Patent Application
20130209522
The present invention relates to the fields of medical devices, biology and medicine, and more particularly to the field of implantable biomaterials and combination medical devices.
Each year in the United States alone, over 500,000 orthopedic surgeries are performed (Bostrom and Seigerman, Hss J, vol. 1, pp. 9-18, 2005; Giannoudis et al., Injury 36, suppl. 3, ppS20-7, 2005), many of which require the use of a natural or engineered bone graft to fill a traumatic or surgically-induced wound or defect (Early Radiological Diagnosis and Differential Diagnosis of Infection in Orthopaedic Surgery,” in Infection and Local Treatment in Orthopedic Surgery, E. Meani, C. Romanò, L. Crosby, G. Hofmann, and G. Calonego, Eds.: Springer, 2007). Cadaveric-sourced allograft bone (e.g., cancellous allograft fragment and morsellized, micron-sized particulate matter) is often used due to not only its high surface area, that provides an appropriate cellular environment to enhance tissue integration and bone remodeling (Nandi et al., Indian J. Med. Res. 132: 15-30, 2010; Kundu et al., J. Mater. Sci. Mater. Med 21: 2955-2969, 2010) but also its wound packing efficiency that minimizes the occurrence of avascular spaces susceptible to opportunistic bacterial colonization (Nandi et al., supra; Aronin et al., Biomaterials 31: 6417-6424; Kanellakopoulou and Giamarellos-Bourboulis, Drugs, vol. 59, pp. 1223-32, 2000; and Winkler et al., J Bone Joint Surg Br, vol. 90, pp. 1580-4, 2008).
Synthetic bone fillers (e.g., calcium phosphate granules, calcium sulfate-based granules, wafers, pastes, and polymers), or naturally derived bone replacement materials such as ProOsteon 500R (BioMet), a porous hybrid calcium carbonate/calcium phosphate coralline ceramic bone graft (Parikh, S. N., J Postgrad Med, 48 (2002) 142-148), and bioactive bone-based technologies (e.g., osteo-inductive growth factors, drug carriers) provide new surgical options with novel, ‘ala carte’ orthopedic solutions for trauma, revision surgeries and major repairs (Giannoudis et al., supra; McLaren, A. C., Clin Orthop Relat Res, (2004) 101-106).
Regardless of the type of implant, clinical success of bone graft void fillers relies on their ability to properly pack the orthopedic defect and allow adequate vascularization for graft integration via tissue, primarily bone, regeneration. Importantly, the intrinsic low vascularity of bone and persisting presence of susceptible avascular spaces provides a favorable niche for acute and chronic bacterial infection. Over 40,000 infected orthopedic surgeries occur per year, many of which are related to bacterial biofilm (see
The high revision infection rates are particularly troublesome. For hip and knee replacement infections, 1-3% of primary joint replacements become infected, but 8-15% of revision arthroplasty surgeries become infected. For knee revisions, over 54,000 primary surgeries occur per year, with a 9% revision rate (compare “TKA” lines of
These infectious events, particularly those that lead to biofilm formation, can further inhibit graft revascularization and proper cortical blood supply, leading not only to tissue necrosis (sequestra) but also to additional avascular spaces (Costerton, J. W., Rev Infect Dis, 6 (1984) 608-616). The porosity and resulting high surface area enables cancellous allograft bone fragments or morselized allograft bone as well as their synthetic surrogates (McKee et al., J. Orthop Trauma, 16 (2002) 622-627; Koort et al., Acta Orthop, 79 (2008) 295-301; and Koort et al., J. Biomed Mater Res A, 78 (2006) 532-540) to be exploited clinically both as a suitable bone substitute and filler and importantly as a local drug delivery vehicle to prevent or treat osteomyelitis, with the degree of porosity directly correlating to antibiotic loading efficiency (Nandi et al., supra; Aronin et al., Biomaterials 31: 6417-6424; Kanellakopoulou and Giamarellos-Bourboulis, Drugs, vol. 59, pp. 1223-32, 2000; and Winkler et al., J Bone Joint Surg Br, vol. 90, pp. 1580-4, 2008). Cancellous autogenic bone grafts and synthetic bone grafts impregrated with antibiotic prior to implantation have been reported to show a reduction in infection with no clinical contraindications, further indication that endowing clinically familiar cancellous allograft bone tissue with application-tailored, polymer-controlled antibiotic release may provide a delivery vehicle to effectively treat osteomyelitis (see, e.g., Borkhuu et al., Spine (Phila Pa 1976), 33 (2008) 2300-2304; Buttaro et al., Acta Orthop Scand, 74 (2003) 505-513; Finley, J. M., J West Soc Periodontol Periodontal Abstr, 49 (2001) 5-9; Winkler et al., Cell Tissue Bank, 7 (2006) 319-323; Witso et al., Acta Orthop Scand, 70 (1999) 298-304; Ketonis et al., Clin Orthop Relat Res, 468 (2010) 2113-2121).
However, simple antibiotic adsorption often used with bone graft materials produces rapid bolus release and limited therapeutic duration of up to a few days maximum (Jiranek et al., J Bone Joint Surg Am, 88 (2006) 2487-2500; Ketonis, et al., Tissue Eng Part A, 16 (2010) 2041-2049; M. Diefenbeck et al., Injury, 37 Suppl 2 (2006) S95-104; Ciampolini and Harding, Postgrad Med J, 76 (2000) 479-483; Levin, P. D., J Bone Joint Surg Br, 57 (1975) 234-237; and Miclau et al., J Orthop Res, 11 (1993) 627-632). This rapid bolus drug release often cannot stem long-term infections where opportunistic pathogens reside in the wound site several weeks to months.
As shown in
Currently used treatments fill bone defects with antibiotic-containing cement. However, this short release duration of the antibiotic is inadequate to resolve infection. Moreover, the current rapid bolus release approaches have resulted in acute local tissue toxicity and development of drug-resistant and multi-drug resistant organisms (
Thus, there is a need to find improved compositions and methods for treating orthopedic injuries (e.g., bone and joint injuries, and implant wounds) to produce better healing of tissues and bone implant sites while reducing the risks of long-term infection.
The invention provides compositions and methods for treating bone and joint injuries while reducing the risks of long-term infection.
Accordingly, in a first aspect, the invention provides an implant comprising, consisting, or consisting essentially of a uniform mixture of degradable polymer component, a bone component, and a drug component. In some embodiments, the drug component comprises, consists, or consists essentially of an antibiotic.
In some embodiments, the implant is configured so that upon implantation of the implant into a host at an implantation site, the drug diffuses from the implant at a therapeutic level (e.g., a level that will inhibit or prevent infection at the implantation site). In some embodiments, the host is a vertebrate animal. In some embodiments, diffusion of the drug from the implant at a therapeutic level is maintained for at least eight weeks post-implantation. In some embodiments, diffusion of the drug from the implant at a therapeutic level is maintained for at least ten weeks post-implantation. In some embodiments, diffusion of the drug from the implant at a therapeutic level is maintained for at least twelve weeks post-implantation. In some embodiments, the therapeutic level is maintained at an implantation site of the implant.
In some embodiments, the bone component is natural bone. In some embodiments, the bone component is synthetic bone. In some embodiments, the bone component comprises, consists, or consists essentially of bone fragments (e.g., fragments of synthetic bone or fragments of natural bone). In some embodiments, the bone component comprises ground or morselized bone (e.g., morselized synthetic bone or morselized natural bone).
In some embodiments, the implant is a solid. In some embodiments, the implant is molded (e.g., is a putty that can be molded). In some embodiments, the implant is injected as a paste. In some embodiments, the liquid or paste implant hardens upon implantation. In some embodiments, the implant is carveable, so that it may be shaped prior to implantation. In some embodiments, the implant is shaped for use with an implantable prosthesis. In some embodiments, the prosthesis is a fixation tooling, a plate, a screw, a rod, a pin, a nail, or a total arthroplasty of various forms used clinically in orthopedic surgery.
In some embodiments, the implant is a liquid. In some embodiments, the liquid implant is a coating on an implantable prosthesis. In some embodiments, the prosthesis is of a material selected from the group consisting of a metal (including, for example, a metal oxide), a ceramic, a porcelain, an alloy, and a combination of two or more of the foregoing.
In some embodiments, the implant is configured so that upon implantation of the implant, the drug diffuses from the implant in a manner to provide a first bolus after a first period of time following implantation and a second bolus after a second period of time following implantation. In some embodiments, the first period is about one week and the second period is about five weeks. In some embodiments, the first period is about one day and the second period is between about three weeks and about six weeks.
In some embodiments, the degradable polymer comprises, consists, or consists essentially of a polycaprolactone (PCL) polymer (e.g., PCL of various different molecular weights). In some embodiments, the degradable polymer comprises, consists, or consists essentially of a polyethylene glycol (PEG) polymer. In some embodiments, the degradable polymer comprises, consists, or consists essentially of a poly(lactide-co-glycolide) polymer. In some embodiments, the implant further comprises a poragen such as calcium chloride. In some embodiments, the implant is contiguously porous.
In some embodiments of the implant, the bone is present in the uniform mixture in a first quantity by weight and the degradable polymer is present in the uniform mixture in a second quantity by weight, wherein the first quantity is greater than the second quantity. In some embodiments, the first quantity is at least 1.125 times larger than the second quantity, or is at least 1.25 times larger than the second quantity, or is at least 1.5 times larger than the second quantity, or is at least two times larger than the second quantity, or is at least 2.25 times larger than the second quantity, or is at least 2.5 times larger than the second quantity, or is at least 4 times larger than the second quantity, or is at least 5 times larger than the second quantity.
