 48, 72 and 96 h post-seeding. Cell viability was determined using a CellTiter 96Ò Aqueous One Solution Cell Proliferation Assay (Promega, Madison, WI). The MTS tetrazolium compound [3-(4,5dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-7 sulfophenyl)-2H-tetrazolium, inner salt] is bio-reduced by cells into a colored formazan product that is soluble in tissue culture medium. Following an appropriate incubation time 20 ll of the MTS reagent was added to 100 ll of culture medium and further incubated for 4 h. The absorbance was then recorded with a microplate reader (Model 680, Bio-Rad, Hercules, CA) at 490 nm. Cell numbers were calculated based upon a standard curve created 24 h after original seeding. Statistical analysis was performed with SPSS (originally Statistical Package for the Social Sciences) software (v.15.0 for Windows, SPSS, Chicago, IL). ANOVA followed by pair-wise comparison with Scheffe’s post hoc test allowed for pair-wise comparison of each of the polymers to the DMSO medium control and medium alone. 2.3. In vitro primary nerve cell viability The cytocompatibility of the polymer blend of poly(o-CPX)– PLAA in a mimetic NGC environment was compared with two controls: an uncoated glass coverslip and a coverslip coated with PLGA (50:50). The polymers were dissolved in methylene chloride (50 mg ml1) (Fisher) and spin-coated onto glass coverslips (18 mm diameter, 0.15 mm thickness) at 2000 rpm. for 30 s using a spin-coater (Headway Research, Garland, TX). Prior to spin-coating the coverslips were cleaned using Alconox (Alconox, New York, NY) and H2SO4:H2O2 (10:1 vol.%) solutions and stored in 70% ethanol. After spin-coating the methylene chloride solution was evaporated in a Pyrex dessicator under vacuum (Fisher, UK) for 24 h. Coverslips were placed in 12-well plates (Fisher) and sterilized under UV light at 254 nm for 900 s using a Spectrolinker XL-1500 UV cross-linker system (Spectronics Corp., Westbury, NY). To mimic in vivo conditions a collagen gel was prepared and self-assembled on top of the spin-coated polymer surface. The collagen gel was prepared by mixing 20 ll of Medium 199 (Sigma), 1 ll penicillin/streptomycin, 10 ll L-glutamine (Sigma) and 677 ll Vitrogen collagen solution in 0.02 M acetic acid (3.0 g collagen ml1) (Cohesion Corp., Temecula, CA). An aliquot of 500 ll of the collagen network self-assembled into a gel in the well of the 12-well plate when incubated at 37 °C under 5% CO2. The experimental design of the mimetic NGC environment is shown in Fig. 2. Dorsal root ganglia (DRG) were prepared from embryonic day 8 (E8) White Leghorn chick embryos (Charles River, Wilmington,
Fig. 1. Chemical structures of (A) poly(o-carboxyphenoxy)-p-xylene) [poly(o-CPX)] and (B) poly(lactide-anhydride) [PLAA].
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
1919
Fig. 2. Theoretical cross-section of the NGC (A) correlated with a representative in vitro experimental model in a cell culture well plate (B) with a collagen depth of 1.0 mm on top of a thin polymer surface on top of a glass coverslip.
MA) and seeded into the three-dimensional collagen gel [20,21]. DRGs were uniformly placed in the center of the well plate in 500 ll of viscous liquid collagen matrix (prior to gelling). DRGs were cultured in medium composed of DMEM containing 10 vol.% fetal bovine serum (Atlanta Biologicals), 2 mM L-glutamine (Sigma), 50 U ml1 penicillin/streptomycin (Sigma) and 12 ng ml1 nerve growth factor (Gibco, Carlsbad, CA). Three bright field images were taken of the DRGs to demonstrate their incremental neurite outgrowth (day 1, Fig. 6A; day 4, Fig. 6B; day 7, Fig. 6C). DRG growth was quantified as percent increase in the neurite halo compared with the original body size (D1, D4 or D7/Dd), where D1, D4 and D7 represent the average equivalent diameter of the neurite halo on days 1, 4 and 7, respectively, and Dd is the diameter of the original dense DRG mass. Neurite extension was measured on days 1, 4 and 7 and compared with the original circumscribed DRG ‘body’ region. A larger halo of neurites typically represented a more viable environment for neuron growth. Further morphological observation was performed by immunostaining the DRG neurons using antibodies against neurofilament 68 and neurofilament 200, abundant neuron-specific cytoskeletal proteins. Neurons were fixed in 4% paraformaldehyde solution (Sigma) for 20 min at 25 °C and washed (three times) with an immunobuffer solution [1% bovine serum albumin (Sigma) and 0.5% Triton X-100 (Fisher) in phosphate-buffered saline]. Neurons were blocked in 10% goat serum (Sigma, St Louis, MO) in immunobuffer for 1 h. Samples were incubated in primary antibody solution to neurofilament 68 (1:1000) and neurofilament 200 (1:200) (Sigma) for 1 h at 25 °C, followed by secondary antibody incubation with anti-mouse Alexafluor 546 (1:400) (Molecular Probes, Carlsbad, CA) for 45 min at 25 °C.
