WO2013177147A2 - Copolymer-xerogel nanocomposites useful for drug delivery - Google Patents

Abstract

Biocompatible, biodegradable eopolyraer-xerogel nanocoraposites contain a biodegradable, biocompatible copolymer, silica nanoparticles and one or more therapeutic agents, that are useful for wound dressings, medical devices and drug delivery applications.

Description

COPOLYMER-XEROGEL A OCOMPOSITES USEFUL FOR DRUG DELIVERY

CROSS-REFERENCE TO RELATED APPLICATIONS

ίϊ01 This application is a continuation-in-part under 35 U.S.C § J 20 of U.S. Patent Application No,. 12/476.009, filed on June 1 , 2009, which claims the benefit of priority under 35 U.S.C § 1 19(e) of US Provisional Application No. 61/057,642, filed on May 30, 2008. This application also claims the benefit o priority under 35 U.S.C 1 19(e) of US Provisional Application No. 6.1/649,643, filed on May 21 , 2012. The disclosures of all of the above documents are incorporated herein by reference in their entireties.

STATEMENT OF GOVERNMENT RIGHTS

(08021 Research leading to the disclosed inventions was funded, in part, by the United States Army, CDM.RP grant number WS i WH -07-1 -0438. Accordingly, the United States Government has certain rights in the invention.

FIELD OF TECHNOLOGY

[00031 The present invention is directed to narioco^'rnposites useful for drug delivery applications which provide controlled delivery of therapeutic agents, including drug "depots", wound dressings, stents and tissue scaffolds.

BACKGROUND OF THE INVENTION

[0004j There is an enormous unmet need for biomaterials than, can provide effective local deli very of therapeutic agents with controlled release kinetics. These biomaterials must, be able to meet challenging requirements for both mechanical and biological functionality in such criti-cal applications as implantable drug delivery depots, wound dressings and stents. The challenge is particularly acute for hydrophobic drugs with limited aqueous solubility that limits formulation concentrations and hence limits drug diffusion to target substrates such as pathogenic biofllm infections in chronic wounds and on orthopedic fracture fixation devices, peripheral nerves involved in chronic pain syndromes, solid tumors or cardiovascular stents experiencing restenosis.

(00051 Biodegradable composites that combine the processability and viseoelasticity of biodegradable organic polymers with the mechanical strength of biodegradable ceramic filler materials offer significant potential, to meet these biomaterial performance requirements when, as is often the case, the properties of polymers or ceramics alone are inadequate. High mechanical.

- I - strength is imparted to composites through effective load transfer between the continuous polymer matrix and the discontinuous inorganic particles. This requires effective mterfacial bonding, either physical or covalent, between the ceramic and polymeric components. The interfacial properties of -polymer-inorganic composites also exert a strong influence on other properties including gas permeability, water uptake, drug release kinetics and cellular responses. Both the physical and chemical properties of bio aterials ca strongly affect the performance of and biological responses to drug delivery devices, wound dressings and tissue engineering scaffolds. When the polymers used to form the composites are biodegradable, the biodegradatio rates are significantly altered by the presence of the inorganic components, their concentration in the composite matrix and whether they are physically or covalent!y bonded to the organic polymer components,

iMMIoj Early treatment of bodily wounds is generally limited to hemostasia and administration of pain medication. For example, for severe battlefield, wounds, the initial treatment consists of applying hemostatic agents such as eh osan bandages and Quick-Clot ^{, M} zeolite. Wound dressings being deployed on the battlefield, however, are not designed to deliver pain medication. Existing injectable hydrogelx, such as Durecfs SABER.' ^M system for delivery of bupivacaine, are not designed for battlefield applications because they cannot withstand the conditions that occur during transport of patients to medical facilities. Further, traditional anesthetic delivery systems such as direct injection, epidural catheters, and intra-articular indwelling catheters are not designed or convenient for battlefield applications. These modalities of local delivery of analgesics are not designed to withstand the conditions present during the transport of patients, have limited efficacy, have potential adverse clinical complications, and require highly trained medical personnel. As a .result, on-the-fleld pain treatment of wounds is usually delivered in the form of systemic morphine injections, which have numerous unwanted and serious side effects. Similarly, in the case of wound infections infections, effective treatment is exacerbated by multi-drug resistant strains of microrga isms (MDRO's) such as rnethiciilin- resistant. Staphylococcus aureus (MRS A), Pseudomonas aeruginosa, Enteroc.ocaisfae.cium, Escherichia cob, Klebsiella pneumoniae, Enterobaeter species, and Acmeiabact.er bamnanni. Topical delivery of antimicrobials and other therapeutic agents is advantageous because systemic toxicity is avoided and high local concentrations can be achieved thai are often necessar to eradicate drug-resistant microbial biofiJms, particularly in cases where systemic delivery resulting from ischemia at wound sites can limit other parenteral or oral drug delivery routes. For chronic open wounds, it is widely recognized that a moist environment promotes wound healing and this is generally accomplished in the clinic by ihe application of occlusi ve polymeric foydrogel wound dressings,

j 0097 J Severe combat wounds, particularly blast wounds resulting from explosive devices, involve substantial tissue damage that produces sustained and often intense levels of pain throughout and beyond the. early tissue healing process. If the pain is left untreated, the pain signals can be imprinted in the central nervous system, resulting in chronic pain, Continuous peripheral nerve block by local delivery of anesthetics immediately following trauma or surgical procedures has been suggested to have potential to prevent chronic pain, including syndromes such as phantom, limb pain. Thus, it is highly desirable to provide controlled delivery of local anesthetics directly to a wound,

(00981 In some instances, severe combat wounds, particularly blast wounds, also result in compartment syndrome. Compartment syndrome occurs when elevated intramuscular pressure decreases vascular perfusion of a muscle compartment to a point no longer sufficient to maintain viability of the muscle and neural tissue contained within the compartment.

Compartment syndromes can result from multiple types of injuries including orthopedic

(traumatic), vascular, iatrogenic, and soft tissue. Blast injuries now seem to fall in this category as well. In some cases the blast injury can only be part of soft tissue injury or it can be a combination of the other etiologies including components of orthopedic, vascular and/or soft tissue, More recently with the increasing number of casualties from blast injury, it is

hypothesized that the blast causes a direct injury to the muscle thai result in swelling and secondary compartment syndrome.

[0099} Generally, in compartment syndromes, there is an increasing pressure within a. tissue compartment that needs to be released -as soon as possible, often within 4 to 6 hours.

Compartment syndromes must be treated early in the time line of wound care that begins at the battlefield and ends in the hospital. If a compartment syndrome is not diagnosed early, a Volk- mann contracture can occur with massive loss of all tissues within the compartment. Untreated compartment syndrome can lead to tissue necrosis, permanent functional impairment, renal failure, and death. However, the standard diagnosis of compartment syndrome by clinical signs - including myoneural pain with passive stretch, paresthesia, and paresis - is often masked by other injuries in patients with blast injuries who suffe polytrauma, [0010| The treatment of compartment syndrome requires the release of the fascia that enclose the compartments within the first, three to six hours to prevent irreversible injury to the nerves and muscles. Once the compartments are released the open wounds are treated with dressings to prevent infection and protect the wound. In some cases a specialized Vacuum

Assisted Closure System is used to cover and protect the wound. The open wounds are then kept dressed for 48 to 72 hours until the patients are returned, to the operating room for a second look to allow further debridement of non-viable muscle tissue if Indicated. Fasciotomies,. however, extend hospital stays and change a closed injury to an. open injur)', greatly increasing the chance of ^'infection. Further, there is s me debate about the criterion for performing a fasciotomy, with recommendations varying from prophylactic fasciotomy at norma! pressure to finding a pressure from 30 mm H to 45 mm Hg.

{Mill ^ has been suggested that impeding the early cellular events leading to ischemia and pressure build up in the compartment can. be the first line of defense. Thus, it, would be desirable to provide controlled deli very of therapeutic agents to prevent the late-stage problems of compartment syndrome and initiate regeneration of healthy tissue.

[0012 | There remains a great need for materials for the treatment of wounds that effect the controlled release of pharmaceutically active molecules. Controlled release focuses on delivering biologically active agents locally over extended time periods. The site specificity of the delivery reduces the potential side effects that can be associated with general administration of drugs through oral, or parenteral therapy. Prevalent mechanisms for the delivery of biological agents by controlled release devices are either resorption of the drug carrier material or diffusion. The resorption of these devices can, however, cause an inflammatory tissue response which interferes with the treatment sought for with the biomotecules.