In another aspect, the invention provides a method of making a solid implant, the method comprising: making a uniform mixture including degradable polymer, bone, and a drug; forming the mixture into a desired shape; and curing the shaped mixture to form a solid implant. In some embodiments, the curing step includes subjecting the shaped mixture to heat. In some embodiments, the curing step includes subjecting the shaped mixture to sterilization. In some embodiments, the implant is contiguously porous.
In various embodiments, the bone is present in the uniform mixture in a first quantity by weight and the degradable polymer is present in the uniform mixture in a second quantity by weight, wherein the first quantity is greater than the second quantity. In some embodiments, the first quantity is at least 1.125 times larger than the second quantity, or is at least 1.25 times larger than the second quantity, or is at least 1.5 times larger than the second quantity, or is at least two times larger than the second quantity, or is at least 2.25 times larger than the second quantity, or is at least 2.5 times larger than the second quantity, or is at least 4 times larger than the second quantity, or is at least 5 times larger than the second quantity.
In another aspect, the invention provides an implantable bone void filler comprising, consisting, or consisting essentially of a polymer component, an antibiotic, and a bone fragment (e.g., a bone fragment from a natural cadaver bone source or a synthetic bone fragment). In some embodiments, the filler further comprises a poragen such as calcium chloride. In some embodiments, the filler is contiguously porous. In some embodiments, the polymer component comprises or consists ofpolycaprolactone (PCL) (e.g., PCL of various different molecular weights). In some embodiments, the polymer component comprises, consists, or consists essentially of polyethylene glycol. In some embodiments, the polymer component comprises, consists, or consists essentially of poly(lactide-co-glycolide) polymer. In some embodiments, the polymer component comprises, consists, or consists essentially of aa combination of PEG and PCL. In some embodiments, the polymer component comprises, consists, or consists essentially of a combination of PEG, PCL, and poly(lactide-co-glycolide) In some embodiments, the antibiotic is selected from the group consisting of tobramycin, ciprofloxacin, and vancomycin.
The patent or application file contains at least one drawing executed in color. Copies of this patent with color drawing(s) will be provided by the Patent and Trademark Office upon request and payment of the necessary fee.
The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee.
The foregoing features of embodiments will be more readily understood by reference to the following detailed description, taken with reference to the accompanying drawings, in which:
The present invention is based upon the development of methods and systems for the long-term treatment of orthopedic injuries and conditions.
The published patents, patent applications, websites, company names, and scientific literature referred to herein establish the knowledge that is available to those with skill in the art and are hereby incorporated by reference in their entirety to the same extent as if each was specifically and individually indicated to be incorporated by reference. Any conflict between any reference cited herein and the specific teachings of this specification shall be resolved in favor of the latter.
Definitions. As used in this description and the accompanying claims, the following terms shall have the meanings indicated, unless the context otherwise requires:
By “therapeutic level” is meant a level of a drug required to have a therapeutic effect. For example, if the drug is a growth factor, a therapeutic level of that drug is a level required to enact growth by a cell expressing the receptor to the growth factor. If the drug is an antibiotic, a therapeutic level of that drug is a level required to inhibit and/or prevent growth of a pathogen susceptible (i.e., responsive) to that antibiotic. In some embodiments, the therapeutic level is maintained at the implantation site. In other words, the therapeutic level is maintained throughout the implant itself and within an area of at least 1 centimeter, or at least 2 centimeters, or at least 5 centimeters, or at least 7 centimeters, or at least 10 centimeters, or at least 12 centimeters, or at least 15 centimeters, or at least 20 centimeters from the outside edge of the implanted implant.
By “bolus” is meant level of diffusion from the implant that is greater than a therapeutic level of diffusion. A bolus can sustain its level of diffusion for at least 24 hours, or at least 48 hours, or at least 72 hours following initiation of the bolus, after which the level of diffusion returns to a therapeutic level.
By “implant” is meant an object that can be or has already been implanted into a vertebrate host for medical or therapeutic purposes (including, for example, experimental medical or therapeutic purposes). In some embodiments, the vertebrate host animal is one in need of an implant (e.g., a patient). In some embodiments, the implant is a solid that may be carved to fit a vertebrate animal in need of the implant prior to implantation. For example, if the vertebrate animal is a rabbit and the implant is to treat a damaged radial bone of the rabbit, the implant may start as a 2 mm×2 mm×6 mm solid block that then can be carved (e.g., during the surgery by a medical or veterinary practitioner) to fit the defect precisely prior to implantation. Likewise, if the vertebrate animal is a human and the implant is designed to fill vacant space resulting from removal of an infected total knee arthroplasty and re-insertion of a new knee prostheses, the implant may start as a 1 cm×1 cm×4 cm solid block and then may be carved by the surgical team during the surgery to fit the vacant space. In some embodiments, the implant may be an injectable paste. In some embodiments, the injectable paste may harden in the host following implantation.
As used herein, by “vertebrate host” or “vertebrate animal” is meant any animal that has a backbone. Included as vertebrate hosts are amphibians, reptiles, birds, fish, mammals, and any other type of animal that has a backbone. In some embodiments, the vertebrate host is a mammal including, without limitation, a domesticated animal (e.g., cow, sheep, goat, llama, horse, donkey, pig, camel, ostrich, chicken, emu, etc.), a laboratory animal (e.g., a chimpanzee, a baboon, a rabbit, a rat, a mouse, a hamster, etc.), a pet animal (e.g., cat, dog, parrot, etc.), an endangered animal (e.g., polar bear, tiger, lion, elephant, rhinoceros, blue whale, hippopotamus, etc.), and a human.
By “curing” is meant the toughening or hardening of a polymer-containing material by cross-linking polymer chains. Curing can be accomplished by a number of methods including, without limitation, the addition of a chemical, ultraviolet radiation, electronic beam, or heat. In some embodiments, a solid implant is solidified by curing through the addition of heat a uniform mixture comprising a degradable polymer, bone, and a drug.
By “uniform mixture” is meant a mixture comprising two or more components, where the components are in approximately the same ratio to one another throughout the mixture, when considered at a macro level, even if at a molecular level the distribution of components is not even. For example, if the mixture is a solid block, a full thickness chip taken out of the solid will have approximately the same ratio of components as the larger block. In another example, a centimeter-sized chip removed from the bulk solid will have approximately the same ratio of components as the bulk solid. Likewise, if the mixture is in a liquid state, an aliquot of the mixture will have approximately the same ratio of components as the larger volume from which the aliquot was taken. In some embodiments, where the mixture is in a liquid state, the uniform mixture does not have to be stirred or agitated to maintain the uniformity of its components in the aliquot. In some embodiments, where the mixture is in a liquid state, the uniform mixture must be stirred or agitated for at least one minute prior to taking an aliquot to maintain the uniformity of its components in the aliquot.
By “degradable” is meant that the structure of a polymer can be broken down by hydrolysis and/or by the host's cells or enzymes following implantation of the implant into the host.
By “bone” or “bone graft” is meant synthetic bone or natural bone collected from living vertebrate animals or cadavers. The synthetic bone may be fabricated or synthetic by man, or may be obtained commercially (e.g., the ProOsteon synthetic bone). Natural bone may be collected or harvested from living vertebrate animals, or collected or harvested from cadavers (e.g., cadaver-derived bone, which is allogeneic to the recipient of the implant). In some embodiments, the synthetic bone or natural bone is fragmented or pulverized into micron-sized particulates. In some embodiments, the bone is sterilized (e.g., in an autoclave). In some embodiments, the bone or bone graft is intended to be used clinically as a replacement filler to fill defects and allow production (e.g., by providing a scaffold) of new autologous bone by the host receiving the implant. No MHC-expressing cells or antigen are included in bone (e.g., processed bone). Synthetic bone solids can be readily directly and routinely synthesized from calcium or strontium-based precursors in large batches in commercial ovens, and pulverized (i.e., ground) into pieces and granules, sterilized and certified to be clinical grade filler biomaterial.
By “drug” is meant any type of molecule, or a mixture or complex of molecules, that may be administered to a host with the intention of that molecule or mixture having a therapeutic effect on that host. A therapeutic effect may be a stimulatory effect on autologous or allograft cells (e.g., stimulating growth of cells that repair wounds) or an inhibitory effect on pathogenic cells or agents (e.g., inhibiting growth of bacteria or viruses). Thus, a drug shall include, without limitation, an antibiotic, a growth factor, a vasodilator, a vasoconstrictor, an angiogenesis factor, a chemotactic factor, a cytokine, a pharmaceutical small molecule, a pharmaceutical biological, an enzyme, an antibody, or a mixture or two or more of the preceding. In some embodiments, the drug is water-soluble.
In some embodiments, the drug is thermostable. By “thermostable” is meant that the drug's activity after heating the drug for at least one minute to a temperature higher than 37° C. is at least 80% or 85% or 90% or 95% or 99% of the activity of that drug at 37° C. For example, a thermostable drug is one that has an activity of at least 80% or 85% or 90% or 95% or 99% of a the activity of the drug at 37° C. when the drug is heated for at least one minute or at least two minutes or at least five minutes to a temperature that is at least 55° C. at least 60° C. or at least 65° C. or at least 70° C. or at least 75° C. or at least 80° C. or at least 85° C. or at least 90° C. or at least 95° C., or at least 98° C., or at boiling point, or at the thermal processing point used for drug processing in the graft filler. Some drugs are thermostable (e.g., the thermostable antibiotics described below). For example, thermostable drugs for this intent include, without limitation, tobramycin, gentamicin, vancomycin, and the cephalosporins.
