Pursuant to 35 USC §120, this application is a continuation of prior U.S. application Ser. No. 14/694,466, filed Apr. 23, 2015, which is a continuation of prior U.S. application Ser. No. 14/496,742, filed Sep. 25, 2014, now U.S. Pat. No. 9,101,695, which is a continuation of prior U.S. application Ser. No. 14/177,888, filed Feb. 11, 2014, now U.S. Pat. No. 8,858,966, which is a continuation of prior U.S. application Ser. No. 13/741,901, filed Jan. 15, 2013, now U.S. Pat. No. 8,685,434, which in turn is a divisional application and claims the benefit of prior U.S. application Ser. No. 11/789,538, filed Apr. 25, 2007, now U.S. Pat. No. 8,377,463, which in turn claims the benefit of prior U.S. Provisional Application 60/794,986, filed Apr. 25, 2006. Each of these applications are incorporated by reference in its entirety.
- Top of Page
OF THE INVENTION
Many spinal cord injuries (SCIs) are a result of the spinal cord being compressed, not cut. Insult to the spinal cord often results in vertebrae, nerve and blood vessel damage. Bleeding, fluid accumulation, and swelling can occur inside the spinal cord or outside the spinal cord but within the vertebral canal. The pressure from the surrounding bone and meninges structure can further damage the spinal cord. Moreover, edema of the cord itself can additionally accelerate secondary tissue loss. There is considerable evidence that the primary mechanical injury initiates a cascade of secondary injury mechanisms including excessive excitatory neurotransmitter accumulation; edema formation; electrolyte shifts, including increased intracellular calcium; free radical production, especially oxidant-free radicals; and eicosanoid production. Therefore, SCIs can be viewed as a two-step process. The primary injury is mechanical, resulting from impact, compression or some other insult to the spinal column. The secondary injury is cellular and biochemical, wherein cellular/molecular reactions cause tissue destruction. By interrupting this second process and diffusing any compression resulting from the primary mechanical lesion, as well as any cord edema, healing is expedited.
As discussed above, spinal cord injury involves not only initial tissue injury, but also devastating secondary injuries. These pathological events, caused by excitotoxicity, free-radical formation and lack of neurotrophic support, include glial scarring, myelin-related axonal growth inhibition, demyelination, secondary cell death such as apoptosis. For example, oligodendrocyte death continues for weeks after many SCIs. An environment antagonistic to axonal regeneration is subsequently formed. In addition to damaged regeneration pathways, reflexia hyperexcitability and muscle spasticity, there are further complications of respiratory and bladder dysfunction, for example. Over time, muscle mass is lost as a result of loss of innervations and non-use. The end result of these spinal cord insults invariably is lost function, the extent of which is determined by the severity of the spinal cord primary lesion as well as by secondary injuries. Even in the case of incomplete motor function loss, common problems include posture, reduced walking speed, abnormal balance and gait, and lack of sufficient weight-bearing.
Surgical decompression of the spinal cord is often used to relieve any pressure from surrounding bone (by removing fractured or dislocated vertebrae or disks). However, the timing of surgical decompression has been a controversial topic. While rat studies have shown early decompression to reduce secondary injury, the results in human clinical trials have been less than consistent. It has been difficult to determine a time window for the effective application of surgical decompression intervention in the clinical setting. Furthermore, there are no technologies which can be used to effectively control the increase in intra-parenchyma pressure resulting from the primary SCI. The absence of such a technology renders surgical decompression surgery, in many cases, ineffective. The removal of bone and soft tissue structures do not address the underlying problem of secondary intrinsic pressure at the SCI site. Therefore, there exists a need to provide alternative devices and methods to impede the process that drive secondary injury at the primary spinal cord injury site. These alternative methods can be used to complement decompression surgical protocols.