2.4. In vivo immune response The polymer blend of poly(o-CPX)–PLAA was placed intramuscularly (gluteus maximus) in adult female Sprague–Dawley rats (220–300 g) (n = 3). The immune response was compared with the immune response to PLGA (50:50). The animals were anesthetized by intraperitoneal injection of 75 mg kg1 ketamine/ 10 mg kg1 xylazine. Gluteal muscle splitting incisions were made in the animals and melt extruded solid rods (1.0 mm diameter, 4.0 mm length, 8.0 mg) were placed intramuscularly (gluteus maximus) in the animals, followed by musculature suturing and skin closure with surgical clips. Immediately following surgery the animals were subcutaneously administered the analgesic buprenorphine (0.05 mg kg1), to alleviate potential post-operative pain, the antibiotic enrofloxacin (5 mg kg1) (Baytril), to manage and prevent infections, and Ringer’s lactate solution (5 ml) to manage dehydration. Following the surgical procedure, animals were allowed to recover in clean
cages placed on top of 37 °C heating pads. All animals were observed for infection and/or lethargy following surgery, then again treated subcutaneously with 0.05 mg kg1 buprenorphine and 5 mg kg1 enrofloxacin 12 and 24 h following surgery. All animal procedures were performed in strict accordance with institutional guidelines (Institutional Animal Care and Use Committee Approval no. I08-071-09). Seven days after implantation the animals were transcardially perfused under deep terminal anesthesia (100 mg kg1 sodium pentobarbital intraperitoneal) with 4% paraformaldehyde in 0.1 mol l1 sodium phosphate buffer. The polymer implants were excised with the surrounding tissue and immersed in 4% paraformaldehyde for 24 h, followed by serial immersion for 24 h in each of three sucrose solutions (10%, 20% and 30% sucrose in 0.05 M sodium phosphate buffer) for cryoprotection. The tissues and implants were then sagittally sectioned on a cryostat at 20 lm. Sections were mounted on plus (+) gold slides (Fisher Scientific) and histological analysis was performed on the extracted tissue to evaluate the in vivo response to the biomaterials. The tissue was stained with Hoechst 33258 nucleic acid stain (1:200 dilution for 15 min at room temperature) (Molecular Probes, Eugene, OR) to detect cell nuclei and double immunolabeled with: (1) a mouse monoclonal antibody ED1 (AbD Serotec, Raleigh, NC) (1:400 dilution overnight at room temperature) followed by a CY3-conjugated goat anti-mouse secondary antibody (Jackson ImmunoResearch, West Grove, PA) (1:500 dilution for 1 h at room temperature) to identify lysosomal antigens in macrophages and giant cells; (2) a rabbit polyclonal antibody against collagen IV (Research Diagnostics, Concord, MA) (1:400 dilution overnight at room temperature) followed by a CY2-conjugated goat anti-rabbit secondary antibody (1:500 dilution for 1 h at room temperature) to detect angiogenesis. A second set of tissue sections was double immunolabeled with: (1) a mouse monoclonal antibody ED1 (1:400 dilution overnight at room temperature) followed by a CY3-conjugated goat anti-mouse secondary (1:500 dilution for 1 h at room temperature) to identify lysomal antigens in macrophages and giant cells; (2) a rabbit polyclonal antibody against fibronectin (Chemicon, Temecula, CA) (1:400 dilution overnight at room temperature) followed by a CY2conjugated goat anti-rabbit secondary antibody to identify fibrotic scarring. All antibodies and Hoechst stain were diluted in phosphate-buffered saline containing 0.3% Triton X-100. Several sections of the tissue were examined for each animal using an epifluorescent Olympus IX81 inverted microscope (Olympus, Melville, NY).
3. Results and discussion Blends of poly(o-CPX) (15 wt.%) (Fig. 1A) and PLAA (85 wt.%) (Fig. 1B) were investigated for their potential as nerve guidance conduits.
1920
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
Both polymers were blended and extruded as hollow tubes with an outside diameter of 2.5 mm and an inside diameter of 2.3 mm. Fig. 3 shows an example of the outside surface and inside lumen morphology of the melt extruded conduits. Suture holes for in vivo studies were created by heating a 21 gauge needle (Becton Dickinson, Franklin Lakes, NJ) and pressing it through the polymer 1 mm from the end of the conduit. 3.1. In vitro cytotoxicity: fibroblasts Cytocompatibility was demonstrated by two approaches: (i) fibroblast growth in polymer-containing medium to mimic late stage polymer degradation; (ii) primary neuron growth on collagen gel-coated polymer substrates to mimic in vivo conditions. Cytocompatibility of the polymer blend of poly(o-CPX)–PLAA was first evaluated by culturing L929 mouse fibroblasts cells in polymercontaining medium at concentrations of 0.10 and 0.01 mg polymer ml medium1. Cells were cultured over a 3 day period, during which cell proliferation and morphology were observed. The intent of this study was to determine cell viability in medium containing the polymer blend of poly(o-CPX)–PLAA. The best mechanical integrity and handling properties were observed with the blend – the homopolymers did not meet this criterion for an NGC. PLGA was chosen as the control because it is a common co-polymer that is used in FDA-approved devices and in a variety of biomedical implants, including nerve guidance conduits. For polymers at both concentrations fibroblasts demonstrated increased cell proliferation throughout the 96 h studied (Fig. 4). As indicated by pair-wise comparison with Scheffe’s post hoc test, the medium containing polymer degradation products at 0.01 mg polymer ml medium1 did not statistically differ from the of DMSO medium and medium alone controls after 96 h cell culture (P < 0.05). For the 0.10 mg ml1 concentration, significant differences were found between the polymer blend of poly(o-CPX)–PLAA and the medium controls at 72 and 96 h (indicated by an asterisk in Fig. 4). Although lower cell proliferation (cell count) was evident at the higher concentration (0.10 mg ml1) of polymer in comparison with the DMSO medium and medium controls, the morphological images (Fig. 5) illustrate cells with stellate extended filipodia at all concentrations, indicating healthy fibroblast cell growth. 3.2. In vitro primary nerve cell viability To model in vivo conditions a collagen gel was prepared on the polymer surface. The experimental design of the NGC mimetic
environment is shown in Fig. 2. For simplicity a NGC of 2 mm inner diameter and 2.2 mm outer diameter was chosen as the system to model in the tissue culture plate. E8 chick DRG were seeded into the three-dimensional collagen gel and cultured in medium containing nerve growth factor for 7 days. Under bright field microscopy, once seeded into the collagen gel the DRG initially appeared to be a dark mass of diameter Dd with no protruding neurites (Fig. 6A). Neurite extension was measured on days 1, 4, and 7 and compared with the original circumscribed DRG ‘body’ region. A DRG that failed to demonstrate neurite extension after 7 days of culture would be indicative of an environment which is unfavorable for neuron growth, i.e. non-viable. An ever increasing halo of neurons extending from the dense DRG body (Dd) over 7 days culture is indicative of a viable environment/surface for neuron growth. Fig. 6A–C demonstrates the technique utilized to evaluate the effect of the degradation products of the polymer blend of poly(o-CPX)–PLAA on neurite outgrowth. Fig. 6E shows the overall increase in neurite outgrowth compared with the size of the original DRG body for the polymer blend of poly(o-CPX)–PLAA compared with PLGA (50:50) and an uncoated glass control. After 7 days culture the DRGs grew as follows: 827 ± 174% on the glass surface, 776 ± 221% on the PLGA and 641 ± 153% on the polymer blend of poly(o-CPX)–PLAA. As indicated by pair-wise comparison with Scheffe’s post hoc test, significant variance among the initial polymer conditions did not result in statistical differences in the extent of neuritic field extension after 7 days (P < 0.05). No additives were included within the chosen polymer blend of poly(o-CPX)–PLAA, PLGA or glass control to further encourage neurite outgrowth. While this system looks promising, we are pursuing further studies involving patterned surfaces, collagen NGC lumen fillers and additional growth factors (nerve growth factor, fibroblast growth factor, etc.) in an attempt to develop an optimal environment for neurite growth. 3.3. In vivo immune response Following 7 days vivo intramuscular implantation in Sprague– Dawley rats, the polymer blend of poly(o-CPX)–PLAA favorably compared with the FDA-approved PLGA (50:50). No profound immunohistological differences were observed between the tissue surrounding the two polymers, as shown in Fig. 7. Excised tissue was stained for an inflammatory response (ED1, macrophages), fibrous scarring (fibronectin), angiogenesis (collagen IV) and cell nuclei (Hoechst 33258 nucleic acid stain). The
Fig. 3. SEM images of the melt extruded polymer blend of poly(o-CPX)–PLAA conduit. Suture holes were melt pressed with a 21 gauge needle into the material for in vivo studies. Scale bars 1 mm.