In addition, methods are known for making almost any protein thermostable (see, e.g., Chautard et al., Nature Methods 4(11): 919-921, 2007; Hoseki et al., J. Biochem. 126(5): 951-956, 1999; Liao et al., Proc. Natl. Acad. Sci. USA 83(3): 576-580, 1986; Iwamoto et al., Appl. Environ. Microbiol. 73(17): 5676-5678, 2007).
In some embodiments, the drug is selected based on the need of the host. For example, for periprosthetic infections following an implant, many involve pathogens such as gram positive organisms such as Staphylococcus aureus and Staphylococcus epidermidis, both of which are inhibited by tobramycin. Likewise, gentamicin (another antibiotic) will inhibit E. coli Enteroacteriaeceae and Pseudomonas aeruginosa. Additional antibiotics can be used to address antibiotic resistant strains of these bacteria. For example, vanomycin can inhibit methicillin-resistant Staphylococcus aureus (see Cui et al., J. of Bone and Joint Surgery 89(4): 871-882, 2007).
Despite clinical, material, and pharmaceutical advances, infection remains a major obstacle in orthopedic surgeries. Successful solutions must extend beyond bulk biomaterial and device modifications, integrating locally delivered pharmaceuticals and physiological cues at the implant site, or within large bone defects with prominent avascular spaces. In some embodiments, the invention provides an approach involving coating clinically familiar allograft bone with an antibiotic-releasing rate-controlling polymer membrane for use as a matrix for local drug release in bone in the context of clinical bone graft fillers used in bone-filling and wound space filling functions. The kinetics of drug release from this system can be tailored via alterations in the substrate or the polymeric coating. Drug-loaded degradable polymer (e.g., polycaprolactone and its copolymers) coatings releases bioactive tobramycin from both cadaveric-sourced cancellous allograft fragments and synthetic hybrid coralline or other ceramic bone graft fragments with similar kinetics over a clinically-relevant 6-week timeframe. However the micron-sized allograft particulate provides extended bioactive tobramycin release. Surprisingly, the addition of a GRAS water-soluble polymer (e.g., polyethylene glycol as an example coating poragen) in different amounts to the graft coating formulation dramatically changes tobramycin release kinetics without a significant impact on released antibiotic bioactivity. Incorporation of pre-formulated lipid-microencapsulated tobramycin into the polymer coating did not significantly modify tobramycin release kinetics. In addition to releasing bactericidal concentrations of tobramycin, antibiotic-loaded allograft bone provides recognized beneficial osteoconductive potential, encouraging bone in-growth and tissue neogenesis, possibly decreasing orthopedic surgical infections with improved filling of dead space and new bone formation.
However, despite significant multi-disciplinary clinical innovations combined with biomaterial and pharmaceutical approaches, including now-standard systemic antibiotic prophylaxis and new bone grafting biomaterials, infection remains a major complication in total joint revision surgery, with rates ranging from 8-15% and relapsing infection representing a significant threat (20-30%) to wound healing (Conterno and da Silva Filho, Cochrane Database Syst Rev, (2009) CD004439; Landersdorfer et al., Clin Pharmacokinet, 48 (2009) 89-124). This represents a considerable healthcare burden as the demand for total joint replacements continues to rise with the aging population. The number of post-arthroplasty infectious complications is projected to increase from current levels of 17,000 cases to an overwhelming 266,000 cases annually by 2030 (Bernthal et al., PLoS One, 5 (2010) e12580.
Despite clinical routine of filling avascular dead spaces with bone graft filler materials and bone cements and prophylactic systemic antibiotics, infection remains the second most prevalent complication associated with orthopedic surgeries. The most common infection is osteomyelitis (Aronin et al., Biomaterials, vol. 31, pp. 6417-6424), acutely caused by perioperative introduction of pervasive pathogens such as staphylococcus (Roald et al., Blood Coagul. Fibrinolysis, vol. 5, pp. 355-63, 1994) into the avascular spaces surrounding the orthopedic graft. In fact, Stapholococcal strains account for more than 90% of osteomyelitis cases (Gogia et al., Semin Plast Surg, vol. 23, pp. 100-7, 2009).
In some cases, the high infection rate may be due to antibiotic resistance that develops from interactions between the environmental conditions, natural selection pressures, and antibiotic misuse (Peters et al., J Infect Dis, vol. 197, pp. 1087-93, 2008). The emergence of antimicrobial resistance based on natural selective pressures and complicated by clinical antibiotic overuse is typified by the increasing number of impotent antibiotics due to the widespread use of single systemic antibiotic therapy. Subsequently, local antibiotic combination therapy may serve to alleviate antibiotic resistance concerns. Combination therapy aims to use multiple antibiotics that produce the analogous or even enhanced therapeutic effects with lower doses of each antibiotic. The utilization of combination therapy has proven helpful in clinical practice when treating chronic staphylococcal infections and the subsequent decreased antibiotic susceptibility of the infection (Bernard et al., J Antimicrob Chemother, vol. 53, pp. 127-9, 2004). In addition to widespread systemic single antibiotic overuse, local antibiotic delivery can also spur antibiotic resistance. Antibiotic concentrations below the therapeutic dose, such as that observed with commercially available drug releasing bone cement (PMMA) and allograft bone soaked in antibiotic solutions in the surgical theater (off-label), inadvertently promote drug-resistance in bacteria (Diefenbeck et al., Injury, vol. 37 Suppl 2, pp. S95-104, 2006). “Off label” preparations such as this are unpredictable, lacking quality controls that may inadvertently promote antibiotic-resistance (Kanellakopoulou and Giamarellos-Bourboulis, Drugs, vol. 59, pp. 1223-32, 2000; Ayers, N., J Polym Sci A Polym Chem, vol. 46, pp. 7713, 2008). Indeed, current methods of simple physical absorption of drug to allograft bone can result in only a bolus release (Kanellakopoulou and Giamarellos-Bourboulis, supra), with no sustained release to combat further infections. Polymer-controlled antibiotic release from allograft bone material is a desirable alternative as it allows for long-term antibiotic success by tailoring antibiotic dose, combination drug therapies, control of drug dose release kinetics and local drug delivery from an osteoconductive clinically approved bone filler biomaterial device.
In combating infection, an antibiotic administered with a bone implant will, in some embodiments of the present invention, hit the “sweet spot” in combating initial infection, while continuing to elute the antibiotic for a prolonged period of time (e.g., 6-8 weeks post-implantation, or even up to 10 weeks or longer post-implantation).
To overcome this problem, in some embodiments, the invention described herein resulted from efforts to investigate the beneficial effect of drug/polymer and drug/physiological fluid solubility and miscibility on drug-release profile, water-soluble antibiotics (e.g., vancomycin or oxacillin) and non-water soluble antibiotics (e.g., ciprofloxacin or rifampicin) were released from a polymer-loaded controlled releasing membrane with and without the incorporation of polymer non-solvent (e.g., water) into the device formulation. Moreover, the danger of multi-drug resistant pathogens will be investigated by considering the combinatorial therapeutic efficacy of other clinically important antibiotics (e.g., ciprofloxacin, rifampicin, and vancomycin) against S. aureus as released from a local drug delivery system using a polymer-controlled releasing membrane coated onto allograft bone. This can be studied as two different graft filler materials formulated each with a single antibiotic and controlling polymer membrane, and mixed in the wound site in varying proportions, or alternatively by combining more than one drug together into a single bone graft filler within a polymer rate-controlling membrane. Tailoring the release kinetics as a function of antibiotic solubility should provide clinicians with a long-term, antibiotic delivery system that can be customized and combined to fit each patient's needs while concurrently mitigating the development of antibiotic resistance.
By developing a local drug delivery system that controls the release of antibiotic via a degradable polymer (e.g., polycaprolactone (PCL) and its copolymers) membrane coated on an implantable allograft bone delivery vehicle, an increase in the bioactive longevity of antibiotic therapy is anticipated (Roald et al., Blood Coagul. Fibrinolysis, vol. 5, pp. 355-63, 1994). Furthermore, combinatorial antibiotic controlled release formulations on these bone graft fillers prepared for each antibiotic will provide clinicians with an “a la carte” method for customizing antibiotic treatment that can be tailored to meet each patient's needs while mitigating the development of bacterial resistance due to systemic and prolonged antibiotic overuse and poor patient compliance.
Autograft bone, or patient-harvested bone, is the gold standard for bone grafting, providing a highly compatible, bioactive, structural matrix as the basis for wound healing. However, cellular death during transplantation, inadequate sourcing due to other pathologies, harvest site morbidity, pain, and cosmetic disfigurement, culminate in a substantial 8.5-20% complication risk, including acute and chronic or recurring infection (Nandi et al., Indian J Med Res, vol. 132, pp. 15-30, 2010; Kundu et al., J Mater Sci Mater Med, vol. 21, pp. 2955-69, 2010; Aronin et al., Biomaterials, vol. 31, pp. 6417-24). Thus, allograft or cadaveric-sourced bone tissue has become an increasingly popular defect and wound packing material, increasing 15-fold over the past decade to now account for almost a third of the over 500,000 orthopedic graft procedures performed annually in the United States to treat traumatic or other boney defects (Aronin et al., supra; Kanellakopoulou and E. J. Giamarellos-Bourboulis, supra). Importantly, allograft bone is processed to remove all cellular and proteinaceous components, leaving only the osteoconductive, and to a more limited extent, osteoinductive mineral component of the graft to provide a structural template for orthopedic repair, and promote integration and turnover by the patient's natural osteoclast and osteoblast populations.