There has been scant, if any, therapeutic attention given to the intrinsic nature of the injured/compressed spinal cord (i.e. the injured/compressed cord itself). As mentioned above, decompression surgery is directed to the extrinsic nature of the injury (i.e. removal of bone or fluid surrounding, and causing, the injury) in hopes of alleviating consequences of intra-tissue pressure build-up. Secondary injury will often impede the nerve regeneration and/or nerve regrowth process. Consequently, there exists a need for devices and methods that alleviate the primary spinal cord injury from, for example, secondary tissue destruction, edema formation, and an influx of inflammatory factors.
Furthermore, it is well known that penetrating spinal cord injuries (SCIs) are the most deadly neurotrauma encountered by people. Reports on combat related open wound SCIs during the Vietnam war indicate that this type of injury leads to close to 100% lethality. While there have been advances in the protective ability of bullet-proof vests, the neck region of persons wearing many of today's vests is often vulnerable to many high velocity weapons. More than 90% of SCIs are initially diagnosed as “incomplete,” wherein the injury does not result in complete severing of the spinal cord. Technology which can protect the spared tissue and promote endogenous healing and repair will mitigate functional deficits resulting from both penetrating and contusion traumatic SCIs.
- Top of Page
OF THE INVENTION
Certain embodiments of the present invention are directed to biocompatible polymeric materials which can be fabricated into “mini-tubes,” or “tubular articles.” These mini-tubes can be used to treat any localized SCI. In one embodiment, the mini-tube is inserted into the epicenter of the injury, wherein the hollow tube runs through the injury site. See FIG. 1. The mini-tube can be inserted through a surgical incision made rostral or caudal to the lesion to be treated. The mini-tube creates a new interface within the compressed spinal cord parenchyma. This new interface relieves the site of pressure and protects tissue that has been spared from injury. Pressure resulting from the compression force exerted on the cord is alleviated by (1) diffusing or redirecting the force down the surface of the mini-tube and away from the initial compressed site, and (2) absorbing the compression energy into the biocompatible material of the mini-tube. See FIG. 1. Furthermore, by providing a structure between the injured site and surrounding tissue (the new interface), inflammation may be mitigated in the adjacent area where functionally relevant residual cord tissue can be spared.
In another embodiment, the present invention relates to biocompatible polymers fabricated into hollow mini-tubes, or tubular articles, having an inner surface, an outer surface and two opposing ends. The mini-tubes may be fabricated into any geometrical shape and size. For example, the size and the shape of the mini-tube may be varied in order to deliver more effective relief. A thin, elongated cylinder is one possible configuration, but other shapes, such as elongated rectangular tubes, spheres, helical structures, and others are possible. Additional alterations in configuration, such as the number, orientation, and shape of the mini-tubes may be varied in order to deliver more effective relief. For instance, the mini-tubes may be rectangular, or any other useful shape, and may be distributed along and/or around epicenter of the spinal cord injury. The size will vary accordingly with the spinal cord lesion to be treated. The mini-tube can be smaller than, the same size as, or longer than the lesion to be treated. In preferred embodiment, the mini-tube will be longer than the length of the injured site. In another preferred embodiment, the length of the mini-tube to be surgically implanted will be approximately between 1.2 and 3 times the length of the injured site or lesion running lengthwise along the spinal cord. In yet another preferred embodiment, the mini-tube will extend beyond the caudal and rostral sides of the injured site at a distance of approximately ¼ the length of the injured site. In a preferred embodiment the mini-tube will extend equally beyond the caudal and rostral sides of the injured site.
The diameter of the mini-tube (outer surface to outer surface; or “outside diameter”) can range from 0.1 microns to 10 millimeters. In a preferred embodiment, the overall diameter of the mini-tube (outer surface to outer surface) is between about 5 and 200 microns. In other embodiments the diameter of the mini-tube (outer surface to outer surface) is between about 20 and 200 microns, between about 50 and 175 microns, between about 100 and 200 microns, and between about 150 and 300 microns. In another embodiment, the diameter of the mini-tube (outer surface to outer surface) is between about 0.5 millimeters and 20 millimeters. In other embodiments, the diameter of the mini-tube (outer surface to outer surface) is between about 1 millimeter and 10 millimeters, between about 1 millimeter and 5 millimeters, and between about 1 millimeter and 3 millimeters.