BRIEF SUMMARY OF THE INVENTION

[0 1-3 | It has now been discovered that highly tunable mechanical and controlled drug delivery properties are accessible with novel biodegradable nanocomposites prepared by non~cova.te.nt binding of si lica xerogels and biodegradable, biocompatible polymers,

[001 } Further, it has also been discovered thai sustained controlled release of clinically significant drugs, including antibiotics and local anaesthetics, can be obtained, from these nanocomposites. Such nanocomposites are thus attractive bioactive biomatertals for appl.icalio.ns including w un dressings, tissue engineering applications;, cardiovascular stents and nerve guides or conduits. One embodiment of the present invention is directed to a

copolymer-xeroge! nanocomposite, comprising a biodegradable, biocompatible copolymer, silica nanoparticles and one or more therapeutic agen ts,

{08151 A further embodiment of the present invention, is directed to a method of forming a therapeutic agent-loaded copolymer-xerogel nanocomposite, comprising the steps of :

(a) providing a silica sol;

(b) adding said silica sol to a po!y(desamin.otyrosyi tyrosine ester-eo-FEG carbonate) to form a mixture;

(c) adding one or more therapeutic agents to said .mixture to .form a drug mixture; and

(d) removing the solvents from said drug mixture to f rm the therapeutic agent-loaded copolymer-xeroge! nanocoraposite.

{00161 Another embodiment of the present invention provides drag depots, wound dressings, tissue scaffolds, cardiovascular stents and nerve guides containing the copolymer- xeroge! nanoeomposites. The devices are adapted to bind and release therapeutic agents, thereby providing controlled delivery of the therapeutic agents for healthcare applications.

[0017J One embodiment of the invention is directed to a copolymer-xerogel nanocomposite, comprising a biodegradable, biocompatible copolymer, silica nanoparticles and one or more therapeutic agents, in a preferred embodiment, the biodegradable, biocompatible copolymer has a molecular weight greater than 20,000 Da! ons. In a further embodiment, the therapeutic agent comprises a compound selected from the group consisting of antibiotics, local anesthetics, and combinations of two or more thereof. In one embodiment, the therapeutic agent is rifempiefn and/or bupivacaine.

{08181 ϊ» another embodiment of the above copolymer-xeroge! nanocomposite, live biodegradable copolymer comprises a copolymer of tyrosme-pofy(alkyl.ene g3yeol}-derived poiyfether carbonate). In one embodiment, the biodegradable copolyme comprises a polycarbonate comprising desaminotyrosyl tyrosine ester and poiyietnySene glycol). In. a further embodiment, the above copolymer-xerogel nanocomposite is adapted to provide controlled release of the therapeutic agent(s). [0019| Another embodiment of the invention is directed to a method of forming a therapeutic agent-loaded eopolymer-xerogel nanocomposite, comprising:

(a) providing a silica sol;

(b) adding said silica sol to a po! {desaras.iiot rosyi tyrosine ester-co-PEG carbonate} to form a mixture;

(c) adding one or more therapeutic agents to said mixture to form a drug mixture; and

(d) removing the solvents from said drug mixture to form the therapeutic agent-loaded eopolymer-xerogel nanocomposite.

[ 020| In a further embodiment of the method the therapeutic agent is selected from the group consisting of antibiotics, local anesthetics, and combinations of two or more thereof. In one embodiment the therapeutic agent is rifampicin and/or bupivacaine.

1002 Ϊ Other embodiments of the invention are directed to:

a drug depot comprising the above nanocomposite;

a wound dressing comprising the above nanocomposite;

a tissue scaffold comprising the above nanocomposite; and

a cardiovascular stent comprising the above nanocomposite.

In one embodiment of the cardiovascular stent, the one or more therapeutic agents of the nanocomposite are selected from the group consisting of drugs which control restenosis. In a specific embodiment the cardiovascular stent is adapted to provide controlled release of the drugs which control restenosis. In one embodiment the therapeutic agents are selected from the group consisting of everolinms, sirolimus frapamycin), otaroiimus, and paclitaxei.

(O022J Another embodiment of the invention is directed to a method of treating a wound comprising applying to the wound a eopolymer-xerogel nanocomposite, comprising a biodegradable, biocompatible copolymer, silica nan.oparticl.es and one or more therapeutic agents. In a further embodiment, the method is adapted to provide controlled release of the therapeutic agent(s). In one embodiment of the method of treatin a wound, the therapeutic agent is selected from the group consistin of antibiotics, local anesthetics, and combinations of two or more thereof In a specific embodiment said therapeutic agent is rifampicin and/or bupivacaine. [0023} Another embodiment of the invention is directed to a hollo tube nerve guide comprising the above nanocomposite. In one embodiment, the one or more therapeutic agents of the nanocomposite are selected from the group consisting of neurotrophic factors. m one embodiment the hollow tube nerve guide is adapted to provide controlled release of the neurotrophic factors.

[002 1 The present invention can be understood, more readily by reference to the following detailed description taken in connection with the accompanying figures and examples, which form a part of this disclosure, ft is to be understood that this invention is not limited to the specific products, methods, eotidi lions or parameters described and/or shown herein, and that the terminology used herein is for the purpose of describing particular embodiments by way of example only and is not intended to be limiting of the claimed invention.

BRIEF DESCRIPTION OF THE .DRAWINGS

[0025} Figure 1 shows the chemical structure of tyrosine-? EG -derived poly(ether carbonate). Adjustable parameters are; x, the mole fraction, of the desaminotyrosyl tyrosine- derived monomer (DTK) (x is fixed at 90% in one embodiment) and y, the mole fraction of PEG (y is fixed at 10% in one embodiment); , the pendent alkyl chain length (i.e. the monomer is "DTE" when the pendent group is ethyl, or "DTO" when it is oct l); and the PEG molecular weight is fixed at 1 ,000 Daiions (degree of po!ymenzation^:::23) in one embodiment.

[0026} Figure 2 shows a nanocomposite film of the invention (QO01G/ 25) containing 25% silica xerogel before (A) and after (8) heating the film at 700^CC to bum off the poIy(DTO 10%.l^>EGcarbonate) copolymer matrix,

[0027} Figure 3 shows TEM images of poly(DTE- 10%PEG carbonate)

nanocoinposites at 3% (A) and 10% (B) silica xerogel loading,

[0028} Figure 4 shows the glass transition temperatures of composites as a function of silica xerogel particle sixe distributions. B001 Of Ik.) (first bar) s poiy(DT£- 10%PEG 1 k)- carbonate. Composites of HOC⁾ 10( 1 ) with micron-scale silica xerogel composites (gray bars) and E0010/N30 nanocomposite (black bar) are all at a fixed silica xerogel content of 30% (w/w). j 29J figure 5 displays the effect of silica xerogel content on the glass transition temperature, Tg, of nanocomposites. Copolymer, poly(DTE-1 %f EG I k) carbonate and nanocomposites, j0030| Figure 6 displays the eff ect of particle size in the composite formulations on the Young's modulus. The xerogel content was fixed at 30 wt%, Polymer is poly(DTE-

1 %PEGI k) carbonate (EOOI 0).

[0631 J Figure 7 shows the stress-elongation curves change with the fraction of sil ica in the E0010 nanocomposites. Low amounts of silica toughen the polymer (A), while higher loadings make it stronger hut more brittle (8).

[0632 j figure 8 displays_, the equilibrium, water content for nanocomposites are a function of silica xerogel concentration. Experimental EWC values (black, bars) where

calculated EWC values (open bars) are based on the assumption that only the .mass of copolymer _.present lakes up water.

}0033] Figure 9 displays the hy irolytic mass loss of nancomposites as a function of silica content; v ~ copolymer (OO l 0); A™ nanocomposite with 5% xerogel;□™ naiiocomposite with 25% xerogel; « ^« nanocomposite with 50% xerogel.

[0034) Figure .10 shows a graph of bupivacaine release from copolymer,

microcomposite and nanocomposite, versus time: o - - copolymer poly(DTO-iO%PBG l k) carbonate (O0010), 0 ~ microcomposite (O0010/MSO) and A- nanocomposite (O0010/N50). The microcomposite and nanocomposite contain the same silica loading, 50 wt%₅ and the bupivacaine content is fixed at 8 wi% for ail of the samples.