Similarly, synthetic bone (e.g., comprising calcium and/or strontium based ceramic filler biomaterials) also promotes integration and turnover by the patient's natural osteoclast and osteoblast populations. On such synthetic bone is the commercially available ProOsteon substrate (available from, for example, Biomet, Inc., Warsaw, Ind.).
Successful solutions to implant-centered infection might best integrate local, rate-controlled drug delivery with appropriate wound and defect filler materials, particularly for implants and large bone defects with prominent avascular spaces or where penetration from systemic antibiotic administration is compromised (Landersdorfer et al., Clin Pharmacokinet, 48 (2009) 89-124; C. Ketonis et al., Tissue Eng Part A, 16 (2010) 2041-2049; C. Ketonis et al., Clin Orthop Relat Res, 468 (2010) 2113-2121; C. Ketonis et al., Antimicrob Agents Chemother, 55 (2011) 487-494; C. Ketonis et al., Bone, 48 (2011) 631-638; N. M. Mathijssen et al., BMC Musculoskelet Disord, 11 (2010) 96).
Treating bone infections is intrinsically complicated by poor bioavailability and drug pharmacokinetics in bone that limit efficacy of systemically administered antibiotic therapy. Bone vascular physiology enables a niche for diverse types of opportunistic pathogens introduced at the time of injury, intraoperatively, or later by hematogenous sourcing to produce difficult-to-treat infections. Antibiotic penetration into the bone as well as the limited vasculature of the affected bone must be considered when designing a clinical treatment strategy (Landersdorfer et al., Clin Pharmacokinet, 48 (2009): 89-124; Chen et al., Arch Orthop Trauma Surg, 125 (2005): 369-375). Although systemic intravenous antibiotics are often sufficient in combating these opportunistic pathogens, the negative impact of a standard 4-6 week course of antibiotics cannot be neglected. Inappropriate use of antibiotic therapies, such as poor selection, inadequate dosing, broad-spectrum antibiotic overuse, and poor patient therapy follow-through, have all accelerated pressure towards multi-drug resistant microbes. The CDC reports an alarming rise in the antibiotic resistance of the major pathogen, Staphylococcus aureus, to at least one of the most common antibiotics from 2% in 1972 to 63% by 2004 (MRSA vs. MSSA) (see Office of the Associate Director for Communication, 2006, http://www.cdc.gov/media/pressrel/r061019.htm?s_cid=mediarel_r061019_x, accessed Jun. 23, 2011). Furthermore, some systemically administered antibiotics may not achieve therapeutic levels in bone, inadvertently supporting the development of resistance. Therefore, options for local sustained antimicrobial therapies are increasingly attractive. A local drug delivery mechanism overcomes bioavailability and systemic delivery issues, limits development of systemic antibiotic resistance while delivering sustained amounts of drug sufficient to both resist and eliminate microbial infection beyond an acute time course (Patzakis and Zalavras, J Am Acad Orthop Surg, 13 (2005): 417-427). Local delivery of antibiotics offers effective killing using higher doses (up to 1000-fold greater than systemically delivered (Costerton, J. W., Rev Infect Dis, 6 (1984): 608-616; Diefenbeck et al., Injury, 37 Suppl 2 (2006): S95-104) precisely at the site of infection while avoiding systemic toxicity associated with high doses (Diefenbeck et al., Injury, 37 Suppl 2 (2006): S95-104). Unfortunately, many approaches to achieve local antibiotic release from bone grafts with desirable therapeutic kinetics—either actively or passively—are often characterized by an early bolus release and subsequent slow leaching of antibiotic at sub-therapeutic levels that may also promote antibiotic-resistance (Kanellakopoulou and Giamarellos-Bourboulis, Drugs, 59 (2000): 1223-1232; Diefenbeck et al., Injury, 37 Suppl 2 (2006): S95-104). Thus, improved control over local drug release in terms of dose control and duration is likely a necessity for efficacious long-term delivery and antimicrobial efficacy.
Drug delivery directly to bone in general and also to avascular traumatized or infected bone presents a pharmaceutical and pharmacokinetic challenge. Currently, bone grafts are used for musculoskeletal mechanical support as well as space filling and osteoconductive foundation for new bone deposition and healing. Incorporating a space filling material with the controlled degradation of a synthetic polymer may provide features appropriate for prophylactic controlled drug delivery. Importantly, most synthetic polymers alone are inappropriate as bulk materials for orthopedic needs that may require mechanical integrity for up to a year or more and also bone regeneration/healing induction. As a clinically recognized biomaterial, resorbable polycaprolactone (PCL) and its copolymers may exhibit the requisite enhanced temporary structural functionality sufficient for bone implant use while also providing appropriate characteristics for rate-controlled drug delivery and degradability (Lowry et al., J Biomed Mater Res, 36 (1997): 536-541; Coombes et al., Biomaterials, 25 (2004): 315-325; H. I. Chang et al., J Control Release, 110 (2006): 414-421). Thus, with PCL's precedent use in bone implants, PCL and its copolymers may offer a significant opportunity to endow clinically familiar bone graft filler materials with an antibiotic-releasing, rate-controlling coating for extended drug delivery.
The controlled hydrolytic degradation of PCL (Lam et al., Biomed Mater, 3 (2008): 034108; Hutmacher et al, J Tissue Eng Regen Med, 1 (2007): 245-260) offers a versatile range of times for extended release kinetics under certain physiological circumstances in tissues. However, a mixed multi-polymer barrier might be more appropriate (Wei et al., Int. J. Pharm, 381 (2009): 1-18). In this regard, polyethylene glycol (PEG), a common biomaterial generally regarded as safe by the FDA may be incorporated into the PCL (co)polymer coating formulation as a poragen and to improve drug loading and solubility (i.e., for certain poorly water-soluble antibiotics also not miscible with PCL and its copolymers), but may also provide more versatile release kinetics for different dosings or applications. Thus, antibiotic loading and subsequent release kinetics might be adjusted and tailored via the rate-controlling polymer coat formulation.
PCT Publication No. WO 2011/127149 (from International Application No. PCT/US2011/031394), and US Patent Publication No. 2009/0324683; (both of which are hereby incorporated by reference in their entireties) describe
The concept of some of the non-limiting bone graft implants of the invention is depicted schematically in
The graph depicted in
Currently, while synthetic bone graft materials have a lengthy clinical pedigree, no FDA approved allograft bone therapies incorporate an integrated antibiotic release scheme as combination medical devices. In accordance with various embodiments of the present invention, the bone graft is acting in its primary mode of action as a medical device (bone graft filler) and the drug-releasing modality is a secondary mode of action. In addition to polymer barrier coating characteristics, other factors can be exploited in this modular combination device approach. Graft surface area (micron-scale morselized bone can be milled to have a higher surface area for drug release than cancellous crouton fragments), diverse differential implant packing (i.e., mixing of large allograft cortical croutons with morselized allograft cancellous granules either as separate coated formulations or within a single coated preparation) and antibiotic solid microencapsulation (e.g., in common, clinically routine starch or solid-dosage form encapsulating matrices) prior to drug dose loading all provide a range of customizable drug loading and release options appropriate for tailoring and customizing bone defect combination devices for better mitigating infectious risks in orthopedic and connective tissue surgical implant and repair sites. While directly soaking allograft bone filler materials in antibiotic has been studied extensively (see, e.g., Witso et al., Acta Orthop, 76 (2005): 481-486; Witso et al., Acta Orthop Scand, 70 (1999): 298-304; Witso et al., Acta Orthop Scand, 71 (2000): 80-84; Darley and MacGowan, J Antimicrob Chemother, 53 (2004): 928-935; Rhyu et al., Int Orthop, 27 (2003): 53-55; Winkler et al., J Antimicrob Chemother, 46 (2000): 423-428; Witso et al., Acta Orthop Scand, 75 (2004): 339-346; Lindsey et al., Clin Orthop Relat Res. 291: 303-312 (1993), the idea of endowing this matrix with a true local controlled release strategy has not.
In some embodiments, the non-limiting allograft and synthetic bone matrix-antibiotic-polymer combination devices (i.e., the bone implants) described herein (and shown schematically in the Figures) permit precise, uniform tobramycin drug loading (via the polymer overcoat) to retain the drug release depot at the surgical site controlled by polymer (PCL (co)polymer±PEG mixtures) coating swelling, porosity and degradation by hydrolysis. The bone implants described herein exhibit long-term antibiotic release at the wound or implantation site and maintenance of therapeutic antimicrobial drug concentrations at the implantation site beyond 6 weeks, beyond 8 weeks, or even beyond 10 weeks post-implantation. In some embodiments, the release (i.e., diffusion) of the drug at a therapeutic levels (e.g., locally at the implantation site) is maintained for at least eight weeks. In some embodiments, the release of the drug at a therapeutic level (e.g., at the implantation site) is maintained for a time longer than the amount of time a pathogen can remain in either a metabolically active or a senescent state (e.g., in a biofilm).