The diameter of the mini-tube (inner surface to inner surface; or the “lumen diameter”) can also range from microns to millimeters. In a preferred embodiment, the diameter of the mini-tube (lumen diameter) is between about 5 and 200 microns. In other embodiments the diameter of the mini-tube (lumen) is between about 20 and 200 microns, between about 50 and 175 microns, between about 100 and 200 microns, and between about 150 and 300 microns. In another embodiment, the diameter of the mini-tube (lumen) is between about 0.5 millimeters and 15 millimeters. In other embodiments, the diameter of the mini-tube (lumen) is between about 1 millimeter and 10 millimeters, between about 1 millimeter and 5 millimeters, and between about 1 millimeter and 3 millimeters.
In another embodiment of the present invention, formable, moldable, biocompatible polymeric materials are disclosed herein. Advantageously, the polymeric material may be fabricated as a putty. By “putty” it is meant that the material has a dough-like consistency that is formable or moldable. These materials are sufficiently and readily moldable and can be formed into flexible three-dimensional structures or shapes complementary to a target site to be treated.
In yet another embodiment, the biocompatible polymeric materials of the present invention can be fabricated into readily formable or moldable bandages, or neuropatches. In one embodiment, a SCI is localized and the bandage or neuropatch is hand-formed to complement the injured site (for example, a hemi-sected spinal cord). The hand formed bandage is then implanted into the epicenter of the injury, wherein the bandage fills in the injury site. The implanted bandage bridges any gap formed by the spinal cord lesion and functions as an artificial pathway, nurturing regrowing neurons, reorganizing neurites and helping to form functional synapses. This new bandage interface allows for interactions between endogenous neural cells (including neural stem cells, if incorporated onto the bandage) and the inhibitory molecule-free polymer implant environment to promote cell survival. Furthermore, by providing a structure between the injured site and surrounding tissue (the new interface), inflammation may be mitigated in the adjacent area where functionally relevant residual cord tissue can be spared.
In another embodiment, the present invention relates to biocompatible polymeric bandages, which can be readily fabricated/formed into any shape and size, comprising a single polymeric scaffold having an inner surface and an outer surface. See example 15. The formed bandages may be fabricated into any geometrical shape and size. For example, the size and the shape of the bandage may be varied in order to deliver more effective relief. A thin, elongated bandage is one possible configuration, but other shapes, such as elongated rectangular bandages, spheres, helical structures, and others are possible. Additional alterations in configuration, such as the number, orientation, and shape of the bandages may be varied in order to deliver more effective relief. For instance, the bandages may be rectangular, or any other useful shape, and may be distributed within and/or around epicenter of the spinal cord injury. In addition, the bandage may have a textured surface including a plurality of pores and/or microgrooves on its inner and/or outer surface. In one embodiment, the pores have diameters between about 0.5 μm to 4 μm and depths of at least 0.5 μm. The microgrooves may have widths of between about 0.5 μm and 4 μm and depths of at least 0.5 μm. The sizes of the bandage, and the sizes and diameters of its pores and microgrooves, will vary accordingly with the spinal cord lesion to be treated. The pores and/or microgrooves on the inner and/or outer surface may be seeded with one or more medicinal agents, for example human neuronal stem cells to provide cellular replacement and trophic support. In preferred embodiment, the bandage will act as a filler (i.e. fill the lesion) after implantation of the bandage within the lesioned area of the spinal cord, for example. In one embodiment, the bandage inner surface is flush with the lesioned spinal cord, i.e. contacts the lesion, when it is implanted.