J003SJ Figure 1 1 shows a graph of the nanocomposite release of rifamplcin versus time. Rifarnpicin loading was 10% (wt:wt) in the poly(.DTO-10%PEG Ik) carbonate/10% silica (O0010/N 10) nanocomposite.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

[0036] The efficacy of prior controlled delivery devices for therapeutic agents is generally limited by the problem of so-called burst release .kinetics. The nanocomposites of the present invention rtxiuce or eliminate the burst release, instead. roviding continuous and constant rates of release of the therapeutic agent that are essential for sustained, effective therapeutic activity. The .nanocomposites uniquely combine materia! and drug delivery properties that are essential for wound dressings and various drug delivery applications. j 037| The inventive nanocomposites are biocompatible {i.e. substantiaUy non- cytotoxic and non-inflammatory), biodegradable, flexible, mechanically robust, and in particular are formabie into various devices which are capable of providing continuous controlled release of a wide array of therapeutic agents for a useful period, of time. Further, the robust, flexible and formabie nature of the nanocomposites enables their use as implanted, depots or wound dressings not only in hospitals and civilian uses but also in the far more demanding conditions of military uses such as on a battlefield or field hospital.

[0038 j The silica nanopartieles are preferably biodegradable silica-baaed glass nanoparticS.es. A preferred route of processing the particles is by a sol-gel methodology, although other methods can be used. The polymers are preferably high molecular weight, biocompatible, biodegradable amphophilic hydrogels comprised of, for example, polyi vinyl alcohol) (FV^'A); poiy(al.kylene oxides), including poly(ethylene oxide) (PEG, or PEG); polycaprolaetone (PCI.,); polymers of desaniinotyrosyl tyrosine; poly(lactic acid) (FLA), polygiyeolie acid (PGA), copolymers of lactic and glycolic acid f PLGA); polysaccharides; peptides; and linear, block or graft copolymers of these. High molecular weight of greater than, about 20,000 Daltons for the polymer is preferred for effective mechanical properties of the nanocomposites. Amphophilic polymer properties are conferred by the presence of one or more hydrophilic monomer units and one or more hydrophobic monomer units. Amphophilic properties enable sustained, controlled deliver of both hydrophilic and hydrophobic drugs.

{00391 Examples o desaminotyr syl tyrosine polymers include polycarbonates, poly- ary!ates, poiylminiocarbonates, polyethers, poiyuretlianes, po!ycarbamates, polythiocarboaates, polycarbonodithionates, poiyphosphoesters, polyphosphazmes and polythiocarbamales of this monomer family. Polycarbonates, specifically poJyfarnide carbonates), as well as polyurelhanes, polycarbamates, po!ythiocarbonates, polycarbonodithionates and poiythiocarbamat.es are prepared by the process disclosed by U.S. Patent No, 5, 198,507, the disclosure of which is incorporated by reference. Methods adaptable for use to prepare polyary late- polymers of the present invention are disclosed in U.S. Patent Nos. 5,317,077 and 5,658,995, the disclosures of which are incorporated herein by reference. Polyesters, specifically poly(ester amides), are prepared by the process disclosed by U.S. Patent No. 5,2.16,1 15, the disclosure of which is incorporated herein by reference.

(0040J Polyiminocarhonates are prepared b the process disclosed by US 4,980,449, the disclosure of which is incorporated by reference. Polyethers are prepared b the process disclosed b U.S. Patent No. 6,602,497, the disclosure of which is incorporated by reference, Polyphosphoesters and polyphosphazin.es are prepared by the process disclosed in U.S. Patent No, 5,912,225. the disclosure of which is incorporated by reference. The other phosphate polymers disclosed in U.S. Patent No. 5,912,225 are likewise suitable for use with the present invention. Desammotyrosyl tyrosine monomers are prepared according to the methods disclosed by U.S. Patent No. 5,099,060, the entire disclosure of which is incorporated, herein by reference.

(0041 Random block copolymers of these polymers with poly(alkylcne oxides) can be prepared as described in U.S. Patent No. 5,658,995, the disclosure of which is incorporated by reference. Radio-opaque versions of the foregoing polymers are prepared according to the methods disclosed by U.S. Patent No. 6,475,477, the entire disclosure of which is incorporated herein by reference.

[0042} The polymers and copolymers can be cross-linked, either by covatent or ionic bonding, to form the hydrogels or to otherwise promote critical performance properties including gelling, fluid adsorption and increased mechanical strength. Versions of these polymers with free pendant earboxylic acid groups available for cross-linking are prepared according to the methods disclosed by U.S. Patent No. 6, 120,491 , the entire disclosure of which is incorporated herein, by reference. Cross-linked versions of these polymers are prepared according to the methods disclosed by U.S. Patent No. 7,368,169, the entire disclosure of which is incorporated herein by reference.

[0043] The nanocomposites provide controllable binding and release of therapeutic agents, thereby providing controlled delivery of the therapeutic agents for healthcare applications. The polymers and the silica iianoparticles independently can contain therapeutic agents,, are independently capable of binding these agents, and can. i ndependently release such agents, it is sufficient that either the polymer or the nanoparticles contains therapeutic agents, although both can contain them. j0044f The nanocomposites of nanoparticles in various biocompatible, biodegradable polymers provides a unique matrix that enables better control of the kinetics of delivery of the therapeutic agents than can be attained by either the polymers or nanoparticles alone. In one embodiment, the nanoparticles are embedded in . polymers having the form of a film, which enables the use of the outstanding release properties of the nanoparticles in applications where a solid sheet is needed for treatment, such as in wound dressings,

10945} in another embodiment, the polymer-nanopartiele- nanocomposite is fabricated for use in the depot delivery of therapeutic agents such as organic drug compounds, genes, oligonucleotides, and proteins, and in wound treatment applications such as for compartment syndrome* chronic and phantom pain treatment_* hemostasis and infectio control.. Thus a wide variety of therapeutic agents including, without limitation, antibiotics, local anesthetics, analgesics, vasodialaiors, and vasoconstrictors can be so delivered. Furthermore, the

nanocomposites can be formulated for pseudo-first order release of one or more therapeutic agents therefrom.

[0046} Therefore, according to one embodiment, biodegradable, biocompatible silica- polymeric nanocomposites are provided that control delivery of local anesthetics and antibiotics directly to the wound site to provide pain relief and infection control. The biomaterial nanocomposite films provide sustained treatment of the peripheral nerves located at the wound site with a local anesthetic that functions as a sodium channel blocker to shut, down the tiring of the afferent axons that cam- the pain signals back to the brain. This- educes or eliminates the imprinting process in the central nervous system that is recognized as a key component of chronic pain. Similarly, sustained delivery at the wound site of ami microbial agents eliminates infections caused by pathogenic biofUms that might otherwise lead to osteomyelitis, non-healing of bone fractures and other serious complications. Further, delivery of pain and antimicrobial medication by the robust formable nanocomposites, beginning on the battlefield or in combat support hospitals or in surgical procedures at veterans and civilian hospitals, leads to reduced morbidity, decreased postoperative narcotic usage, and the attenuation of chronic pain and infection syndromes.

10047} in accordance with another aspect of the invention, there is provided a biocompatible nanocomposite designed to counteract the effects of compartment sy ndrome of the tissues. Thus the present invention provides nanocomposites of biocompatible polymers and bio- resorbable silica-based sol-gels that deliver antt-apoptotie and pro-angiogenie factors to seal damaged cell membranes and thereby repair damaged tissues. The nanocomposites also absorb extracellular fluid within the compartment to reduce hydrostatic pressure and minimize the extent of damaged tissue. These treatments can be used prophylactic-ally to reduce, if not eliminate the need for fasciotoroies. When required, the treatments can be used to accelerate healing after lasciotomies.

[0948} Another aspect of the invention, is directed to a cardiovascular stent comprising the inventive nanoeomposite. Preferably for stent applications, the therapeutic agents of the nanocomposi te are selected from the group consisting of drugs which control restenosis; These can include agents selected from the group consisting of everoiimus, sirolinius (rapamycm), zotarolimus, and paelitaxel. The cardiovascular stents of the invention are- adapted to provide controlled release of these drugs, which thereby eliminates, reduces, delays or otherwise controls the restenosis process. This control of restenosis can last from months to years, preferably 6 months to 5 years.

[0049} Another aspect of the invention is directed to a hollow tube nerve guide or conduit comprising the inventive nanocomposite. Preferably for nerve guide applications, the therapeutic agents of the .nanocomposite are selected from the group consisting of neurotrophic factors. The nerve guides of the invention are adapted to provide controlled release of the neurotrophic factors, thereby stimulating regeneration of nerve tissue.

Nmweomp ite imtrphelogy.