In some embodiments, the versatility of at least some of the bone implants of the invention (and fabrication methods thereof) is depicted schematically in
In yet another embodiment (
In some embodiments, an implant (e.g., generated using the generation 3 fabrication method) can be made of a polymer component, a bone component, and a drug component. In some embodiments, for example, the bone component may be ground or morselized and/or may be natural bone or synthetic bone (e.g., ProOsteon). In some embodiments, for example, the drug may be an antibiotic such as tobramycin. In some embodiments, for example, the polymer component may be PCL, or may be a PCL and a PEG combination, or may be a PCL, a PEG, and a poly(lactide-co-glycolide) combination. In some embodiments, the polymer component may also include a poragen such as calcium chloride. Calcium chloride is a biocompatible water-soluble salt, which is being looked at as pore former in the modified formulation. This water-soluble salt is expected to dissolve in less than 24 hours to create initial porosity to allow ingress of fluid and cells. The poly(lactide-co-glycolide) will degrade faster than the PCL and will release the initial tobromycin load. It is expect the slower degrading PCL will then deliver the later drug load. Poly(lactide-co-glycolide) is a biocompatible degradable polymer used commonly in sutures, fracture fixation deices and microsphere in drug delivery, It is more hydrophilic than PCL and as such degrades faster than PCL.
Note that in
In
As shown in
In accordance with the present disclosure, the implants described herein (e.g., those depicted in
Accordingly, in a first aspect, the invention provides an implant comprising, consisting, or consisting essentially of a uniform mixture of degradable polymer, bone, and a drug. In some embodiments, the drug comprises an antibiotic. In some embodiments, the implant is configured so that upon implantation of the implant into a host at an implantation site, the drug diffuses from the implant at a therapeutic level. In some embodiments, the host is a vertebrate animal. In some embodiments, diffusion of the drug from the implant at a therapeutic level is maintained for at least eight weeks post-implantation. In some embodiments, diffusion of the drug from the implant at a therapeutic level is maintained for at least ten weeks post-implantation. In some embodiments, diffusion of the drug from the implant at a therapeutic level is maintained for at least twelve weeks post-implantation. In some embodiments, the therapeutic level is maintained at an implantation site of the implant.
In some embodiments, the implant is a solid. In some embodiments, the implant is molded. In some embodiments, the implant is carvable, so that it may be shaped prior to implantation. In some embodiments, the implant is shaped for use with an implantable prosthesis. In some embodiments, the implant is shaped for use with an implantable prosthesis. In some embodiments, the prosthesis is a fixation tooling, a plate, a screw, a rod, a pin, a nail, or a total arthroplasty of various forms used clinically in orthopedic surgery.
In some embodiments, the implant is a liquid. In some embodiments, the implant is a paste. In some embodiments, the implant is a putty. In some embodiments, the implant is a coating on an implantable prosthesis. In some embodiments, the prosthesis is of a material selected from the group consisting of a metal (including, for example, a metal oxide), a ceramic, a porcelain, an alloy, and a combination of two or more of the foregoing.
In some embodiments, the implant is configured so that upon implantation of the implant, the drug diffuses from the implant in a manner to provide a first bolus after a first period of time following implantation and a second bolus after a second period of time following implantation. In some embodiments, the first period is about one week and the second period is about five weeks. In some embodiments, the first period is about one day and the second period is between about three weeks and about six weeks.
In some embodiments of the implant, the bone is present in the uniform mixture in a first quantity by weight and the degradable polymer is present in the uniform mixture in a second quantity by weight, wherein the first quantity is greater than the second quantity. In some embodiments, the first quantity is at least 1.125 times larger than the second quantity, or is at least 1.25 times larger than the second quantity, or is at least 1.5 times larger than the second quantity, or is at least two times larger than the second quantity, or is at least 2.25 times larger than the second quantity, or is at least 2.5 times larger than the second quantity.
In another aspect, the invention provides a method of making a solid implant, the method comprising: making a uniform mixture including degradable polymer, bone, and a drug; forming the mixture into a desired shape; and curing the shaped mixture to form a solid implant. In some embodiments, the curing step includes subjecting the shaped mixture to heat.
In various embodiments, the bone is present in the uniform mixture in a first quantity by weight and the degradable polymer is present in the uniform mixture in a second quantity by weight, wherein the first quantity is greater than the second quantity. In some embodiments, the first quantity is at least 1.125 times larger than the second quantity, or is at least 1.25 times larger than the second quantity, or is at least 1.5 times larger than the second quantity, or is at least two times larger than the second quantity, or is at least 2.25 times larger than the second quantity, or is at least 2.5 times larger than the second quantity.
In another aspect, the invention provides an implantable bone void filler comprising a polycaprolactone (PCL) polymer, an antibiotic, and a bone fragment. In some embodiments, the antibiotic is selected from the group consisting of tobramycin, ciprofloxacin, and vancomycin.
In some embodiments, the implant is in contact with (e.g., in combination with or coated onto) a prosthetic. In some embodiments, the prosthetic is implanted. By “prosthetic” is meant a wholly artificial structure that is or can be implanted into a vertebrate host animal to aid in functional restoration of a tissue, including bone. Prosthetics include, without limitation, metal prosthetics (e.g., titanium, steel, gold, platinum, etc.), ceramic, and porcelain in the form of multiple tools and stabilizing, or structural aids, including plates, screws, rods, cannulae, fusion cages, nails, pins, meshes, cups, sutures, and joint arthroplasty devices. The prosthetic need not be solid. For example, a prosthetic may be porous. A prosthesis may also be flexible, or may be both porous and flexible. In some embodiments where the prosthesis is porous and the implant is liquid, the liquid implant may coat the surfaces or walls of the pores of the prosthesis. Such coating may be done prior to implantation, or during implantation.
In this example, an implant was fabricated and tested for its ability to diffuse drug for a prolonged amount of time in vitro.
For these studies, the following methods were used.
Fabrication of polymer-coated allograft fragments. To do this, cancellous allograft bone fragments (Miami Tissue Bank) or ProOsteon 500R (BioMet, Warsaw, Ind., USA) were weighed and like-size and mass fragments were selected for each cohort (n=3). Alternatively, micron-size allograft bone particulate matter (Miami Tissue Bank, Miami, Fla., USA) was partitioned into 100 mg aliquots for polymer-drug coating. PCL (10 kD, Sigma Aldrich, St. Louis, Mo., USA) (60 mg/ml) was dissolved in acetone (Thermo Fisher Scientific, Waltham, Mass., USA) at 45° C. Tobramycin (MP Biomedicals, Solon, Ohio, USA) was suspended as “free” (i.e., unencapsulated) drug in PCL acetone solutions at 10% weight/volume. Alternatively, certain PCL coating formulations included tobramycin commercially microencapsulated in vegetable triglycerides (70 w/w % tobramycin, lot#TM150-70-30, Maxx Performance Inc., Chester, N.Y., USA). Formulations and cohorts are detailed in Table 1.
Each cohort was made with 10 kD PCL dissolved in acetone at 45 C. If PEG was included then it was dissolved in water first and then the PEG water solution was added to the PCL acetone solution. All tobramycin (regardless of encapsulation state) was added to the polymer solution(s) as a dry powder to create the formulation used to either dip coat allograft or ProOsteon fragments or to solvent cast the particulate as described in the next paragraph and schematically depicted in FIGS. 7B and 9A-9C. The solvent used with the fabrications in all of these cohorts was acetone. These fabrications were generated with the PEG dissolved in water and the PCL being dissolved in acetone. [HOW IS COHORT 6 MADE?] An unloaded polymer bone control was included in all analyses (data not shown). Dip-coated cohorts were prepared by placing allograft bone into the PCL/free tobramycin solution at room temperature.
This fabrication method is schematically depicted in
Note that for cohort 6, allograft particulate was coated in individual aluminum trays with a total of 2 ml of polymer/drug solution (500 ul PCL with free tobramycin, 1 ml PEG with microencapsulated tobramycin, 500 ul PCL with free tobramycin) in a layer-by-layer (LBL) fashion with alternating layers of PCL and PEG. To create a polymer/drug layer, bone graft particulate was mixed twice in each polymer/drug solution and the solvent was allowed to flash off, leaving coated particulate. The dried particulate-containing polymer film was ground to granules again using a weighing spatula and the next layer was applied according to the same protocol.
Drug Release Measurements.
To measure drug release from the fabricated implants, each coated allograft bone sample was placed into 3 ml of phosphate buffered saline pH 7.4 (PBS, Fisher Scientific, Waltham, Mass., USA). The complete volume (called the release volume because it contains the released drug) was drawn off and replaced at 30 minutes, 1 hour, 2 hours, 4 hours, 8 hours, 24 hours, 72 hours, and each week for up to 6 weeks to simulate sink conditions. Kinetics of release from each formulation were assessed via a 96-well fluorescent assay previously reported (Sevy et al., Biomed Sci Instrum, 46 (2010) 136-141). Briefly, 75 ul of each release sample was added to 75 ul of isopropanol in wells within black-masked 96-well plate (Fisher Scientific, Pittsburgh, Pa., USA). OPA working solution (150 ul) (50 ul of o-phthaldehyde (OPA, Sigma Aldrich, St. Louis, Mo., USA) stock solution in 1 ml of 0.5M potassium borate buffer pH 10.5) was added to each well and incubated for 30 minutes prior to assessing the fluorescence of the tobramycin/OPA derivative (Biotek spectrophotometer, ex=360 nm, em=460 nm) using Gen5 1.09 software (BioTek, Winooski, Vt., USA). Each cohort contained a certain number of reference samples (n=3, 6, or 9) from which tobramycin was not released over time but instead the entire coating was dissolved in 1 ml of chloroform (Thermo Fisher Scientific, Waltham, Mass., USA) for approximately 5 minutes and 1 ml of water was used to phase extract tobramycin from the polymer solution by vortexing for 30 seconds and then centrifuging at 15,000 rpm for 2 minutes and 30 seconds. These samples were considered 100% release samples and all amounts of tobramycin released over time from coated grafts were normalized to their cohort-matched 100% release value as well as to the unloaded polymer bone control, and reported as a percent to facilitate direct comparison of release from different polymer formulations. Tobramycin from fragments coated with a PCL-water non-solvent system was phase extracted after 8 weeks of release into PBS using chloroform and water to verify the mass balance of the system (data not shown).