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
1921
Fig. 4. Cell viability/proliferation in culture media with polymers at concentration of (A) 0.01 mg polymer ml medium1 and (B) 0.10 mg polymer ml medium1. *Statistically different at P < 0.05.
Fig. 5. L929 mouse fibroblasts after 96 h cell culture (10). (A) Polymer blend of poly(o-CPX)–PLAA at 0.01 mg ml1; (B) PLGA (50:50) at 0.01 mg ml1; (C) polymer blend of poly(o-CPX)–PLAA at 0.1 mg ml1; (D) PLGA (50:50) at 0.10 mg ml1. Scale bars 200 lm.
1922
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
Fig. 6. Bright field images of incremental neurite outgrowth on days 1 (A), 4 (B), and 7 (C) on a polymer blend of poly(o-CPX)–PLAA. (D) Immunostained DRG (as C) against neurofilament 68 and 200. DRG growth was quantified as percentage increase in the neurite halo compared with the original body size (D1, D4 or D7/Dd). Scale bar: (A) 1000 lm; (B–D) 2000 lm. (E) Comparison of the percentage increase in neurite outgrowth compared with the original DRG body size after 7 days culture on a polymer blend of poly(o-CPX)–PLAA, PLGA and glass (control) collagen substrates.
results of the immunocytochemical analysis for ED1, collagen IV and fibronectin is presented in Table 1. Immunohistological staining of the tissue was scored in a range from deficient to very abundant for each molecule. After 7 days the polymer blend of poly(o-CPX)–PLAA and PLGA implants were both surrounded by minimal fibrin matrix (Fibronectin, +). The surrounding tissue also indicated a healthy accepting response, containing as it did normal amounts of fibrin. Compared with the FDA-approved PLGA implant, minimal macrophage migration (ED1, ±) was visible in the polymer blend of
poly(o-CPX)–PLAA implant over the 7 day implantation period. A pronounced inflammatory tissue reaction should result in a much greater concentration of macrophages at the site of implantation and would be clearly visible on the histological sections [22]. The visible tissue reactions were more consistent with local phagocytosis, rather than a large scale inflammatory response. With the ultimate goal of serving as a hollow conduit to support regenerating nerves, this biomaterial must support angiogenesis. The abundance of collagen IV in this immunostaining study for the polymer blend of poly(o-CPX)–PLAA (collagen IV, +++) indi-
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
1923
Fig. 7. Immunohistological staining for an immune response. (A, C) ED1, red; collagen IV, green; cell nuclei, blue. (B, D) Fibronectin, green; cell nuclei, blue. Scale bar 200 lm.
Table 1 Immunocytochemical analysis of subcutaneously implanted polymer rods. ED1, collagen IV and fibronectin were scored in the tissues surrounding the material: , deficient; ±, sparse; +, moderate; ++, clearly expressed; +++, abundant; ++++, very abundant. Polymer rod composition
ED1
Collagen IV
Fibronectin
Poly(o-CPX) [15 wt.%]:PLAA [85 wt.%] PLGA (50:50)
± ±
+++ +++
+ ++
cates that the material does not adversely affect the formation of capillaries surrounding the implant. This further indicates that hollow conduits of the polymer blend of poly(o-CPX)–PLAA will be conducive to providing the well-vascularized environment that is necessary to support nerve regeneration [23].
such polyanhydrides have the capability to have therapeutics chemically incorporated into the polymer backbone with additional labile ester bonds, which could allow for controlled release of bioactives, if desired [18,26]. The polyanhydrides were melt extruded into conduits of appropriate size to evaluate their in vivo regenerative properties and characteristics using the sciatic nerve model in Sprague– Dawley rats. The polyanhydride blends evoked a mild foreign body reaction, with biocompatibility comparable with PLGA implants. Future studies will involve additional functionality arising from an impregnated collagen matrix. Based upon these positive results, conduits fabricated from the polymer blend of poly(oCPX)–PLAA will be more extensively evaluated as tissue engineering devices. Acknowledgements
4. Conclusion NGCs comprised of polyanhydrides such as PLAA and poly(oCPX) are advantageous compared with the biologically derived, commercially available NGCs. NGCs made from collagen or other complex biological materials have the potential for batch-to-batch variability [24], whereas synthetic NGCs (such as the polyanhydrides) have a well-defined chemical composition that directly controls polymer degradation, as well as device degradation [9]. In addition, these polyanhydrides have exceptional thermal properties, with glass transition temperatures (Tg) well above the physiological temperature, which is desirable to avoid loss of mechanical integrity when plasticizers or admixtures of other polymers are utilized [9]. Poly(o-CPX) and related polyanhydrides were chosen since their components and resulting degradation products are known to be biocompatible [10–13,25]. Additionally,
This research was supported by US Army contract no. W81XWH-04-2-0031. The authors thank Olex Hnojewyj (BTI, Menlo Park, CA) for providing melt extruded conduits, Natasha Piracha (Department of Cell Biology and Neuroscience, Rutgers University) for assisting with the fibroblast cytotoxicity study and Susan Harris (Department of Pharmacology, Robert Wood Johnson Medical School) for her help with the immunohistological study. References [1] IJkema-Paassen J, Jansen K, Gramsbergena A, Meek M. Transection of peripheral nerves, bridging strategies and effect evaluation. Biomaterials 2003;25:1583–92. [2] Committee on Capturing the Full Power of Biomaterials for Military Medical Needs, National Research Council. Capturing the full power of biomaterials for military medical needs. Washington, DC: National Academic Press; 2004.
1924
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
[3] Ichihara S, Inada Y, Nakamura T. Artificial nerve tubes and their application for repair of peripheral nerve injury: an update of current concepts. Injury: Int J Care Injured 2008;39S4:S29–39. [4] Maquet V, Martin D, Malgrange B, Franzen R, Schoenen J, Moonen G, et al. Peripheral nerve regeneration using bioresorbable macroporous polylactide scaffolds. J Biomed Mater Res 2000;52:639–51. [5] Dellon A, Mackinnon S. An alternative to the classical nerve graft for the management of the short nerve gap. Plast Reconstr Surg 1988;82:849–56. [6] Evans G. Challenges to nerve regeneration. Semin Surg Oncol. 2000;19:312–8. [7] den Dunnen W, Stokroos I, Blaauw E, Holwerda A, Pennings A, Robinson P, et al. Light-microscopic and electron-microscopic evaluation of short-term nerve regeneration using a biodegradable poly (dl-lactide–caprolacton) nerve guide. J Biomed Mater Res A 1998;31:105–15. [8] Nicoli Aldini N, Perego G, Cella G, Maltarello M, Fini M, Rocca M, et al. Effectiveness of a bioabsorbable conduit in the repair of peripheral nerves. Biomaterials 1996;17:959–62. [9] Anastasiou T, Uhrich K. Novel polyanhydrides with enhanced thermal and solubility. Macromolecules 2000;33:6217–21. [10] Leong K, Brott B, Langer R. Bioerodable polyanhydride as drug-carrier matrices I: characterization, degradation, and release characteristics. J Biomed Mater Res 1985;19:941–55. [11] Domb A, Langer R. Solid state and solution stability of poly(anhydrides) and poly(esters). Macromolecules 1989;22:2117–22. [12] Chasin M, Lewis D, Langer R. Polyanhydrides for controlled release. Biopharm Manuf 1988;1:33–46. [13] Campo C, Anastasiou T, Uhrich K. Polyanhydrides: the effects of ring substitution changes on polymer properties. Polym Bull 1999;42:61–8. [14] Uhrich K, Cannizzaro S, Langer R, Shakesheff K. Polymeric systems for controlled drug release. Chem Rev 1999;99:3181–98.