[00501 When tsn-modifted inorganic particles are mixed into a polymer matrix without the hel of a compatibilixer, the dispersed inorganic phase often tends to agglomerate, which results in opaque films with poor mechanical properties. The present invention avoids this problem by preparing nanocomposites in situ from fully miscible precursors. Tetraethyl orihosilicaie (TEOS) is hydroiyzed under acidic catalysis conditions and the condensation reaction is allowed to proceed only so far as to form a sol but not a gel state. jOOS!J The sol is then mixed vigorously with the copolymer dissolved in acetic acid such that a clear, homogenous solution is obtained. Upon solvent casting and drying this solution, the resulting films are. found to be optically transparent (Figure 2), which is indicative of dispersion of silica at sub-micron particle size. Afte burning off the copolymer at 700 °C, the residual silica maintains the original shape of the film sample, which is further indicative of the uniform dispersion of the silica throughout the copolymer matrix. The observed shrinking of the film is expected given that only 25% of the mass remains afterburning the copolymer.

{00521 That these films are indeed nanocomposites is confirmed, by TEM

micrographs of composites containing 3% silica xerogei (.E00I0 N3) and 10% silica xerogei (EO01O/N1 O) (Figure 3), In both samples the silica xerogei particles are well below 50 nm diameter and are uniformly distributed throughout, the polymer matrix with no micron-scale aggregates present. For E0 10 N3, polymer-rich domains (lighter regions) of approximately 50- 100 fifii in length can be observed, while silica-rich domains are also present (darker regions). As the s.tliea content is increased, the silica distribution remains fairly uniform and the silica xerogei network appears continuous,

Glass transition temperature.

100531 The glass transition temperature, Tg, is a measure of the motion of polymer chain segments and is dependent on chain rigidity, cohesi ve energy density, polarity, molecular weight and cross-linking between chains. Above the Tg the cooperative movement of a certain number of backbone units is allowed and the polymer chains can slide past each other when a force is applied. For the micron-scale xerogei composites containing 30% silica particles ranging in size from .10 to 105 μτη. the Tg^*s are all 38~39^w€, the same as that of the copolymer alone (Figure 4). T his; is indicative of minimal perturbation of polymer chain motions by the micron- scale silica particles and hence of weak kterfaciai interactions between the copolymers and silic particles. For the 30% silica-containing nanocomposite, however, the Tg is 8.5 °C, which is 46 °C higher than that of the copolymer or the micron-scale xerogei particle composites. This is indicative of a significant interfacial interaction between, the copolymer chains and the nana- scale silica particles that significantly restricts copolymer chain segment mobility,

[005 { The glass transition temperature of the nanocomposites i ncreases as the weight fraction of silica xerogei increases (Figure 5). The Tg values for these nanocomposites do not scale linearly with the silica particle surface areas; that is, assuming spherical silica nanoparticles of constant density and volume directly proportional to the weight percentage in the nanocom-posites, the calculated ratio of Tg/A, where A is the particle surface area, is not constant but rather decreases from 2.9 for the 5% xerogei nanocomposite to 1.5 for the 50% xerogei nanocomposite. [0055J While the. compositions and reaction conditions for forming the

nanocomposites preclude covalent. bonding between the copolymer chains and the silica, the observed Tg behavior with increasing xeroge! content is consistent with an increasing number of interfacial non-covaleni binding interactions including hydrogen bonding between silica-derived hydroxy 1 groups and the copolymer's PEG chain oxygen atoms and DTE amide group nitrogen atoms

Mechanical properties,

[00561 The Young's moduli of the nanoconiposites exceeded by factors of 5 to 20 times those of the copolymers or of composites made with micron-scale silica particles, increasing the fraction of xerogei in the nanocomposites increased, the glass transition temperature and the mechanical strength but decreased the equilibrium water content, which were all indicative of strong non-covalent mierfaciai interactions between the copolymers and the silica nanopar cles.

[0057J The Young's moduli for all of the micron-scale xerogei particle composites and the nanocomposites were significantly greater than that of the po!yCDTE- 10%

PEG ifc)carbo.nate copolymer alone and increased with decreasing xerogei particle size (Figure 6). At a fixed 30 wt% xerogei content, the composite with the smallest micron-scale particle size range ( 10-20 pm) had a modulus of 664 MPA compared to modulus of 430 MPa for the composite with the larger particle size range (70-105 pm). In contrast, the modulus of the nanocomposiie at this same 30 wt% xerogei concentration was greatly increased to 920 MPa. The Young's moduli for nanocomposites at silica xerogei concentrations between 0.5% and 5% were between 16 MPa and 384 MPa, about twice that of the po!yi DTE- 10%PEG 1 k.) carbonate copolymer alone, and increase rapidly to 920 MPa at 30% xerogei (Table I }.

[O058| The elongation at fracture, measured here as (L-Lo) Lo, where L is the final and Lo the initial sample lengths, for the nanocomposites was essentially the same as that of the copolymer. 1230%, for xerogei concentrations up to 3% but began to decrease above that concentration, diminishing to 5% at 30% xerogei, which is indicative of the brittle, inflexible nature of the nanocomposites at very high xerogei. loadings. Similarly, the Young's modulus of nanocomposites based on poly(DTO- 10%P£G t k carbonate) increased rapidly from 4.4 to 80 MPa when the silica loading was varied from 5 to 10 wt (Table I). Table .1. Tensile properties of the pol.y(BTE~10%PEG1.k carbonate) and poJy(DTO« 10%PEGlk carbonate) nanocomposites de end on the silica xerogel concentrations.

00591 The tensile behavior of the nanoconiposites changed gradually as the silica content increased. When small fractions of silica were present in the oanocomposites they became stronger but remained just as ductile as the copolymer alone (Table I and Figure 7, top). From 0.5 to 5 wt% of silica loading, the stress at yield stayed almost constant at around 22 MPa, about 50% higher than that of the copol mer.

[OOoiij When the nauo-silica .fraction was increased {E00.I0 M25 in. Table 1 and Figure 7, bottom) the nanoconipostte became much stronger as evidenced by the three-fold increase of the stress at yield compared to the copolymer (48 MPa. and 14 MPa, respectively). However, the polymer chain movement under stress was restricted by the silica and the nanocomposite was much less ductile (only 12% elongation at fracture). As the silica loading reached 30 wt%, the nanocomposif.es became even less ductile and did not yield (B00I0/N30 in Figure 7„ bottom).

Equilibrium witter uptake

fOOoi] As the silica loading in the nancomposite increased, the equilibrium water content (BWC) decreased (Figure 8). The copolymer itself was a weakly absorbent hydrogel with an EWC of 18%. If the silica in the nanocomposites were inert and did. not take up water or interact with the copolymer chains, then the amount of water uptake would be directly proportional to the mass fraction of copolymer present. [0O62[ For example, for the 25% silica xerogel-containing nanoeoraposiie, the mass fraction of copolymer in the nanocomposite was 75% and that mass of copolymer would by itself have an EWC of 13.5% (Figure 8), However, the observed BWC for thai nanocomposite was only 5,9%, which further demonstrated a significant mierfaciai interaction between the silica and the copolymer such that copolymer chain mobility was greatly restricted and water uptake as thereby reduced. For the microeon posites, the EWC was also Jbund to decrease as the silica concentration increased, a trend that was seen with many but not ail composites and which depended upon the nature of the polymers nd inorganic components, their particle volume fraction and any non-cova^'lent or c valent bonding between the components.

Hydrotytie Degradation ami Er km,

I0063J The mass loss of the pory(DTE-l 0¾ΡΕ<31 k)carbonate copoly mer by itself during incubation in buffer solution at 37 C was negligible for 6 days and then increased slowly over several weeks (Figure 9). The mass loss for the nanocomposit.es during incubation was significantly greater than, that of the copolymer alone at every time interval and increased with increasing silica xerogei concentration. The kinetics of nanocomposite degradation after the initial 24 hr period were essentially linear with time and increased only slightly as the silica weight fraction increased.

[00641 Nanocomposite degradation was faster than for i e copolymer alone, which was ascribed to the rapid dissolution of nano-scale silica particles. There was a significant mass loss in the first 24 hr for the nanocomposites of up to 8% for the 50% silica-containing specimen and, since the copolymer itself did not significantly degrade in that time frame, this mass loss of the nanocomposit.es was attributed to the rapid dissol ution of the silic nanopatticies adsorbed on or near the outer surfaces of the specimens. The water uptake and degradation rate of the composites can be increased by increasing the hydrophilic PEG content of the copolymers and can be decreased by substitution of the more hydrophobic DTO mononier for the DTE monomer.

Drug release fro n n compos es

[00651 The release kinetics of hnpivaeaine from the microcomposites and nanoeom- posites were compared to the poly(DTO- IO%PEGI k)ca.rbonate copolymer alone at a fixed loading of 8 wt% bupivacaine. The copolymer alone exhibited an initial very large burst release stage of 50% of the drug load in the 24 hr followed by a second stage of relatively constant slow release (hereafter (Figure 10), The nanocomposite also exhibited 2-stage behavior but had a

- J 6 - much reduced ini tial stage release of only 10% of the drug in the first 24 hours- and thereafter continued to release the drug at a rate comparable to that of the copolymer.