As shown in
In various embodiments of the invention, the ability to modulate drug release kinetics is useful for combating bacterial infection. Indeed, drug release can be tailored to match the rate of bone growth and remodeling. As shown in
Data Analysis:
The amount of tobramycin released in each sample was calculated based on the linear regression of the fluorescent units (FU) for each standard. Percent drug release was calculated by dividing the amount of tobramycin released by the amount of tobramycin detected in the dissolved coating (100% release) multiplied by 100. All formulations were tested in triplicate (biological and technical replicates) and Excel was used to calculate the propagating standard deviation. Pairwise one-way ANOVAs were used to identify significant differences (p<0.05 for significance). Particular comparisons to be tested were selected in advance and were reported individually rather than as a group and therefore a multiple comparison correction was not necessary (see Dunnett and Goldsmith, “When and how to do multiple comparisons”, in: C. R. Buncher, J.-Y. Tsay (Eds.) Statistics in the Pharmaceutical Industry, Chapman and Hall/CRC New York, 2006, pp. 421-452).
High Performance Liquid Chromatography (HPLC):
Standard concentrations of tobramycin were resuspended in acetonitrile-water (52:48). All samples were analyzed in triplicate using high pressure liquid chromatography (HPLC), with a pre-column OPA derivatization (see Sevy et al., Biomed Sci. Instrum., 46 (2010) 136-141). Data was collected from both a fluorescence detector (ex=350 nm, em=450 nm) as well as UV-V is detector (340 nm). Samples were analyzed using a Hypersil GOLD HPLC column (100×4.6 mm, Thermo Fisher Scientific, Waltham, Mass., USA) and ChromQuest 5.0 (Thermo Fisher Scientific, Waltham, Mass., USA) software on a Finnigan Surveyor (Thermo Fisher Scientific, Waltham, Mass., USA) system. Each sample (10 ul) was injected using a 2 ml/min flow rate. The mobile phase was mixed 0.02M phosphate (pH 6.5):acetonitrile (52:48). The area under the tobramycin peak was plotted against standard concentration and data were fit by linear regression as a standard curve, used to calculate the concentration of unknown drug release samples.
Microbiology:
Release samples (500 ul per experiment) for all microbiology studies were concentrated in a vacuum centrifuge (Labconoco Centrivap, Kansas City, Mo., USA) overnight at ambient temperature and prepared in low-bind, non-tissue culture-treated 96-well microtiter plates according to their subsequent experimental use (i.e., MIC: round bottom, ZOI: flat bottom). All samples were stored dry at 4° C. until use. Antimicrobial activity after concentration as well as storage was confirmed with control conditions.
Bacteriostatic Assay:
LB broth (100 ul, Becton Dickinson, Franklin Lakes, N.J., USA) was added to each well of the round bottom 96-well plate to reconstitute the dried drug release samples. Each well was inoculated with 105 CFU in 200 ul of a liquid culture of E. coli (ATCC 25922, American Type Culture Collection, Manassas, Va., USA). Liquid bacterial cultures were prepared using a sterile swab to select 1-3 isolated colonies from a blood agar plate (Remel, Lenexa, Kans., USA). Inoculated plates were incubated overnight at 37° C. Released drug activity, assessed by bacterial growth inhibition, was visually determined by comparing known standard tobramycin concentrations. Growth inhibition was positive if the visual turbidity of bacterial growth media differed from the positive control by 80%. Negative growth was designated when the well was free of a visible bacterial pellet.
Zone of Inhibition (ZOI):
For ZOI experiments, release samples were dried onto 6 mm Whatman 1 filter paper disks. Muller Hinton agar plates (Fisher Scientific, Waltham, Mass., USA) were prepared by streaking E. coli (ATCC 25922 from American Type Culture Collection, Manassas, Va.) to create a confluent lawn of bacterial growth (turbidity adjusted to a 0.5 McFarland standard using a nephelometer (Phoenix Spec, BD Diagnostic Systems, Franklin Lakes, N.J., USA)). Disks containing the dried-down drug from release samples were then placed with a minimum distance of 24 mm between each disk and the side of the plate. Plates were incubated overnight at 37° C. Calipers were used to measure the diameter of the zone of inhibition around each disk.
Sample Fabrication and Drug Release Assay. Tobramycin is a clinical drug of choice used to treat orthopedic infections; however, due to associated nephro- and ototoxicity (Begg and Barclay, B. J. Clin. Pharmacol. 39: 597-703, 1995), maintaining adequate drug concentration to combat opportunistic microbes in bone using traditional delivery mechanisms (intravenous) may be unachievable. Therefore, a local, polymer-controlled delivery of tobramycin directly to the bone using a bone graft delivery vehicle was investigated. Samples were fabricated according to Table 1 with all allograft fragments (approx. 6 mm×5 mm×4.5 mm) and ProOsteon 500R® fragments (approxl Omm×8 mm×7 mm) being dip-coated to add approximately 22 mg of drug-releasing coating to the fragment's initial weight. Importantly, cohorts of bone graft fragments were weight matched to limit their variability. Alternatively, micron-size allograft particulate (cohorts 3-6) was coated in individual aluminum trays via a solvent evaporation procedure. The particulate-containing polymer film was ground prior to release. Theoretical amounts of tobramycin applied to each sample were calculated based on the weight of coating applied to the allograft bone material and the percent of tobramycin included in the formulation. Tobramycin was released from the polymer coating on each cohort into PBS. PBS was sampled at designated time points and replaced to simulate sink conditions. For the data depicted in
Release kinetics. Tobramycin is very water-soluble and thermostable during formulation as evidenced by no loss in bioactivity after included in a PCL formulation (data not shown) (Mousset et al., Int Orthop, 19 (1995): 157-161). However, detection of this small molecule aminoglycoside antimicrobial in a sample is complicated by lack of a unique optical signature. Therefore, tobramycin was derivatized with o-phthaldehyde (OPA) (Sevy et al., Biomed Sci Instrum, 46 (2010): 136-141). This reaction yields a chromophore by chemically coupling with primary amines on the drug, producing fluorescence signals with a dynamic range from 0 to 8 mg/ml and a limit of detection of 62.5 ug/ml (Sevy et al., Biomed Sci Instrum, 46 (2010): 136-141). The OPA derivatization reaction was verified via HPLC detection of tobramycin in the presence or absence of OPA (data not shown). In the absence of OPA, tobramycin did not elicit any absorbance or fluorescence signal. Furthermore, inherent OPA fluorescence was not detected in the absence of tobramycin. Thus, release of tobramycin from a variety of polymer formulations was compared using an OPA derivatization in a 96-well assay format as previously reported and validated by mass spectrometry (Sevy et al., Biomed Sci Instrum, 46 (2010): 136-141). To facilitate comparison, the measured amount of drug released at each time point was normalized, in a cohort specific manner, by the average (n=3-9) of the detected tobramycin after complete dissolution of the polymer coating in chloroform and subsequent extraction of tobramycin from the chloroform polymer solution with water. This amount was assumed to be 100% of the tobramycin added to the polymer coating system (100% release). The phase extraction procedure was controlled by determining the percent recovery of both free drug (approximately 100% recovery) and also microencapsulated (approximately 82% recovery) tobramycin. Thus, for microencapsulated samples, 18% of the amount of the 100% drug release sample was added and this value was used for all subsequent calculations. Moreover, the validated coating-dissolution, phase-extraction method was applied to time course release samples, allowing determination of the mass balance. After 8 weeks of release into PBS, between 97-100% of the drug was recovered from a PCL-tobramycin coating (data not shown).
As shown in
As shown in
The superior release performance of coated allograft particulate led to its use in all subsequent experiments in this Example 1 with the notable exception of the synthetic bone graft filler, ProOsteon 500R® particulate, which were not micron-sized and had a porous structure reported similar to cancellous bone (pore size reported to be 280-770 um with 55% porosity; cancellous allograft pore size reported to average between 400-500 um with porosity ranging between 60-77% (Bloebaum et al, Clin Orthop Relat Res, (1994): 2-10). Not surprisingly, when ProOsteon 500R® fragments were dip-coated and vacuum dried analogously to the allograft fragments (cohort 1), tobramycin release did not differ significantly in its kinetics when compared to allograft coated crouton fragments (FIG. 14A), demonstrating the substrate independence of the drug-releasing coating system described on graft materials with similar microstructural features.
Impact of Coating Formulation on Tobramycin Release Kinetics.
In solution-based drug formulating and polymer vehicle coating, component compatibility issues in drug-solvent-polymer solubility, mutual miscibility and controlled solution stability are important design criteria (J. Liu et al, J. Pharm. Sci. 93: 132-143, 2004). Consideration of thermodynamics predictors of these properties (i.e., matching appropriate Hildebrand solubility parameters of the drug, polymer, and solvent(s), such as with PCL (δ (delta)=20.2 (Bordes et al., Int. J. Pharm. 383: 236-243, 2010) in acetone and tobramycin) facilitate sustained and controlled drug release and can help avoid phase separation (J. Liu et al., supra; Huang et al., J. of Applied Polymer Sci., 100: 2002-2009, 2006). Unfortunately, tobramycin is highly soluble in water and marginally soluble in alcohols while PCL is insoluble in water, highly soluble in chloroform, and soluble in acetone at elevated temperature. The dissimilarity in component solubility (i.e., powdered tobramycin added to acetone solvated PCL creates a tobramycin suspension upon mixing) and resulting solution heterogeneity and possible phase separation as a drug delivery system will impact the kinetics of drug release.