[15] Duncan E, Lemons J, Mayesh J, Saha P, Saha S, Scranton M. New Products and standards in biomaterials science. New York: Elsevier Academic Press; 2004. [16] Schmeltzer R, Schmalenberg KE, Uhrich KE. Biomacromolecules 2005;6:359–67. [17] Carbone A, Song M, Uhrich K. Iodinated sacilcylate-based poly(anhydrideesters) as radiopaque biomaterials. Biomacromolecules 2008;9:1604–12. [18] Fischer D, Li Y, Ahlemeyer B, Krieglstein J, Kissel T. In vitro cytotoxicity testing of polycations: influence of polymer structure on cell viability and hemolysis. Biomaterials 2003;24:1121–31. [19] International Standards Organization. Biological evaluation for medical devices-part 5: tests for cytotoxicity in vitro methods. ISO-10993-5 (EN 30993-5);1992. [20] Luduena RF, Woodward DO. Isolation and partial characterization of alpha and beta-tubulin from outer doublets of sea-urchin sperm and microtubules of chick-embryo brain. Proc Natl Acad Sci USA 1973;70:3594–8. [21] Schmalenberg KE, Buettner HM, Uhrich KE. Microcontact printing of proteins on oxygen plasma-activated poly(methyl methacrylate). Biomaterials 2004;25:1851–7. [22] Jansen K, van der Werff JFA, van Wachem PB, Nicolai J-PA, de Leij LFMH, van Luyn MJA. A hyaluronan-based nerve guide: in vitro cytotoxicity, subcutaneous tissue reactions, and degradation in the rat. Biomaterials 2004;25:483–9. [23] Hobson MI, Brown R, Green CJ, Terenghi G. Inter-relationships between angiogenesis and nerve regeneration: a histochemical study. Br J Plast Surg 1997;50:125–31. [24] Lutolf M. Integration column: artificial ECM: expanding the cell biology toolbox in 3D. The Royal Soc Chem: Integr Biol 2009;1:235–41. [25] Merck Research. The Merck Index. Whitehouse Station, NJ: Merck Research Laboratories; 2001. [26] Prudencio A, Schmeltzer R, Uhrich K. Using the ‘‘linker” to control polymer properties. Macromolecules 2005;38:6895–901.
Contents lists available at ScienceDirect
Acta Biomaterialia journal homepage: www.elsevier.com/locate/actabiomat
Design and evaluation of novel polyanhydride blends as nerve guidance conduits Jeremy Griffin a, Ashley Carbone b, Roberto Delgado-Rivera b, Sally Meiners c, Kathryn E. Uhrich a,b,* a
Department of Biomedical Engineering, Rutgers University, 599 Taylor Road, Piscataway, NJ 08854, USA Department of Chemistry and Chemical Biology, Rutgers University, 610 Taylor Road, Piscataway, NJ 08854, USA c Department of Pharmacology, Robert Wood Johnson Medical School, 675 Hoes Lane, Piscataway, NJ 08854, USA b
a r t i c l e
i n f o
Article history: Received 2 July 2009 Received in revised form 16 November 2009 Accepted 17 November 2009 Available online 23 November 2009 Keywords: Nerve regeneration Nerve conduit Polyanhydride Tissue engineering Xylene
a b s t r a c t Implantable biodegradable nerve guidance conduits (NGCs) have the potential to align and support regenerating cells, as well as prevent scar formation. In this study in vitro bioassays and in vivo material evaluations were performed using a nerve guidance conduit material made from a novel polyanhydride blend. In vitro cytotoxicity studies with both fibroblasts and primary chick neurons demonstrated that the proposed polyanhydride blend was non-cytotoxic. Subcutaneous implantation for 7 days in rats resulted in an initial fibrin matrix, minimal macrophage presence and angiogenesis in the surrounding tissues. Nerve guidance conduits fabricated from the proposed polyanhydride blend material may serve as favorable biocompatible tissue engineering devices. Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Peripheral nerves consist of parallel bundles of neuronal axons bound together by support tissue into long cables that extend from the brain and spinal cord for the purpose of innervating the numerous regions of the body. Damage to a peripheral nerve cable results in loss of sensation and motor control in the region that it innervates and is typically debilitating. Following laceration, natural nerve regeneration is an inefficient process involving obstructive scar formation, random axon regrowth and possible neuroma growth [1]. Recovery from nerve injury requires repair of the damaged nerve tissue, a challenge that has only been partially met by inserting autografts of functional nerve to bridge the injury. Success rates remain less than 80%, with the prognosis dropping off as the severity of the injury increases [2]. Autogenous nerve grafts may serve as ideal nerve conduits by providing a stimulating scaffold for regeneration; however, the number and length of potential donor nerves is limited [3]. Additionally, secondary surgical procedures to harvest the autogenous tissue results in permanent sensory deficiencies at the donor site [4–6]. Synthetic nerve guidance conduits (NGCs) are promising alternatives to autograft nerve for tissue regeneration because they have the potential to align and support regenerating cells, as well as prevent scar formation. An essential process that must occur
* Corresponding author. Address: Department of Chemistry and Chemical Biology, Rutgers University, 610 Taylor Road, Piscataway, NJ 08854, USA. Tel.: +1 732 445 0361; fax: +1 732 445 7036. E-mail address: [email protected] (K.E. Uhrich).