{ 0 61 The relatively small difference in the hydrolytic degradation of the copolymer and the nanocoraposite over the first 24 he cannot explain the significant reduction in the initial release from the nanoeomposite. Rather, this difference was ascribed to differences in water influx and drug efflux from the samples thai were determined by the different physical states of the co-polymer chains.

[0067} When no silica particles are present, the drug-loaded copolymer is in a rubbery state (the Tg is 2°C). In contrast, the nanocomposites are in a glassy state as evidenced by the higher Tg (59°C). Therefore the copolymer chain mobilities in. the nanocomposites are more restricted and water uptake is reduced, which slows drug solubilization and diffusion out of the nanocomposite compared to the pristine copolymer. The silica nanoparticl.es appear also to impede efflux from the nanocomposites by binding the drugs and/or by acting as physical barriers to flow.

[0068} In contrast, in the microeomposites the drug is initially confined entirely within naiiopores of the xerogel particles and the copolymer matrix acts as a barrier membrane to further control water influx and drug efflux. The porosity of the micron-scale xerogel particles and the hydrophobieil of the copolymer matrix determine the drug release kinetics of the microeomposites. which for the O00I0/M50 is fester than for the nanocomposite and essentially zero-order, i.e., pseudo-zero order over the first 72 hr. (Figure 10)

{00691 "Pseudo-zero order" release is a well-known term of art referring to a kinetic drug release profile equivalent to essentially zero order release obtained by balancing diffusiona! slow-down and acceleration of the release rate by erosion. For purposes of the present invention, "essentially zero-order release" and "near zero-order release" refer to a drug release rate at or near zero order over the sustained release phase of drug delivery under physio logical, conditions. Compositions with drug release at or near zero order have drug release coefficients thai are essentially unchanged relative to the arithmetic mean over the sustained release phase of drug delivery under physiological conditions,

{08701 For example, In one embodiment "essentially zero-order .release" and "near zero-order release" refer to the release kinetics of polymer compositions under physiological conditions, in which the release rate of drug from the composition varies by no more than ± 10% over the sustained release phase following the Initial burst for a period of about 1 week to about 4 years. One embodiment had a sustained release for a period between about one month to about three years. Additional, embodiments included compositions in which the release rate of drug from the composition varied by no more than ±9%, ±7,5%, or ±5% over the sustained release phase following the initial burst,

(0071 J As guided by the present specification, one of skill in the art can manipulate the release profile by adjusting certain features of the composition, for example, the polymer(s), drug(s), level of drug loading, surface area, etc. Furthermore, the initial burst can be shortened to less than one week by subsequent processing such as rinsing the blend to remove drug at or -near the surface or by coating the composition with a bioerodib!e polymer that is either drug free ot has a reduced drug content.

[0072} The release rate of the antibiotic, rifarapicin, from the nanoeomposite (Figure 1 1) was similar to that of bupivacaine from the nanocomposite. The initial loading of rifampicin wa 10% wt:wt in the po!y(DT010% PEGifc) carbonate (Ο0010 Ϊ.0) nanocomposite. The initial rifampicin release over the first 24 hr was about 10% of the rifampicin loading and this was followed by a slower second stage release rate.

[0073J This is consistent with the similar physical properties of the two drugs:

rifampicin is hydrophobic, with an octanol/water partition coefficient of log -2,72 and water solubility of 1.4 mg ml; similarly, for bupivacaine, the log P is 3.41 and the water solubility is 2.4 mg ml. When the cumulative rifampicin release data are plotted as a function of t^!" they can be fit by a single straight line (correlation coefficient of 0.98) which is consistent with the Higuchi model for diffusion controlled drug release.

(00741 Thus, optical iy clear nanoeomposites of tyrosine-PEG-derived polyether carbonate copolymers and silica xeroge!s were obtained by the simple process of raising the silica sol into the copolymer solution. Based on TEM micrographs, the particle size of the silica xerogels was about 5 to about 50 nm homogeneously distributed throughout the copolymer matrix. Material properties of the nanoeomposites depended on the amount of the silica xerogel present. Increasing amounts of nano-sized xerogel increased the Tg and the mechanical strength, but decreased the equilibrium water content (EWC). The increase in the Tg, and the fact that the experimental EWC's for various compositions were significantly lower than the theoretical values, is indicative of a strong interfacial interaction between the copolymer and the silica nanoparticl.es. j 075| The hydrogen bonding interfacial interactions between the large number ethylene oxide units in the copolymer backbone and the silica nanoparticles can act as physical cross-linkers and explain the reduced polymer chain mobility reflected by the increased Tg. The Tg behavior of the present nanocomposites contrasts with, similarly prepared nanocomposites based upon poly{K-eaproIactone) and TEOS-derived silica, where no significant increase in Tg is observed with increased silica content in the nanocomposites. The difference between the poiyfg- caprolactone) nanocomposites and poly(DTE~ 10%PEG I k) carbonate nanocomposites can be ascribed to the large number of PEG oxygen atoms present in po.ly(! T ~ 1 ()%PHG I k) carbonate copolymers compared to the poiy(e-capfolactone), which provides only a very limited number of ester group oxygen atoms for hydrogen bonding to the silica-derived hydroxy! groups, and hence there is no significant increase in interfacial. hydrogen bondin as the silica nanoparticie content is increased in the poly(e~caprolactone) nanocomposites.

[0076} Sustained controlled delivery of two clinically important drugs, ri ampicin and bupivacaine, was obtained from the drug-loaded nanocomposites. After a small initial burst release that could, be attributed, to the dissolution of the loosely bound drug, bupivacaine was released at a constant rate for a total of ? days. The amount of rifampicin released from the nanocomposite in the first 24 hr was 0.06 mg ml, which exceeds the minimum inhibitory concentration (MIC) for planktonic mewicUlm-resistam Staph, aureus (MRS A) infections and for Staph, epidermidis biofilms. The release of rifampicin from the nanocomposite wa i accord with the Hugichi model (i.e., follows linear behavior when plotted as t ^; ) so higher drug loading levels would be expected to result in greater solution concentrations such that treatment of biofilms in vivo would be possible.

[0077| .By varying the silica nanoparticie loading and the copolymer matrix compositions it. has now been demonstrated that nanocomposites of silica xerogels and tyrosine- po!yC ethylene glycol )-derived po!y(ether carbonates) provide a broad, tunable range of mechanical properties and bio.-degradabil.ity under physiological conditions. The strong tensile properties of the nanocomposites. and their controlled release of hydrophobic drugs make these biomateria!s highly attractive for applications such implantable drug delivery depots and wound dressings for treating ai ), and orthopedic infection, for tissue engineering substrates, for cardiovascular stents and for nerve guides or conduits.

{00781 The polymers copolymers and silica nanoparticies independently can contain therapeutic agents, are independently capable of binding these agents, and can independently release such agents, it is sufficient that either the polymer or the nanoparticies contains

therapeutic agents, although both can. contain them. The nanocoinposite of the nanoparticks in the polymer pro vides a unique matrix that enables far better control of the kinetics of delivery of the therapeutic agents than can be attained by either the polymers or the nanoparticies alone. These nanoeomposUes provide uniq u control of binding and release of therapeutic agents, thereby providing controlled delivery of the therapeutic agents for healthcare applications. The nanoeomposUes carabine the advantages of the drug binding and release kinetics of silica sol- gels with the mechanical flexibility and drug binding of polycarbonate- films, and further, are uniquely formable into various devices,

[007 | Th drug delivery system of the present invention permits fine tuning of drug loading and drug release kinetics while providing the mechanical strength and stability properties characteristic of heterogeneous nanoeompositea. The nanoeomposUes of the present invention are designed to reduce burst release and provide the continuous and constant rates of release of a therapeutic agent thai is essential for sustained, effective therapeutic activity. The release- of one or more therapeutic agents from the present nanoeomposUes can be pseudo first order release (le., the release kinetics of the present nanoconrposites can be characterized by a substantially constant release of therapeutic agent over time).

{00801 Conditions for synthesizing the silica nanopartkles can be controlled to produce a particular controlled release profile for a therapeutic agent corresponding to a concentration with known therapeutic effect. The drug molecules, incorporated in nano-sized pore channels of the nanopartkles and non-covaiently bound by the copolymers of the biocompatible film, will release by diffusion through the aqueous phase that penetrates into the nanocomposiles.