Biocompatible polymers for the fabrication of the herein described mini-tubes and formable bandage or neuropatch articles are well-known in the art. In a preferred embodiment, the biocompatible polymers are biodegradable (for example, PLGA). As used herein, biodegradable and erodible are used interchangeably. Examples of biocompatible polymers that are biodegradable include, but are not limited to, biodegradable hydrophilic polymers such as polysaccharides, proteinaceous polymers, soluble derivatives of polysaccharides, soluble derivatives of proteinaceous polymers, polypeptides, polyesters, polyorthoesters, and the like. The polysaccharides may be poly-1,4-glucans, e.g., starch glycogen, amylose and amylopectin, and the like. Preferably, biodegradable hydrophilic polymers are water-soluble derivatives of poly-1,4-glucan, including hydrolyzed amylopectin, hydroxyalkyl derivatives of hydrolyzed amylopectin such as hydroxyethyl starch (HES), hydroxyethyl amylase, dialdehyde starch, and the like. Proteinaceous polymers and their soluble derivatives include gelation biodegradable synthetic polypeptides, elastin, alkylated collagen, alkylated elastin, and the like. Biodegradable synthetic polypeptides include poly-(N-hydroxyalkyl)-L-asparagine, poly-(N-hydroxyalkyl)-L-glutamine, copolymers of N-hydroxyalkyl-L-asparagine and N-hydroxyalkyl-L-glutamine with other amino acids. Suggested amino acids include L-alanine, L-lysine, L-phenylalanine, L-leucine, L-valine, L-tyrosine, and the like.
Definitions or further description of any of the foregoing terminology are well known in the art and may be found by referring to any standard biochemistry reference text such as “Biochemistry” by Albert L. Lehninger, Worth Publishers, Inc. and “Biochemistry” by Lubert Stryer, W. H. Freeman and Company, both of which are hereby incorporated by reference.
The aforementioned biodegradable hydrophilic polymers are particularly suited for the methods and compositions of the present invention by reason of their characteristically low human toxicity and virtually complete biodegradability. Of course, it will be understood that the particular polymer utilized is not critical and a variety of biodegradable hydrophilic polymers may be utilized as a consequence of the novel processing methods of the invention.
Electrical signals in the form of action potentials are the means of signaling for billions of cells in the central nervous system. Numerous studies have shown that this electrical activity is not only a means of communication, but also necessary for the normal development of the nervous system and refinement of functional neural circuits. In the case of spinal cord injury, cell-to-cell communication may be interrupted and the mechanisms of normal neurological development imply that electrical activity should be part of the restoration of functional connections. Such activity is important for the survival of existing cells and the incorporation of any transplanted cells (such as neural stem cells) into working circuits. In an embodiment of the present invention, single and double layer scaffolds and minitubes are fabricated from synthetic biomaterials and are capable of conducting electricity and naturally eroding inside the body. In an exemplary embodiment, the single scaffold, double scaffold, or minitube comprises a biocompatible polymer capable of conducting electricity is a polypyrrole polymer. Polyaniline, polyacetyline, poly-p-phenylene, poly-p-phenylene-vinylene, polythiophene, and hemosin are examples of other biocompatible polymers that are capable of conducting electricity and may be used in conjunction with the present invention. Other erodible, conducting polymers are well known (for example, see Zelikin et al., Erodible Conducting Polymers for Potential Biomedical Applications, Angew. Chem. Int. Ed. Engl., 2002, 41(1):141-144). Any of the foregoing electrical conducting polymers can be applied or coated onto a malleable or moldable article. The coated article can be also be used as a bandage, or neuropatch, as described herein.
In a preferred embodiment the biodegradable and/or bioabsorbable polymer contains a monomer selected from the group consisting of a glycolide, lactide, dioxanone, caprolactone, trimethylene carbonate, ethylene glycol and lysine. By the terminology “contains a monomer” is intended a polymer which is produced from the specified monomer(s) or contains the specified monomeric unit(s). The polymer can be a homopolymer, random or block co-polymer or hetero-polymer containing any combination of these monomers. The material can be a random copolymer, block copolymer or blend of homopolymers, copolymers, and/or heteropolymers that contains these monomers.