for functional recovery from nerve injury is the regeneration of damaged axons across the injury site. Schwann cells, the support cells of the peripheral nervous system, aid in this process by clearing debris and producing factors that promote regeneration [3]. The most fundamental function that a nerve graft serves is to provide structural support at the lesion site [1], acting as a physical bridge across the damaged region to direct migrating Schwann cells and regenerating axons to the distal nerve stump. For this purpose a synthetic nerve graft material must be able to interact favorably with the cellular components of nerve tissue and provide appropriate physical strength. Ideally the NGC would biodegrade following regeneration, leaving the biological structure intact and avoiding a chronic foreign body reaction or possible second surgery [7,8]. Most current nerve guidance products on the market are hollow tube structures that serve as a protective environment for nerve regeneration. Examples include the Synovis GEM NeurotubeÒ bioabsorbable poly(glycolic acid) (PGA) mesh and the Integra NeuraWrap™ collagen nerve protector. In this study we fabricated and evaluated an alternative NGC with appropriate handling and chemical properties: a,a0 -bis(o-carboxyphenoxy)-p-xylene-based polyanhydrides [poly(o-CPX)] [9] were admixed with poly(lactide anhydride) (PLAA). NGCs comprised of polyanhydrides are advantageous when compared with the biologically derived NCGs commercially available because the adverse potential for batch to batch variability is eliminated by a well-defined chemical composition and degradation products. The molecules chemically incorporated into the polymer have previously been investigated as components of pharmacological and medicinal compounds [9–
1742-7061/$ - see front matter Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.11.023
1918
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
13]. The enhanced solubility and thermal properties [i.e., glass transition temperature (Tg) and melting temperature (Tm)] of the surface eroding aromatic polyanhydrides allow easy fabrication into NGCs. Furthermore, polymer synthetic approaches potentially allow drugs to be incorporated into the NGC that can be released in a controlled and predictable manner [10–12,14]. 2. Materials and methods 2.1. Polymer preparation and visualization Poly(o-carboxyphenoxy)-p-xylene) [poly(o-CPX)] (Fig. 1A) was prepared as previously described [9]. Briefly, the synthetic approach involved coupling an o-substituted phenol to dibromoxylene to yield a benzylic ether. The methyl esters were hydrolyzed to give a free diacid, which was then acetylated to yield the monomer. The polyanhydrides (Fig. 1A) were then prepared by melt condensation of the monomer using previously established polymerization conditions [13]. Poly(lactide anhydride) (PLAA) (Fig. 1B) was obtained from Bioabsorbable Therapeutics (Menlo Park, CA). Polymers were blended and extruded using a HAAKE MiniLab II Micro Compounder (Thermo Scientific, Waltham, MA) into hollow conduits with an outside diameter of 2.5 mm and inside diameter of 2.3 mm and were cut to a length of 12 mm. Following melt extrusion, the surface morphology was observed on an AMRAY 1830 I scanning electron microscope. Samples were mounted on aluminum studs and sputter coated with gold–palladium using a SCD 004 Sputter Coater (Bal-Tec, Liechtenstein). 2.2. In vitro cytotoxicity Cell compatibility of the polymer blend of poly(o-CPX)–PLAA was performed by culturing NCTC clone 929 (strain L) mouse areolar fibroblast cells (ATCC, Manassas, VA) in medium containing the dissolved polymer. These L929 fibroblast cells are a standard cell type for cytocompatibility testing as recommended by the ASTM [15]. The polymer was dissolved in 10 mg ml1 dimethyl sulfoxide (DMSO) (Sigma, St Louis, MO) as a stock solution and serially diluted with cell culture medium to two concentrations (0.01 and 0.10 mg ml1), based on standard cytotoxicity protocols [16–19]. The cell culture medium consisted of Dulbecco’s modified Eagle’s medium (DMEM) (Sigma-Aldrich, St Louis, MO), 10 vol.% fetal bovine serum (Atlanta Biologicals, Lawrenceville, GA), 1% L-glutamate (Sigma) and 1% penicillin/streptomycin (Sigma). The polymer-containing medium was distributed in a 96-well plate (Fisher, Fair Lawn, NJ) and seeded at an initial concentration of 2000 cells per well (0.33 cm2). The medium with dissolved polymer was compared with three controls: (i) poly(lactic–co-glycolic acid) (PLGA)
(50:50) (BPI, Birmingham, AL) dissolved in medium (0.10 and 0.01 mg ml1); (ii) medium containing DMSO (0.10 and 0.01 mg ml1); (iii) medium alone. Cellular morphology was observed and documented at 10 original magnification (IX81, Olympus, Center Valley, PA) at 48, 72 and 96 h post-seeding. Cell viability was determined using a CellTiter 96Ò Aqueous One Solution Cell Proliferation Assay (Promega, Madison, WI). The MTS tetrazolium compound [3-(4,5dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-7 sulfophenyl)-2H-tetrazolium, inner salt] is bio-reduced by cells into a colored formazan product that is soluble in tissue culture medium. Following an appropriate incubation time 20 ll of the MTS reagent was added to 100 ll of culture medium and further incubated for 4 h. The absorbance was then recorded with a microplate reader (Model 680, Bio-Rad, Hercules, CA) at 490 nm. Cell numbers were calculated based upon a standard curve created 24 h after original seeding. Statistical analysis was performed with SPSS (originally Statistical Package for the Social Sciences) software (v.15.0 for Windows, SPSS, Chicago, IL). ANOVA followed by pair-wise comparison with Scheffe’s post hoc test allowed for pair-wise comparison of each of the polymers to the DMSO medium control and medium alone. 2.3. In vitro primary nerve cell viability The cytocompatibility of the polymer blend of poly(o-CPX)– PLAA in a mimetic NGC environment was compared with two controls: an uncoated glass coverslip and a coverslip coated with PLGA (50:50). The polymers were dissolved in methylene chloride (50 mg ml1) (Fisher) and spin-coated onto glass coverslips (18 mm diameter, 0.15 mm thickness) at 2000 rpm. for 30 s using a spin-coater (Headway Research, Garland, TX). Prior to spin-coating the coverslips were cleaned using Alconox (Alconox, New York, NY) and H2SO4:H2O2 (10:1 vol.%) solutions and stored in 70% ethanol. After spin-coating the methylene chloride solution was evaporated in a Pyrex dessicator under vacuum (Fisher, UK) for 24 h. Coverslips were placed in 12-well plates (Fisher) and sterilized under UV light at 254 nm for 900 s using a Spectrolinker XL-1500 UV cross-linker system (Spectronics Corp., Westbury, NY). To mimic in vivo conditions a collagen gel was prepared and self-assembled on top of the spin-coated polymer surface. The collagen gel was prepared by mixing 20 ll of Medium 199 (Sigma), 1 ll penicillin/streptomycin, 10 ll L-glutamine (Sigma) and 677 ll Vitrogen collagen solution in 0.02 M acetic acid (3.0 g collagen ml1) (Cohesion Corp., Temecula, CA). An aliquot of 500 ll of the collagen network self-assembled into a gel in the well of the 12-well plate when incubated at 37 °C under 5% CO2. The experimental design of the mimetic NGC environment is shown in Fig. 2. Dorsal root ganglia (DRG) were prepared from embryonic day 8 (E8) White Leghorn chick embryos (Charles River, Wilmington,
Fig. 1. Chemical structures of (A) poly(o-carboxyphenoxy)-p-xylene) [poly(o-CPX)] and (B) poly(lactide-anhydride) [PLAA].
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
1919
Fig. 2. Theoretical cross-section of the NGC (A) correlated with a representative in vitro experimental model in a cell culture well plate (B) with a collagen depth of 1.0 mm on top of a thin polymer surface on top of a glass coverslip.