{00811 The parameters of the silica nanopartkle synthesis affects the fundamental properties of the particles that control release of the therapeutic agent. These parameters include specific surface area, granule or powder size, and pore size and porosity. Formation of

nanoeomposites of nanoparticies in polymers, such as in poIy(DTR-co-PBG carbonate), can be by compression molding; the copolymer compositions (pendent ester R chain lengths, PEG molee-ular weight and PEG/DTR molar ratios) can be varied systematically to achieve an optimum loading efficiency of the drug-loaded silic sol-gel nanoparticles .and to improve the mechanical properties of the films, such as tensile and flex strengths.

|l> 82| The .nanocoo posites of the present invention, are useful in depot delivery of therapeutic agents such as organic drug compounds, genes, oligonucleotides, and proteins, and in wound treatment applications such as for compartment syndrome, chronic and phantom pain treatment, heniostasis, and infection control. The nanocomposites of the present invention can be useful in various therapeutic applications, including treatment of pain resulting from wounds and prophylactic^* treatment of compartment syndrome associated, with wounds. For the treatment of pain, silica-based nanoparticles and tyrosine-based copolymers can be synthesized to effectively bind and release therapeutic agents such as bupfvaca e and raepivacaine.

(08831 For the prophylactic treatment of compartment syndrome, sol-gels and copolymers can be synthesized to effectively bind and release anti-apoptotic and pro-angiogenie factors. While the therapeutic nanocoraposiies of the present invention can be described in connection with a single drug, it will be understood by those skilled in the art that the therapeutic nanocomposites are capable of concurrent delivery of multiple drugs.

Pain Treatment

( 8841 Also provided is a. novel approac to the treatment of chro ic pain arising from wounds with severe tissue damage and/or from surgical procedures. This approach entails control led release of a selected local anesthetic from biocompatible, biodegradable

nanoeomposites applied directly to the wound site beginning as soon as possible after the wound or surgery occurs. The biocompatihe nanoeomposites provide sustained treatment of the peripheral nerves located at the wound site with a local anesthetic that functions as a sodium, channel blocker to shut down, the firing of the afferent axons that carry the pain signals back to the brain. This technology can potentially reduce or eliminate the imprinting process in the central nervous system that is recognized as a key component of chronic pain.

(008S1 In accordance with this aspect of the invention, a local anesthetic can be bound to a nanoconiposite matrix, comprised of silica nanoparticles incorporated in a tyrosine based polycarbon te- PEG film to provide controlled release of the anesthetic. The local anesthetic is preferably mepivieame or bupivicaine, because of their high activity with low cardiovascular side effects. The iianoeomposiies are preferably effective for up to 72 boors, permitting easy use on the battlefield, in combat support hospitals, and civilian and veterans' hospitals.

(O086J Bupivacaine and mepivacaine can be incorporated by addition of appropriate solutions to a mixture of the silica oanoparticles (silica sol) and the biodegradable, biocompatible copolymer, preferabl a desaminotyrosyl tyrosine ester-PEG carbonate copolymer. The immediate and sustained delivery of local anesthetic enables quicker recover times, shorter hospital stays, earlier achievement of phy sical therapy milestones, and lo wer rates of narcotic use and abuse among military and civilian patient populations.

Prophylactic Treatment Of Compartment Syndrome

[0 871 in -compartment syndromes, there is a zone of tissue that is between normal and irreversibly damaged, and in this zo e anti-apoptotlc and pro-attgiogenic factors can be useful to restore function. Thus, in accordance with one aspect of the invention, provided is a prophylactic treatment of a wound site to avoid the onset of compartment syndrome and associated fasctotomy treatment Even when faseiotomy is ultimately required, treatment in accordance with the invention provides for more rapid and complete healing of incision and wound sites.

[O08$| in. acute compartment syndrome, .fluid accumulates and the intramuscular pressure (IMP) increases. Removal of only about 1 ml of interstitial fluid can result in a reduction of intramuscular pressure such that intramuscular pressure (IMF) is restored to a normal range. Thus, in accordance with one aspect of the invention, nanocomposites made from polymers such as tyrosine-based block copolymers and silica nanoparticies can be designed and formed as a polymer-nanoparticle wound dressing to remove fluid from injured muscle compartments,

[0089! The biocompatible nanocomposites can be composed of ty rosine-based copolymers and silica sol-gels in the form of nanocomposite films or other shaped devices that are adapted to absorb 100% or more of their weight in body fluid while maintaining their flexibility, adhesion, and mechanical integrity. To provide this type of fluid adsorption, well-established synthetic polymer chemistry methods for forming cross-linked polymers can be employed. jOOWf Further, the nanocomposite dressing is capable of concurrently delivering a selected therapeutic agent to the wound site. The therapeutic agent can be incorporated in the resorbable nanocomposite of silica nanopattieSes and biodegradable, biocompatible copolymer. The therapeutic agent incorporated into the nanocomposite can include one or more of an anti- apoplotic factor, a pro-angi genie factor, and a polymeric surfactant

Tyrosine- ased Polycarbonate-Po!y(Etiiyleiie Glycol) Copolymers

[00911 Degradable polyesters, polyt^'glyeolic acid) (PGA), poly (lactic aeld) (FLA), their copolymers (PLGA), and poiydioxanone, are the predominant synthetic, degradable polymers with extensive regulatory approval histories in the USA. Although the utility of these materials as sutures and in a number of drug delivery applications is well established, these polymers cannot meet many of the material properties required for drug delivery devices. j0092] For example, all of these polyesters release acidic degradation products, limiting their utility to applications where acidity at the implant site is not a concern. They also tend to be .relatively rigid, inflexible materials, a disadvantage when, mechanical compliance with soft tissue or blood vessels is required, Finally, the chemical properties of these polyesters is not substantially tunable, being limited to_. nly a few combinations of fixed monomer structures, which limits thermodynamic and kinetic parameters that control drug binding and release.

[0093| The present invention encompasses a broad class of tunable, desaminoiyrosyl tyrosine ester (DTK) diphenolic monomers that can be used to prepare polycarbonates and other polymer families. Amon these polymers, tyros ine-dersved polycarbonates have been studied most extensively and have been found to be tissue-compatible, strong, tough, hydrophobic materials thai degrade slowly under physiological conditions. Further, it is preferable to use tyrosine-based block copolymers rather than polylactides because of the tar greirter tunability of the tyrosine-based blocks and because the polylactides are known to have inflammatory effects s vivo whereas the tyrosine-based copolymers do. not. When these .tyrosine-derived diphenolic monomers are eopolymerized with blocks of poly(ethylene glycol) (PEG), a class of poly(eth.er carbonate is is obtained that is elasiomeric with remarkable tensi le strengths and elongations. EXAMPLES

Materials and Methods

Materials

{0O9 | Teiraethoxysilane (TEOS) was purchased from Strein Chemicals, Newbury- port, MA. Pyridine 99% was purchased from Ae.ros (MorrisPlains.MI). Polyfethyiene glycol) of molecular weight 1 ,000 (PEG I K) and bis(trichioromet yl)carb4)nate were purchased from Fluka (Milwaukee, T), Methylene chloride HPLC grade and methanol HPLC grade were purchased from Fisher Scientific (Morris Pkras,NJ). Tetrahydrofuran (TUP) high, purity sol vent stabilized with 250 ppm BHT was purchased from BM.D (Gibbstown, NJ), A-propanot bupivacaine hydrochloride, rifamptcin, Dulbecco's phosphate buffer saline, acefonitrile HPLC grade and water solution containing 0.1% (v/v) trifluoroacetic- acid for HPLC were purchased from Sigma ASdrieh (Milwaukee, WT).

Methods

Copolymer synthesis and chftracteriz&tiott,

Synthmis ofPalyfDTR-eo-PEG Ciirhomte)

[0O95| These copolymers are referred to as poly(DTR~co-/P.EG M carbonate) where R represents the type of ester pendent chain, / represents the percent molar fraction of PEG units present within the backbone, and M represents the molecular weight of the PEG blocks. Thus, po!y(DTE-eo~ 5%PEG I (KK) carbonate) refers to a copolymer prepared from the ethyl ester of desam oiyTosyHyrosine containing 5 mo.l% of PEG blocks of a verage molecular weight of 1000 g/^'mol. This molecular design provides tunability through three independent variables to enable optimization of materials properties (i) the pendent chain (ii) overaii PEG content /; and (Hi) length (molecular weight) M of the PEG block.