To limit the tobramycin burst release kinetics within the first 24 hours, potentially due to phase separation, some coating formulations were modified to include a 45% w/v aqueous polyethylene glycol (PEG δ (delta)=22.9 (J. Liu et al., supra) 20 kD) feed solution (cohort 4), which would, in theory, retain tobramycin in a more compatible aqueous phase thereby mitigating phase separation and creating a more homogeneous polymer-drug formulation for coating (see
As shown in
Tobramycin drug release kinetics were further modified using commercially microencapsulated tobramycin to further slow and extend the duration of drug release. The modifications of the fabrication method and the implants resulting therefrom are shown in
Since the addition of microencapsulated tobramycin into a PEG/PCL-based coating formulation retarded the tobramycin release kinetics, the influence of the coating application technique was also investigated (see
Regardless of coating formulation, antimicrobial activity was confirmed via in vitro bacteriostatic assays based on a modification of the standard techniques for determining the minimal inhibitory concentration (MIC) for an antibiotic (data not shown), as well as classic zone of inhibition or radial diffusion assays (see
Tobramycin release was primarily affected by the allograft material morphology (larger porous crouton fragments or micron-sized porous particulate, see
In addition to drug release kinetics dependence on the geometry of the underlying substrate, the polymer formulation can also be engineered to alter the rate of drug release. Based on the disparate release kinetics measured upon addition of the 45% PEG aqueous solution (cohort 4; PCLF/PEG in
Thus, from this Example 1, the degradable polymer-controlled, antibiotic-releasing bone graft system described was shown to be able to successfully deliver tobramycin antibiotic in vitro over 6 weeks, offering a distinct performance advantage over current antibiotic-releasing technologies for bone that may inadvertently promote both infection and bacterial antibiotic resistance. Furthermore, the broad implications of polymer-mediated control over local drug release kinetics with some degree of versatility presents an attractive alternative technique for improved local delivery of different classes of bioactive molecules from tissue implants, particularly in a diffusion-limited tissue such as bone defects. A facile, convenient drug fluorescence assay was developed to evaluate drug release kinetics from a variety of tobramycin-loaded PCL-coated bone graft fillers. ZOI assays confirmed the antimicrobial activity of tobramycin after coating formulation and release, independent of the underlying graft substrate or coating method (see
Regardless of the polymer matrix formulation, micron-sized allograft bone particulate provided the most desirable release profile for tobramcyin. Derivative to its small size and high porosity, micronized allograft and/or synthetic graft may also provide a more efficacious wound packing material to prevent the formation of inadvertent avascular dead spaces, as opposed to larger porous fragments (see
Ideal antibiotic delivery systems would provide killing via a burst release (i.e., a bolus release) within the first 24-hour period, after administration followed by a sustained release above the minimal inhibitory concentration (MIC) to address the remaining microbial threat out to the 6-week time point (previously established by the orthopedic community as important to infection prevention) (Kanellakopoulou and Giamarellos-Bourboulis, supra; Chang et al., J Control Release 2006; 110: 414-21). Targeted antibiotic delivery sustained above the MIC minimizes the selective pressures of antibiotic resistant infections (Patzakis and Zalavras. J Am Acad Orthop Surg 2005; 13: 417-27; Strachan C J., J Antimicrob Chemother 1993; 31 SupplB: 65-78). However, current methods of simple physical absorption of drug to allograft bone can result in only a bolus release (Kanellakopoulou and Giamarellos-Bourboulis, supra), with no sustained release to combat further infections. Polymer-controlled antibiotic release from allograft bone material is a desirable alternative as it allows for long-term antibiotic success.
The use of a pharmaceutical-encapsulating, rate-controlling polymer membrane with defined degradation character to endow allograft bone with antimicrobial activity provides a level of delivery control unattainable with mere physical adsorption. Although polymer-controlled local drug delivery is not a new idea, the successful tailoring of the polymer coating to provide predictable drug release kinetics provides an important twist on a classic idea (Davidoff et al., Biomed Sci Instrum 2010; 46: 184-9; Davidoff et al., Biomed Sci Instrum 47:46-51, 2011; Sevy et al., supra). Previous studies have demonstrated changes in the release kinetic profile of tobramycin from this polymer-controlled drug delivery system based on 1) a change in the molecular weight of the polymer, 2) the incorporation of a water, non-solvent (Davidoff et al., Biomed Sci Instrum 2011; 47:46-51, 2011) or 3) the addition of an aqueous polyethylene glycol (PEG) solution (see Example 1 above) in the coating fabrication processes. The most significant and beneficial alterations in tobramycin release were noted when the miscibility of the drug and the polymer were changed as a function of solvent, most likely as a result of improved drug solubility, hydrophobicity, and charge. To investigate the effect of solubility and miscibility on drug-release profile, water-soluble antibiotics (vancomycin, oxacillin) and non-water soluble antibiotics (ciprofloxacin, rifampicin) were released from a polymer-controlled releasing membrane with and without the incorporation of a water, non-solvent into the formulation. Moreover, the danger of multi-drug resistant pathogens will be investigated by considering the combinatorial therapeutic efficacy of ciprofloxacin, rifampicin, and vancomycin against S. aureus as released from a local drug delivery system using a polymer-controlled releasing membrane coated onto allograft bone. Tailoring the release kinetics as a function of antibiotic solubility should provide clinicians with a long-term, antibiotic delivery system that can be customized and combined to fit each patient's needs while concurrently mitigating the development of antibiotic resistance.
By developing a local drug delivery system that controls the release of antibiotic via a degradable polycaprolactone (PCL) polymer membrane coated on an implantable allograft bone delivery vehicle, an increase in the bioactive longevity of antibiotic therapy is anticipated. Furthermore, combinatorial antibiotic experimentation will provide clinicians with an “ala carte” style of antibiotic treatment that can be tailored to meet each patient's needs while mitigating the development of bacterial resistance due to systemic and prolonged antibiotic overuse and poor patient compliance.
In this example, another non-limiting implant was fabricated and tested for its ability to diffuse drug for a prolonged amount of time. For this example, the following methods were used.
Fabrication of Antimicrobial Allograft Bone Fragments (AABF). Cancellous allograft bone fragments (Miami Tissue Bank) were weighed and similar weights were selected for each cohort. Each cohort was coated with a drug-releasing polymer to investigate release kinetics of different antibiotic-containing formulations (n=6). Each delivery system consisted of allograft bone coated with 18-22 mg of a 60 mg/mL polycaprolactone (PCL, Sigma CAS 24980-41-24, St. Louis, Mo., USA) solution containing 16 mg/mL antibiotic. PCL was resuspended in acetone at 45° C. Subsequently, antibiotic powder was added to PCL solution for cohorts without the water non-solvent. Alternatively, for cohorts containing water non-solvent, water was added to the PCL acetone solution at 4% (v/v) prior to adding the antibiotic powder. The cohorts differed according to 1) which of the antibiotics was incorporated (vancomycin HCl, rifampicin, ciprofloxacin, oxacillin, and ciprofloxacin HCl) and 2) the addition of 4% water, non-solvent to the system. Standards containing only one drug were evaluated as a baseline to determine combinatorial effects.
Release Studies. The incorporated antibiotic was subsequently released from each cohort via incubation in 5 mL of phosphate buffered saline (PBS, cat#BP661-10, Fisher Scientific) at 37° C. Release media was collected and completely exchanged at various time points between 24 hours and 8 weeks. Subsequently, 500 uL of each collected release media was dried in a concentrator and stored at 4° C. for microbial studies.
Scanning Electron Microscopy (SEM). Three fragments from each cohort were analyzed for surface characteristics using low vacuum SEM (FEI Quanta 3D dbFIB). Microanalysis system software displayed the real-time, back-scatter electron detector (BSED) images captured from the microscope and allowed the capture of images between 1 mm and 100 um magnification.
Kinetics Assay. Dried antibiotic release samples were reconstituted in 96-well plates with 100 μL of PBS. An absorbance assay (320 nm) was used to determine the relative concentration of antibiotic in each sample for ciprofloxacin (salt and free-base form), rifampicin, and vancomycin HCl. Plates with rifampicin or ciprofloxacin release media were read for absorbance (320 nm) using a microplate reader (BioTek spectrophotometer and Gen5 1.09 software). Vancomycin containing samples were tested using 2 uL of the reconstituted sample read in a BioTek Take3 quartz plate using the Gen5 1.09 software. Tobramycin sulfate samples were analyzed for comparison using a modified o-phthaldialdehyde (OPA)-based fluorescence assay (100 uL sample, 100 uL isopropanol, and 200 uL OPA reagent (Sigma P-05322)) and read on the Biotek microplate reader at 360 nm excitation and 460 nm emission (Sevy et al., Biomed Sci Instrum, vol. 46, pp. 136-41, 2010).