MA) and seeded into the three-dimensional collagen gel [20,21]. DRGs were uniformly placed in the center of the well plate in 500 ll of viscous liquid collagen matrix (prior to gelling). DRGs were cultured in medium composed of DMEM containing 10 vol.% fetal bovine serum (Atlanta Biologicals), 2 mM L-glutamine (Sigma), 50 U ml1 penicillin/streptomycin (Sigma) and 12 ng ml1 nerve growth factor (Gibco, Carlsbad, CA). Three bright field images were taken of the DRGs to demonstrate their incremental neurite outgrowth (day 1, Fig. 6A; day 4, Fig. 6B; day 7, Fig. 6C). DRG growth was quantified as percent increase in the neurite halo compared with the original body size (D1, D4 or D7/Dd), where D1, D4 and D7 represent the average equivalent diameter of the neurite halo on days 1, 4 and 7, respectively, and Dd is the diameter of the original dense DRG mass. Neurite extension was measured on days 1, 4 and 7 and compared with the original circumscribed DRG ‘body’ region. A larger halo of neurites typically represented a more viable environment for neuron growth. Further morphological observation was performed by immunostaining the DRG neurons using antibodies against neurofilament 68 and neurofilament 200, abundant neuron-specific cytoskeletal proteins. Neurons were fixed in 4% paraformaldehyde solution (Sigma) for 20 min at 25 °C and washed (three times) with an immunobuffer solution [1% bovine serum albumin (Sigma) and 0.5% Triton X-100 (Fisher) in phosphate-buffered saline]. Neurons were blocked in 10% goat serum (Sigma, St Louis, MO) in immunobuffer for 1 h. Samples were incubated in primary antibody solution to neurofilament 68 (1:1000) and neurofilament 200 (1:200) (Sigma) for 1 h at 25 °C, followed by secondary antibody incubation with anti-mouse Alexafluor 546 (1:400) (Molecular Probes, Carlsbad, CA) for 45 min at 25 °C.
2.4. In vivo immune response The polymer blend of poly(o-CPX)–PLAA was placed intramuscularly (gluteus maximus) in adult female Sprague–Dawley rats (220–300 g) (n = 3). The immune response was compared with the immune response to PLGA (50:50). The animals were anesthetized by intraperitoneal injection of 75 mg kg1 ketamine/ 10 mg kg1 xylazine. Gluteal muscle splitting incisions were made in the animals and melt extruded solid rods (1.0 mm diameter, 4.0 mm length, 8.0 mg) were placed intramuscularly (gluteus maximus) in the animals, followed by musculature suturing and skin closure with surgical clips. Immediately following surgery the animals were subcutaneously administered the analgesic buprenorphine (0.05 mg kg1), to alleviate potential post-operative pain, the antibiotic enrofloxacin (5 mg kg1) (Baytril), to manage and prevent infections, and Ringer’s lactate solution (5 ml) to manage dehydration. Following the surgical procedure, animals were allowed to recover in clean
cages placed on top of 37 °C heating pads. All animals were observed for infection and/or lethargy following surgery, then again treated subcutaneously with 0.05 mg kg1 buprenorphine and 5 mg kg1 enrofloxacin 12 and 24 h following surgery. All animal procedures were performed in strict accordance with institutional guidelines (Institutional Animal Care and Use Committee Approval no. I08-071-09). Seven days after implantation the animals were transcardially perfused under deep terminal anesthesia (100 mg kg1 sodium pentobarbital intraperitoneal) with 4% paraformaldehyde in 0.1 mol l1 sodium phosphate buffer. The polymer implants were excised with the surrounding tissue and immersed in 4% paraformaldehyde for 24 h, followed by serial immersion for 24 h in each of three sucrose solutions (10%, 20% and 30% sucrose in 0.05 M sodium phosphate buffer) for cryoprotection. The tissues and implants were then sagittally sectioned on a cryostat at 20 lm. Sections were mounted on plus (+) gold slides (Fisher Scientific) and histological analysis was performed on the extracted tissue to evaluate the in vivo response to the biomaterials. The tissue was stained with Hoechst 33258 nucleic acid stain (1:200 dilution for 15 min at room temperature) (Molecular Probes, Eugene, OR) to detect cell nuclei and double immunolabeled with: (1) a mouse monoclonal antibody ED1 (AbD Serotec, Raleigh, NC) (1:400 dilution overnight at room temperature) followed by a CY3-conjugated goat anti-mouse secondary antibody (Jackson ImmunoResearch, West Grove, PA) (1:500 dilution for 1 h at room temperature) to identify lysosomal antigens in macrophages and giant cells; (2) a rabbit polyclonal antibody against collagen IV (Research Diagnostics, Concord, MA) (1:400 dilution overnight at room temperature) followed by a CY2-conjugated goat anti-rabbit secondary antibody (1:500 dilution for 1 h at room temperature) to detect angiogenesis. A second set of tissue sections was double immunolabeled with: (1) a mouse monoclonal antibody ED1 (1:400 dilution overnight at room temperature) followed by a CY3-conjugated goat anti-mouse secondary (1:500 dilution for 1 h at room temperature) to identify lysomal antigens in macrophages and giant cells; (2) a rabbit polyclonal antibody against fibronectin (Chemicon, Temecula, CA) (1:400 dilution overnight at room temperature) followed by a CY2conjugated goat anti-rabbit secondary antibody to identify fibrotic scarring. All antibodies and Hoechst stain were diluted in phosphate-buffered saline containing 0.3% Triton X-100. Several sections of the tissue were examined for each animal using an epifluorescent Olympus IX81 inverted microscope (Olympus, Melville, NY).
3. Results and discussion Blends of poly(o-CPX) (15 wt.%) (Fig. 1A) and PLAA (85 wt.%) (Fig. 1B) were investigated for their potential as nerve guidance conduits.