There are an enormous numbe of possible structures with this molecular design, including copolymers of poly(DTO-co PEG 1000 carbonate), where/^;-^~; 0%, 10%, 40% and 70% to provide a range of hydrophobic-to-hydrophilic properties, DTO (i.e, the octyl ester) was selected because it has been identified as an ester having superior thermodynamic solubility parameter for binding hydrophobic drug molecules. Synthesis was performed by adding the DTO monomer and PEG to round bottom flasks containing methylene chloride and anhydrous pyridine. At room temperature, phosgene solution in toluene was added over 90 min to the reaction mixture with overhead stirring. Teirahydrofuran (THF) was then added to dilute the reaction mixture to a 5% (w/v) solution. j0097J The copolymer was precipitated by slowly adding the mixture into 10 volumes of ethyl ether. For further purification, copolymers with lower PEG content (<70% by weight) were redissoSved m THF (5% w/v) and repreeipitated by slowly adding the polymer solution into 10 volumes of water. Copolymers with higher PEG content: (70% by weight) were redissob/ed in THF (10% w v) and reprecipitated by slowly adding the polymer solution into 10 volumes of isopropanol in each case, the precipitated copolymer was collected and dried under vacuum,

[009 1 The molecular weight of the copolymers can. be controlled by the duration of the reaction and determined by gel permeation chromatography using THF as the solvent and using polystyrene standards. Chemical structure and. polymer purity can be monitored by FT-I , H-N R, and C-NMR. The glass transition temperatures (T_g), crystallinity, and melting points of each copolymer can be determined by differential scanning c Sorimetry (DSC") and the decomposition temperature obiained by thermognsvimetrie analysis (TGA), with heating rales for both DSC and TGA of 10"C/min using an. average sample size of 15 mg.

[0099} Polycarbonate copolymers of poly(etliylene glycol) (PEG) and desamlno- tyrosyi tyrosine esters (DTR) can be prepared by solution . phosgenation as illustrated in Figure 3. These copolymers have weight-average molecular weights up to about 200,000 and have symmetrical molecular weight distributions. To obtain structure-activity relationships, copolymers were prepared with either 5% PEG 1000 or 5% PEG2000 and different pendent ester chains (R™ E (ethyl), 8 (butyl), H (hexyl), and O (octyl)). Also, the effect of PEG content was determined by preparing a series of poly(DTE~co-PEG 1 00 carbonatej's with PEG content ranging from 1 mot% to 70 moi%. AH of these copolymers were soluble in common organic solvents and those with high PEG content (70 wt%) were also soluble in water. Increasin the length of the hydrophobic pendent R. chain lowers the glass transition temperature, T_¾, in a linear fashion. The copolymers were observed to be thermally stable up to abou 300°C.

jO!O ] The binding and release of organic drug compounds by the copolymers is a function, of the hydrophobic! iy of the drug molecules as well as the hydrophobieity of the copolymer. The relative affinity of the copolymers for a drug can. be predicted by their thermodynamic solubility parameters. }01.ΘΊ] Organosilanes such as tetraethyoxysilane (TEOS) or ietrarnethoxysilane (TMOS) were used as the precursor molecules for the synthesis of the silica sol-gels via hydrolysis and condensation reactions. The hydrolysis reaction, which can be either acid or base catalyzed, replaces alkoxide groups with hydroxy! groups, SiJoxane bonds (Si-O-Si) are formed during subsequent condensation. Alcohol and water are byproducts of the. condensation reaction and evaporate during drying. Theoretically, the overall reaction is as follows: n Si(OR) + 2n H₂0 ~» tt Si0₂ - 4n OH

Ho wever, .in reality, tire completion of the reaction and the chemical composition of the resulting product depend on the excess of water above the stoichiometric l-EO/Si ratio of 2. A number of other sol-gel processing parameters (such as pH of the sol, type and concentration of solvents, temperature, aging and drying schedules, etc.) can also affect the composition, structure, and properties of the resulting product.

{0102] The po!y(eth.er carbonate) copolymer used throughout this study was composed of desammoryrosyl tyrosine ethyl ester (DTE) monomer and polyCethylene glycol) (PEG) of molecular weight 1 ,000 Daltons (Fig. I ), which is referred to as poly(DTE-co- .10%PE<} .1 k carbonate) and abbreviated as E00I 0, Similarly, poly(DTO-10¾PEGI k carbonate) contained desaminotyrosyl tyrosine octyl ester (DTO) monomer and PEG and is abbreviated as O0010. The two copolymers were synthesized following a previously reported method and their structure is illustrated in Figure I . The copolymer composition was confirmed by Ή N R ( MSO-rf^'i), Variaa VNMRS 400M& spectrometer) and Fourier transform infrared

spectroscop (FTIR) (Avatar 380 spectrometer, Thermo Nicotct). The number average (M_n) and weight average (M¾) molecular weights of the copolymer were determined by gel permeation chromatography (GPC; Waters Corp, 5 15 HPLC pump, 717 autosampier, 410 Ri detector, and Empower 2 software) with 103 and 105 Angstrom gel columns (Polymer Laboratories/ Agilent, Santa Clara, CA) in series, with Τί-!F as mobile phase and a flow rate of i ml mm ¹. Calibration was based on polystyrene standards (Polymer Laboratories/ Agilent),

Xerogel mknmsc k particle synthesis,

Comparison method.

(^{11031 Silica xeroge!s were prepared at room temperature via a one-step acid catalyzed sol-gel process using tetraethoxysilane (TEOS) as a precursor at a watenTEOS molar ratio of 10: 1 , Briefly, 1 , M HC1. was added to a mixture of water and TEOS at a 10: 1 molar ratio to a final pi! of 2,2. A clear sol formed after vigorous stirring for 20 mm. The sol was cast into cylindrical polystyrene vials that were sealed and the so! allowed to gel and ate at 3TC for 2 days. Subsequently the vials were opened and the gels were allowed to dry in an oven at 37°C for 3-4 days until the gel weight became constant. The silica gel was dried, crushed into granules, sieved using nylon meshes and sorted according to their particle size in 10-20 μπι, 20- 40 μηι, 40-70 pro and 70-1 5μπ diameters. Gas fN¾ adsorption/BET analysis {Autosorb- L Quamachrome, Boynton Beach, FL) was used to determine the surface area, pore size and pore volume of .xerogels that were first dried and outgassed at 50°C for 20 h. Bupivacaine-coniaimng xerogel microparticles were prepared b dissolving the drug in methanol and adding this directly to the acid catalyzed sol.

Micron-Mete particle composites fabrication.

Comparison method,

(0104] Mierocomposites of the copolymer and micron-scale xerogel particles, referred to herein simply as microcomposites", were prepared via solution blending method. For a typical 500 mg sample o mkrocomposite, 350 mg EQ01.0 copolymer was dissolved in ? m.L THF and 150 mg dry xerogel with the desired particle size was vigourosly mixed in for 2 minutes. The slurry was then poured into a PTFE mold and the solvent was slowly evaporated over 48 h in the fume hood to yield a uniform film. The resulting film was dried under nitrogen flow for 24 h arid in a vacuum oven at 50°C for 24 h. The micron-scale silica particle composites were abbreviated as, e.g., EOOIO/X 30(10), meaning a matrix of the copolymer EGO 10 containing 30% (wt:wt) silica xerogel (X) having a particle size of 10-20 μητκ

Nanocomposke synt sis and morphology.

(ΘΙ05) The nanocomposites were prepared in situ by adding deio.nk.ed water to T.BOS in a 20 raL scintillation vial to obtain a watenTEOS molar ratio, Ms, of 6-1 . The TEOS hydrolysis reaction was catalyzed by adding 1 N HCl to a final concentration of 0.35 M HO. The reaction mixture was stirred at room temperature for about 16 l.rrs to allow complete TEOS hydrolysis without allowing the silica polycondensation reaction to reach the gel point. Volumes of silica sol were transferred into small vials containing 10% solutions of poly(DTE-l 0%PEG1 k carbonate) in glacial acetic acid. The silica sol volumes transferred and the copolymer amounts used were chosen such that the theoretical amount of SiC¾ formed after hydrolysis corresponded to 0.5, 1 , 3, 5, 10, 25, 30 and 50 wi% SiGyoopolyraer In the final nanocomposites, |θί< | When the silica sols were added to the copolymer solutions, all of the samples remained transparent with no macroscopic phase separation or precipitation observed. The nanoeomposite solutions were then stirred for 5 minutes, poured into Teflon Petri dishes and dried under nitrogen .flow overnight, and then placed in a vacuum oven at 40°C for a total of 96 h. The nano-seale silica composites were abbreviated as, e.g., EO010 N3O, meaning a nanocomposite (N) of E0010 copolymer with 30% (wt:wt) silica erogel.

|Θ!Θ7{ Transmission electron microscopy (TBM) of the nanocoraposite films was performed by embedding them in a low viscosity epoxy resin and then cutting 50 nm thick samples using an ultamicrotome equipped with diamond knife. The thin sections were transferred to carbon-coated copper grids (200-mesh) and imaged, in a JEOL lOOC transmission electron microscope operated at accelerating voltage of 100 kV. No heavy metal staining of sections prior to imaging was necessary.