In one embodiment, the biodegradable and/or bioabsorbable polymer contains bioabsorbable and biodegradable linear aliphatic polyesters such as polyglycolide (PGA) and its random copolymer poly(glycolide-co-lactide) (PGA-co-PLA). The FDA has approved these polymers for use in surgical applications, including medical sutures. An advantage of these synthetic absorbable materials is their degradability by simple hydrolysis of the ester backbone in aqueous environments, such as body fluids. The degradation products are ultimately metabolized to carbon dioxide and water or can be excreted via the kidney. These polymers are different from cellulose based materials, which cannot be absorbed by the body.
The molecular weight (MW) of the polymers used in the formable articles of the presently described invention can vary according to the polymers used and the degradation rate desired to be achieved. In one embodiment, the average MW of the polymers in the fabricated bandage is between about 1,000 and about 50,000. In another embodiment, the average MW of the polymers in the fabricated bandage is between about 2,000 and 30,000. In yet another embodiment, the average MW is between about 20,000 and 50,000 for PLGA and between about 1,000 and 3,000 for polylysine.
The herein described mini-tubes and formable articles may be incorporated with any number of medically useful substances. In a preferred embodiment, the inner and/or outer surfaces of the mini-tube is seeded with stem cells; for example, mesenchymal and/or neuronal stem cells, wherein the cells are deposited onto the inner (lumen in the case of the mini-tubes) and/or outer surface(s). See FIG. 3. The incorporation of stem cells provide for trophic support and/or cellular replacement at the site of injury.
In another embodiment, the foregoing described polymeric articles are used in methods for providing controlled tissue healing. These methods comprise, for example, implanting into a target compression injury site in an animal, a system for controlled tissue healing, the system comprising a biodegradable and/or bioabsorbable polymeric hollow tube. The target injury site may be any injury that is susceptible to secondary tissue injury, including but not limited to: glial scarring, myelin inhibition, demyelination, cell death, lack of neurotrophic support, ischemia, free-radical formation, and excitotoxicity. In one embodiment, the injury to be treated is a spinal cord injury, wherein the spinal cord is compressed. The herein described methods may be used in conjunction with decompression surgery; for example, concomitant with decompression surgery, prior to decompression surgery, or subsequent to decompression surgery.
In another embodiment, the foregoing described polymeric articles are used in methods for treating a compression spinal cord injury comprising implanting into a target compression injury site in an animal a biodegradable and/or bioabsorbable polymeric hollow tube. The spinal cord injury may be susceptible to secondary tissue injury, including but not limited to: glial scarring, myelin inhibition, demyelination, cell death, lack of neurotrophic support, ischemia, free-radical formation, and excitotoxicity. The herein described methods may be used in conjunction with decompression surgery; for example, concomitant with decompression surgery, prior to decompression surgery, or subsequent to decompression surgery.
DESCRIPTION OF THE FIGURES
FIG. 1A and FIG. 1B. Two schematic representations (A and B) of the polypyrrole scaffold inserted around the center of the lesion area in order to protect surrounding tissues.
FIG. 2. Electrodeposition of erodible PPy to form mini-tube scaffolds.
FIG. 3A to FIG. 3D. SEM images of microfabricated PPy tubes. A. Murine neural stem cells seeded inside of a 600 μm inner diameter tube (150×). B. High-magnification (350×) view of 25 μm inner diameter tube. Rough surface texture is a result of low electrodeposition temperature (4° C.). C. Lower magnification (150×) view of a 25 μm inner diameter tube created with a smooth surface texture by electrodeposition at 24° C. D. Higher magnification (500×) view of same tube as in C.
FIG. 4. MRI shows reduced fluid filled cyst (appears bright white in the T2 weighted MR image) formation in rodents treated with a PPy scaffold (shown on right) relative to untreated control (shown at left).
FIG. 5. Open-field locomotor scores for polypyrrole minitube-implanted rats (n=8) and lesion-control rats (n=11).
FIG. 6. BBB open-field walking scores for the four groups on the ipsilateral, lesioned side. Hindlimbs were assessed independently to determine the degree of asymmetry. The rate of improvement for the scaffold-treated group was significantly greater than the rate for the stem cells-alone (P<0.001) and lesion-control groups (P<0.004; two-way repeated measures of ANOVA; N=12 each group).