1920
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
Both polymers were blended and extruded as hollow tubes with an outside diameter of 2.5 mm and an inside diameter of 2.3 mm. Fig. 3 shows an example of the outside surface and inside lumen morphology of the melt extruded conduits. Suture holes for in vivo studies were created by heating a 21 gauge needle (Becton Dickinson, Franklin Lakes, NJ) and pressing it through the polymer 1 mm from the end of the conduit. 3.1. In vitro cytotoxicity: fibroblasts Cytocompatibility was demonstrated by two approaches: (i) fibroblast growth in polymer-containing medium to mimic late stage polymer degradation; (ii) primary neuron growth on collagen gel-coated polymer substrates to mimic in vivo conditions. Cytocompatibility of the polymer blend of poly(o-CPX)–PLAA was first evaluated by culturing L929 mouse fibroblasts cells in polymercontaining medium at concentrations of 0.10 and 0.01 mg polymer ml medium1. Cells were cultured over a 3 day period, during which cell proliferation and morphology were observed. The intent of this study was to determine cell viability in medium containing the polymer blend of poly(o-CPX)–PLAA. The best mechanical integrity and handling properties were observed with the blend – the homopolymers did not meet this criterion for an NGC. PLGA was chosen as the control because it is a common co-polymer that is used in FDA-approved devices and in a variety of biomedical implants, including nerve guidance conduits. For polymers at both concentrations fibroblasts demonstrated increased cell proliferation throughout the 96 h studied (Fig. 4). As indicated by pair-wise comparison with Scheffe’s post hoc test, the medium containing polymer degradation products at 0.01 mg polymer ml medium1 did not statistically differ from the of DMSO medium and medium alone controls after 96 h cell culture (P < 0.05). For the 0.10 mg ml1 concentration, significant differences were found between the polymer blend of poly(o-CPX)–PLAA and the medium controls at 72 and 96 h (indicated by an asterisk in Fig. 4). Although lower cell proliferation (cell count) was evident at the higher concentration (0.10 mg ml1) of polymer in comparison with the DMSO medium and medium controls, the morphological images (Fig. 5) illustrate cells with stellate extended filipodia at all concentrations, indicating healthy fibroblast cell growth. 3.2. In vitro primary nerve cell viability To model in vivo conditions a collagen gel was prepared on the polymer surface. The experimental design of the NGC mimetic
environment is shown in Fig. 2. For simplicity a NGC of 2 mm inner diameter and 2.2 mm outer diameter was chosen as the system to model in the tissue culture plate. E8 chick DRG were seeded into the three-dimensional collagen gel and cultured in medium containing nerve growth factor for 7 days. Under bright field microscopy, once seeded into the collagen gel the DRG initially appeared to be a dark mass of diameter Dd with no protruding neurites (Fig. 6A). Neurite extension was measured on days 1, 4, and 7 and compared with the original circumscribed DRG ‘body’ region. A DRG that failed to demonstrate neurite extension after 7 days of culture would be indicative of an environment which is unfavorable for neuron growth, i.e. non-viable. An ever increasing halo of neurons extending from the dense DRG body (Dd) over 7 days culture is indicative of a viable environment/surface for neuron growth. Fig. 6A–C demonstrates the technique utilized to evaluate the effect of the degradation products of the polymer blend of poly(o-CPX)–PLAA on neurite outgrowth. Fig. 6E shows the overall increase in neurite outgrowth compared with the size of the original DRG body for the polymer blend of poly(o-CPX)–PLAA compared with PLGA (50:50) and an uncoated glass control. After 7 days culture the DRGs grew as follows: 827 ± 174% on the glass surface, 776 ± 221% on the PLGA and 641 ± 153% on the polymer blend of poly(o-CPX)–PLAA. As indicated by pair-wise comparison with Scheffe’s post hoc test, significant variance among the initial polymer conditions did not result in statistical differences in the extent of neuritic field extension after 7 days (P < 0.05). No additives were included within the chosen polymer blend of poly(o-CPX)–PLAA, PLGA or glass control to further encourage neurite outgrowth. While this system looks promising, we are pursuing further studies involving patterned surfaces, collagen NGC lumen fillers and additional growth factors (nerve growth factor, fibroblast growth factor, etc.) in an attempt to develop an optimal environment for neurite growth. 3.3. In vivo immune response Following 7 days vivo intramuscular implantation in Sprague– Dawley rats, the polymer blend of poly(o-CPX)–PLAA favorably compared with the FDA-approved PLGA (50:50). No profound immunohistological differences were observed between the tissue surrounding the two polymers, as shown in Fig. 7. Excised tissue was stained for an inflammatory response (ED1, macrophages), fibrous scarring (fibronectin), angiogenesis (collagen IV) and cell nuclei (Hoechst 33258 nucleic acid stain). The
Fig. 3. SEM images of the melt extruded polymer blend of poly(o-CPX)–PLAA conduit. Suture holes were melt pressed with a 21 gauge needle into the material for in vivo studies. Scale bars 1 mm.
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
1921
Fig. 4. Cell viability/proliferation in culture media with polymers at concentration of (A) 0.01 mg polymer ml medium1 and (B) 0.10 mg polymer ml medium1. *Statistically different at P < 0.05.
Fig. 5. L929 mouse fibroblasts after 96 h cell culture (10). (A) Polymer blend of poly(o-CPX)–PLAA at 0.01 mg ml1; (B) PLGA (50:50) at 0.01 mg ml1; (C) polymer blend of poly(o-CPX)–PLAA at 0.1 mg ml1; (D) PLGA (50:50) at 0.10 mg ml1. Scale bars 200 lm.
1922
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
Fig. 6. Bright field images of incremental neurite outgrowth on days 1 (A), 4 (B), and 7 (C) on a polymer blend of poly(o-CPX)–PLAA. (D) Immunostained DRG (as C) against neurofilament 68 and 200. DRG growth was quantified as percentage increase in the neurite halo compared with the original body size (D1, D4 or D7/Dd). Scale bar: (A) 1000 lm; (B–D) 2000 lm. (E) Comparison of the percentage increase in neurite outgrowth compared with the original DRG body size after 7 days culture on a polymer blend of poly(o-CPX)–PLAA, PLGA and glass (control) collagen substrates.
results of the immunocytochemical analysis for ED1, collagen IV and fibronectin is presented in Table 1. Immunohistological staining of the tissue was scored in a range from deficient to very abundant for each molecule. After 7 days the polymer blend of poly(o-CPX)–PLAA and PLGA implants were both surrounded by minimal fibrin matrix (Fibronectin, +). The surrounding tissue also indicated a healthy accepting response, containing as it did normal amounts of fibrin. Compared with the FDA-approved PLGA implant, minimal macrophage migration (ED1, ±) was visible in the polymer blend of
poly(o-CPX)–PLAA implant over the 7 day implantation period. A pronounced inflammatory tissue reaction should result in a much greater concentration of macrophages at the site of implantation and would be clearly visible on the histological sections [22]. The visible tissue reactions were more consistent with local phagocytosis, rather than a large scale inflammatory response. With the ultimate goal of serving as a hollow conduit to support regenerating nerves, this biomaterial must support angiogenesis. The abundance of collagen IV in this immunostaining study for the polymer blend of poly(o-CPX)–PLAA (collagen IV, +++) indi-
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
1923
Fig. 7. Immunohistological staining for an immune response. (A, C) ED1, red; collagen IV, green; cell nuclei, blue. (B, D) Fibronectin, green; cell nuclei, blue. Scale bar 200 lm.