.Drug-eontuning nanoeomposites,

|01 8| Drug-loaded nanocomposhes were prepared b stirring the pre-hydroiize TEOS solution with the copolymer solution for 1 minute and adding appropriate volumes of a. 2 mg/mL bupivacaine or .rifampicm solution in methanol followed by vigorous mixing for another minute. The mixture was then poured into Teflon dishes, dried first under nitrogen fl w and then in a vacuum oven for 96 h at room temperature to avoid any possible drug degradation.

Thermal properties,

jfll0S>] Thermogravimetrie (TGA) experiments were performed in air and the temperature was ramped from 25 to 600 deg C at a 10 deg/min rate. The glass transition temperature (Tg) was determined by differential scanning calorimetry (2 10 Modulated DSC, TA Instruments) on 10- 15 mg samples. Specimens were sealed in aluminum pans and subjected to a heat-cool-reheat temperature program from -50 to 150^aC at a heating rate of lO /mm. The glass transition temperature were taken as the inflection points in the second heating scans of the DSC temperature program.

T ensile, properties,

jOHOJ Tensile properties of thin copoly mer and composite film, strips (30 x 5 x 0.20 mm) were tested according to AST standard D882-91 on a Sintech 5 D tensile tester. Measurements were done in dry state at room temperature, and also in water at 37°C after the samples were pre-ineubated in PBS at the same temperature tor 1 , 3 and 5 days respectively. The results for Young's modulus, elongation (strain) and tensile strength were averaged over 3-4 replicates. The width (5mm) and thickness of each of the specimens tested were averaged over three measurements in different parts of each specimen. The initial grip speed was 2rara/min, allowing for a reliable measurement of the elastic modulus; the yield point was calculated based on the zero slope criterion.

Equilibrium water uptake

θ I 1 J I Rectangular 200 pm-thick samples of polymer and composites were immersed in PBS at 37°C without shaking. The average sample size was 60 mg. At given time points the specimens were removed from the buffer, dried on a paper towel and weighted. The equilibrium water uptake (EWC) was caicu.laied as the ratio of the mass gain during incubation and the weight of the dry specimen:

and was taken as the point at which two consecutive time point samples had the same mass. Samples were run in triplicate.

In vitro degradation

| 112) For the accelerated hydrolyiic degradation study, rectangular 200 pm- thiek. 60 mg copolymer and nanocomposite samples were incubated in 20 mL PBS, pH 7.4 at 37°C and mixed at 100 rpm. The buffer was replaced every two days to maintain sink .conditions for the degradation products and the mass loss of these samples was followed for 90 day s. For each of the nanocomposite formulations (0,5, 1, 3, 5, 1 , 25, 30 and 50% by weight, silica loading) and the copolymer alone, the samples were removed from the buffer at selected time intervals, rinsed with DI water, freeze-dried and weighted. The mass loss was calculated as the ratio between the mas loss during the incubation period and the weight of the sample before incubation.

In vitro drug release and antimicrobial activity

|0i l3] The release rates of the local anesthetic, bupivacaine (BP), from the micron-scale particle composite and the nanocomposite films were measured for up to seven days using 30 mg of samples of the composites incubated in 6 mL PBS at 37^&C in a J labo SW2 water bath shaker at 100 rpm. The incubation medium was completely withdrawn at specified time intervals and replaced with 6 mL fresh buffer. The withdrawn samples were diluted 1 :1. (v/v) with aeetonitriie and analyzed by high performance liquid chromatography (HPLC). All experiments were performed in triplicate. The same procedure was used for the determining the release rate of bupivacaine and rifampicin from a nanocomposite film. Validity of the method was established through a study of specificity, linearity and accuracy according to the

international Conference on Harmonization fiCFI) guidelines. fOI 14] The antimicrobial activity of rifampicin and rifampicin-containing nanocomposttes were determined against Staphylococcus aureus UA S-I (ATCC 49230), a clinical osteomyelitis strain, using a slightly modified Kifby-Bauer zone of inhibition (ZOI) method. Frozen S. aureus UA S- I stock was thawed, and diluted in 4 mi Mueller-Hinton II broth (cation-adjusted) (MR^'BU) to a density of I McFariand unit (-0,25 AU), then used to streak a lawn of bacteria onto Mueller-Hinton agar plates. Three circular 6 mm diameter discs of each nanoeoraposite were placed on the agar plates equidistant from each other and midway between the center and edge of the plate. The plates were incubated overnight at 37 and the circular ZOI (absence of bacterial colonization, which was readily distinguished by visual inspection) measured with an electronic caliper.

[ ilSj Although the invention herein has been described with reference to particular embodiments, it is to be understood that these embodiments are merely illustrative of the principles and applications of the present invention, and. are not intended to limit the

invention in any way. It is therefore to be understood that numerous modifications can be made to the illustrative embodiments provided herein, and that other embodiments can be devised without departing from the spirit, and scope of the present invention as defined by the following claims.

|01 16| Ail references cited herein are incorporated by reference in their entireties.

Claims

What is claimed is:

I . A copolymer-xerogef nanocomposite, comprising a biodegradable, biocompatible copolymer, silka nanopariicles and one or more therapeutic agents.

2 , The nanocomposite of claim I , where in the biodegradable , b iocompatible copolymer has a molecular weight greater than 20,000 Daltons,

3. The nanocomposite of claim 1 or 2, wherein said therapeutic agent comprises a compound selected from the group consisting of antibiotics, focal anesthetics, and combinations of two or more thereof.

4. The nanocomposite of any of claims I to 3, wherein, said therapeutic agent is rifamp cm and/or bupivaeaine.

5. The nanocomposite of any of claims ! to 4, wherein said, biodegradable

copoiymer comprises a copolymer of tyrosine-poly(alky lene glycol )-derived poly( ether carbonate),

6. The nanocomposite of claim 5, wherein said biodegradable copolymer comprises a polycarbonate comprising desamtnotyrosyl tyrosine ester and polytethylene glycol).

?, The nanocomposite of any of clai ms I. to 6, adapted to provide control led release of said therapeutic agent(s),

8. A method of orming a therapeutic agent-loaded copofymer-xerogei

nanocomposite, comprising:

(a) providing a silica sol;

(b) adding said silica so! to a poly(desaminotyrosy] tyrosine esier~eo-PEG carbonate) to form a mixture;

(c) adding one or more therapeutic agents to said mixture to form a drug mixture; and

(d) removing the solvents from said drug mixture to form the therapeutic agent-loaded copoiyroer-xerogel nanocomposite.

9. The method of claim 8, wherein, said, therapeutic agent is selected from ihe group consisting of antibiotics, local anesthetics, and combinations of two or more thereof.

10. The method of claim 8 or 9, wherein said therapeutic agent is rifampkin and/or bupivaeaine,

1 1. A drug depo t comprising the nanocomposite of any of claims 1. to 7.

1 2. A wound dressing comprising the nanocomposite of any of claims I to 7.

1 3. A cardiovascular stent comprising the nanocomposite of an of claims I to 7,

14. A tissue scaffold comprising the nanocomposite of an of claims I to 7.

i 5. A method of treating a wound comprising applying to the wound a copolymer- xerogel nanocomposite comprising a biodegradable, biocompatible copolymer, silica

nanoparticles and one or more therapeutic agents.

16. The method of claim \ 5 adapted to provide controlled release of said therapeutic- agent.

1 7. The method of claim 15 or 16. wherein said therapeutic agent is selected from the group consisting of antibiotics, local anesthetics, and combinations of two or more thereof

18. The method of any of claims 15 to 17, wherein said therapeutic agent is riiampicin and/or buptvaeame,

19. The cardiovascular stent of clai m 13, wherein the one or more therapeutic agents of the nanocomposite are selected from the group consisting of drugs which control restenosis.

20. The cardiovascular stent of claim. 19, wherein said one or more therapeutic agents is selected from the group consisting of evero!lmits, sirohmus (rapamyein), zotarolirous, and paclttaxe!.

21 . The cardiovascular stent of claim 19 or 20, adapted to provide controlled release of said drugs which control restenosis.

22. A hollow tube nerve guide comprising the nanocomposite of any of claims 1 to 7.

23. The hollow tube nerve guide of claim 22, wherein the one or more therapeutic agents of the nanocomposite are selected from the group consisting of neurotrophic factors.

24. The hollow tube nerve guide of -claim 22 or 23. adapted to provide controlled release of said neurotrophic factors.