Table 1 Immunocytochemical analysis of subcutaneously implanted polymer rods. ED1, collagen IV and fibronectin were scored in the tissues surrounding the material: , deficient; ±, sparse; +, moderate; ++, clearly expressed; +++, abundant; ++++, very abundant. Polymer rod composition
ED1
Collagen IV
Fibronectin
Poly(o-CPX) [15 wt.%]:PLAA [85 wt.%] PLGA (50:50)
± ±
+++ +++
+ ++
cates that the material does not adversely affect the formation of capillaries surrounding the implant. This further indicates that hollow conduits of the polymer blend of poly(o-CPX)–PLAA will be conducive to providing the well-vascularized environment that is necessary to support nerve regeneration [23].
such polyanhydrides have the capability to have therapeutics chemically incorporated into the polymer backbone with additional labile ester bonds, which could allow for controlled release of bioactives, if desired [18,26]. The polyanhydrides were melt extruded into conduits of appropriate size to evaluate their in vivo regenerative properties and characteristics using the sciatic nerve model in Sprague– Dawley rats. The polyanhydride blends evoked a mild foreign body reaction, with biocompatibility comparable with PLGA implants. Future studies will involve additional functionality arising from an impregnated collagen matrix. Based upon these positive results, conduits fabricated from the polymer blend of poly(oCPX)–PLAA will be more extensively evaluated as tissue engineering devices. Acknowledgements
4. Conclusion NGCs comprised of polyanhydrides such as PLAA and poly(oCPX) are advantageous compared with the biologically derived, commercially available NGCs. NGCs made from collagen or other complex biological materials have the potential for batch-to-batch variability [24], whereas synthetic NGCs (such as the polyanhydrides) have a well-defined chemical composition that directly controls polymer degradation, as well as device degradation [9]. In addition, these polyanhydrides have exceptional thermal properties, with glass transition temperatures (Tg) well above the physiological temperature, which is desirable to avoid loss of mechanical integrity when plasticizers or admixtures of other polymers are utilized [9]. Poly(o-CPX) and related polyanhydrides were chosen since their components and resulting degradation products are known to be biocompatible [10–13,25]. Additionally,
This research was supported by US Army contract no. W81XWH-04-2-0031. The authors thank Olex Hnojewyj (BTI, Menlo Park, CA) for providing melt extruded conduits, Natasha Piracha (Department of Cell Biology and Neuroscience, Rutgers University) for assisting with the fibroblast cytotoxicity study and Susan Harris (Department of Pharmacology, Robert Wood Johnson Medical School) for her help with the immunohistological study. References [1] IJkema-Paassen J, Jansen K, Gramsbergena A, Meek M. Transection of peripheral nerves, bridging strategies and effect evaluation. Biomaterials 2003;25:1583–92. [2] Committee on Capturing the Full Power of Biomaterials for Military Medical Needs, National Research Council. Capturing the full power of biomaterials for military medical needs. Washington, DC: National Academic Press; 2004.
1924
J. Griffin et al. / Acta Biomaterialia 6 (2010) 1917–1924
[3] Ichihara S, Inada Y, Nakamura T. Artificial nerve tubes and their application for repair of peripheral nerve injury: an update of current concepts. Injury: Int J Care Injured 2008;39S4:S29–39. [4] Maquet V, Martin D, Malgrange B, Franzen R, Schoenen J, Moonen G, et al. Peripheral nerve regeneration using bioresorbable macroporous polylactide scaffolds. J Biomed Mater Res 2000;52:639–51. [5] Dellon A, Mackinnon S. An alternative to the classical nerve graft for the management of the short nerve gap. Plast Reconstr Surg 1988;82:849–56. [6] Evans G. Challenges to nerve regeneration. Semin Surg Oncol. 2000;19:312–8. [7] den Dunnen W, Stokroos I, Blaauw E, Holwerda A, Pennings A, Robinson P, et al. Light-microscopic and electron-microscopic evaluation of short-term nerve regeneration using a biodegradable poly (dl-lactide–caprolacton) nerve guide. J Biomed Mater Res A 1998;31:105–15. [8] Nicoli Aldini N, Perego G, Cella G, Maltarello M, Fini M, Rocca M, et al. Effectiveness of a bioabsorbable conduit in the repair of peripheral nerves. Biomaterials 1996;17:959–62. [9] Anastasiou T, Uhrich K. Novel polyanhydrides with enhanced thermal and solubility. Macromolecules 2000;33:6217–21. [10] Leong K, Brott B, Langer R. Bioerodable polyanhydride as drug-carrier matrices I: characterization, degradation, and release characteristics. J Biomed Mater Res 1985;19:941–55. [11] Domb A, Langer R. Solid state and solution stability of poly(anhydrides) and poly(esters). Macromolecules 1989;22:2117–22. [12] Chasin M, Lewis D, Langer R. Polyanhydrides for controlled release. Biopharm Manuf 1988;1:33–46. [13] Campo C, Anastasiou T, Uhrich K. Polyanhydrides: the effects of ring substitution changes on polymer properties. Polym Bull 1999;42:61–8. [14] Uhrich K, Cannizzaro S, Langer R, Shakesheff K. Polymeric systems for controlled drug release. Chem Rev 1999;99:3181–98.
[15] Duncan E, Lemons J, Mayesh J, Saha P, Saha S, Scranton M. New Products and standards in biomaterials science. New York: Elsevier Academic Press; 2004. [16] Schmeltzer R, Schmalenberg KE, Uhrich KE. Biomacromolecules 2005;6:359–67. [17] Carbone A, Song M, Uhrich K. Iodinated sacilcylate-based poly(anhydrideesters) as radiopaque biomaterials. Biomacromolecules 2008;9:1604–12. [18] Fischer D, Li Y, Ahlemeyer B, Krieglstein J, Kissel T. In vitro cytotoxicity testing of polycations: influence of polymer structure on cell viability and hemolysis. Biomaterials 2003;24:1121–31. [19] International Standards Organization. Biological evaluation for medical devices-part 5: tests for cytotoxicity in vitro methods. ISO-10993-5 (EN 30993-5);1992. [20] Luduena RF, Woodward DO. Isolation and partial characterization of alpha and beta-tubulin from outer doublets of sea-urchin sperm and microtubules of chick-embryo brain. Proc Natl Acad Sci USA 1973;70:3594–8. [21] Schmalenberg KE, Buettner HM, Uhrich KE. Microcontact printing of proteins on oxygen plasma-activated poly(methyl methacrylate). Biomaterials 2004;25:1851–7. [22] Jansen K, van der Werff JFA, van Wachem PB, Nicolai J-PA, de Leij LFMH, van Luyn MJA. A hyaluronan-based nerve guide: in vitro cytotoxicity, subcutaneous tissue reactions, and degradation in the rat. Biomaterials 2004;25:483–9. [23] Hobson MI, Brown R, Green CJ, Terenghi G. Inter-relationships between angiogenesis and nerve regeneration: a histochemical study. Br J Plast Surg 1997;50:125–31. [24] Lutolf M. Integration column: artificial ECM: expanding the cell biology toolbox in 3D. The Royal Soc Chem: Integr Biol 2009;1:235–41. [25] Merck Research. The Merck Index. Whitehouse Station, NJ: Merck Research Laboratories; 2001. [26] Prudencio A, Schmeltzer R, Uhrich K. Using the ‘‘linker” to control polymer properties. Macromolecules 2005;38:6895–901.