Zone of Inhibition. Immediately following the kinetics assay, filter paper disks (diameter=6 mm) were placed in the antibiotic release samples in the 96-well plate and dried overnight. Staphylococcus aureus (ATCC 25923 from American Type Culture Collection, Manassas, Va., USA) streaked Brain-Heart Infusion (BHI, BD cat#237500) agar plates were prepared for zone of inhibition experiments using pooled isolated colonies from blood agar plates diluted to a standard concentration (0.5 McFarland units). Subsequently, each drug-containing disk was placed on the bacterial-streaked agar 24 cm apart. Plates were incubated for 18 hours at 37° C. Electronic calipers were used to measure the cleared (no bacterial growth) diameter surrounding each disk (zone of inhibition). If no measurable zone was present the diameter was recorded as zero.
Combination Therapy. Ciprofloxacin, rifampicin, and vancomycin HCl were evaluated in combination (ciprofloxacin:rifampicin, rifampicin:vancomycin, ciprofloxacin:vancomycin) by combining 100 ul of release media (i.e., media released from the implant) for each time point and testing the bioactivity using zone of inhibition, as described above. 100 uL of uncombined release medias were used as controls. A student's 2-tailed t-test was used to determine if there was a significant increase in bioactivity when antibiotics were tested in combination.
In this Example, the following results were obtained.
Sample Fabrication. Polycaprolactone (PCL) (10 kD PCL, 60 mg/ml) polymer was dissolved in acetone at 45° C. Subsequently antibiotic (ciprofloxacin (salt and free-base forms), rifampicin, vancomycin HCl, oxacillin) and/or 4% water non-solvent was added to the solution to create the final coating formulation. Antibiotic polymer solutions were used to dip-coat cancellous allograft bone fragments (average dimensions 6 mm×5 mm×4.5 mm). Each antibiotic was added to two cohorts: one with a 4% water non-solvent in the formulation, and one without the water non-solvent component according to methods previously described (Davidoff et al., Biomed Sci Instrum, vol. 47: 46-51, 2011). The amount of antibiotic applied to each allograft was determined by using the weight of the applied coating and the antibiotic percent of the formulation. In past studies, tobramycin was used as the drug of choice due to its clinical relevance, high thermostability, and efficacy.
Scanning Electron Microscopy. In order to assess the PCL coating consistency and allograft surface coverage (i.e., to ensure a uniform mixture), the physical characteristics of the different surfaces were analyzed using SEM images (500 nM, 100 nM, and 20 nM). Images (500 nm) revealed that the polymer coatings of all antibiotics and coatings formulations with and without, water exhibit consistency and coverage (see the natural (allograft) croutons shown
Release kinetics. Antibiotics were released from the polymer-coated allograft bone into PBS. PBS was completely removed replenished at regular time points out to 8 weeks in order to simulate sink conditions. A kinetic release curve for each antibiotic was calculated using a standard curve of the absorbance for each antibiotic and normalizing to the amount of antibiotic theoretically applied to each crouton based on the weight of the coating and the amount of antibiotic added to the formulation (see
Table 2 is a chart showing drug and polymer solubility.
In Table 2, note the discrepancy between the solubilities of the drugs and the polymer. This difference has a dramatic effect on the miscibility of the polymer and the drug and forms the basis for the drug release method described here.
Based on differential polymer/drug solubilities, phase extraction methods were used to separate and isolate the antibiotic, polymer, and allograft components of the system. Drug load was subsequently quantified using optical absorbance or fluorescence assays. Methods were validated based on control samples without polymer or bone graft, revealing over 85% drug recovery over the linear range of the specific antibiotic assay. Ultimately, this information can be used to determine the mass balance after drug release so as not to incur some of the pitfalls of current antibiotic-releasing implants, sub-therapeutic drug dosing and development of antibiotic resistant pathogens.
While all samples demonstrated release profiles throughout the 8-week study duration, immediate differences were evident in the kinetics of release based on the type of antibiotic incorporated into the coating. Salt formulations of the antibiotics exhibited an initial bolus release of antibiotic, followed by a slow tapering of drug release at each time point; however, free-base (free-base) formulations of the drug show evidence of first order release kinetics (see
Antimicrobial bioactivity. Bioactivity of each sample was determined using zone of inhibition assays (
The variation in antibiotic release kinetic curves (first-order or zero-order) based on the salt form of the antibiotic prompted an investigation of the antibiotic benefits of combination therapies. The antimicrobial bioactivity of two drugs in combination was accessed using ZOIs. Interestingly, no statistically significant changes were observed in these studies, suggesting that combinatory therapies, while not deemed synergistic or additive, would not be antagonistic.
Salt vs. Free-Base Formulation Comparisons.
The impact of antibiotic form (salt or non-salt) was assessed by determining the antibiotic release kinetic profiles and antimicrobial bioactivity using PCL/acetone coating formulations that contained either ciprofloxacin salt or ciprofloxacin free-base. Release profiles and bioactivity showed significant differences (see
Thus, the formulation appears to affect the release kinetics of antibiotic from a polymer-membrane. Nosocomial osteomyelitis remains a significant clinical challenge associated with orthopedic surgeries due to a combination of bone's inherent avascularity, surgically compromised vascular supply, prevalent void spaces, and necrotic tissue. These factors not only limit the body's ability to combat opportunistic infections but also provide a favorable environment allowing invading microbes to evade the immune response, potentially coordinating the infection via biofilm development. The prevalent use of systemic antibiotics or antibiotic leaching orthopedic products has inadvertently promoted the development of antibiotic resistant microbes. Local antibiotic delivery offers a promising solution to combat these challenging opportunistic infections. During the course of this study, a polymer-coated allograft bone void filler was developed to release a variety of antibiotics locally, in a polymer controlled manner. The polymer coating and antibiotic constituents can be varied to provide either a bolus or sustained antibiotic release over the clinically relevant time frame of 8 weeks. More importantly, this technique has also allowed the development of a combinatorial approach to combating infection that may provide an important advance in reducing surgically introduced infections both acute and chronic following both primary and revision arthroplasty.
By utilizing a combinatory approach, i.e. the use of multiple drugs released from different fragments surgically implanted in a single void, surgeons can potentially minimize antimicrobial drug resistance, provide localized drug application, and provide protection against a broad spectrum of microbes. Over the past decade, drug resistance microbes have emerged as perhaps the greatest threat to surgical success (Peters et al., J Infect Dis 2008; 197: 1087-93). A major factor in drug resistance is leaching of drug at sub-minimal inhibitory concentrations for a sustained period of time. Combinatory approaches would enhance antibiotic protection by providing a more sustained drug release above the MIC while simultaneously decreasing the amount of either drug administered. Furthermore, the drug is localized to site of infection, allowing the host to maintain the native bacterial flora as well as minimize the nephro- and oto-toxic side effects common to many antibiotics during systemic administration.
Past studies utilizing tobramycin in this drug delivery system suggest that by mixing polymer and coating techniques this system can be engineered to not only minimize infections but also promote healing. A localized, sustained release out to 8 weeks provides a significant advancement in controlled drug release. In vivo assessment of this system will provide additional insight into the system's efficacy and practicality for clinical application.
In this example, the generation 2 fabrication was tested in vivo in mice.
For these studies, C57 Black mice were used, with 3-5 animals per group. Morselized allograft bone was coated by solvent casting in a PCL tobramycin solution (PCL—60 mg/ml 10 kD; 10% tobramycin). This bone implant was a generation 2 fabrication made using the method schematically depicted in
Mice were assessed post-implant for appearance and behavior using a modified Petty scale to assess these attributes (see Rao, N. et al., Plast. Reconstr. Surg. 127, Suppl. 1, 177S-187S, 2010). Representative individuals are shown in
Additionally, using a modified Petty scale, the implanted mice (n=9 mic total) were assessed for appearance and behavior up to forty days (i.e., almost 6 weeks) post implant. In this assessment, the higher the score, the more “unnatural” the animal appeared and behaved. As shown in
As described in the Examples above, addition of PEG to the implant helped with the phase separation problem and altered the kinetics of the drug release. However, PEG did not fully cure the phase separation problem. Thus, the generation 3 fabrication method was developed as a molten cast method. To generate approximately twenty solid implants having a uniform mixture (where each implant is 2 mm×2 mm×6 mm in size), the following protocol is used. This process is schematically depicted in
Additionally, although the below protocol uses the thermostable drug tobramycin, any other thermostable drug may be used. If the drug is an antibiotic, such non-limiting thermostable antibiotics include tobramycin, gentamicin, vancomycin, a cephalosporin, or a mixture of two or more of tobramycin, gentamicin, vancomycin, and a cephalosporin
For this protocol, the materials used were:
ProOsteon 500R (commercially available from Biomet, Inc., Warsaw, Ind.)
PCL 10 KD (commercially available from Sigma-Aldrich Co, St. Louis, Mo., catalog #: 440752)
PEG 20 KD (commercially available from Sigma-Aldrich Co, St. Louis, Mo., catalog #: P2263)
Acetone
Tobramycin (commercially available from Research Products International Corp., Mr. Prospect, Ill., catalog #: T45000-1.0) or Microencapsulated Tobramycin (Maxx Performance, Inc., Chester, N.Y.) or any other thermostable drug of interest
Other materials included weigh boats, Glass Petri Dishes, Spatulas (2×), Round Bottom Flask with a stir bar, Slide Molds, Silicone Isolators, Mortar and Pestle, water bath. −20° C. freezer, hot plate, and an external temperature probe.
To generate the implants, the following steps were taken: First, the Water Bath was heated to 45° C. and the hot plate is heated to 80° C. Next, the amount of polymer/drug/bone void filler needed was calculated. For example, if the bone component was about 64% ground, then 0.7 grams of morselized ProOsteon was used, 0.3 grams of PEG/PCL combination (all ratios by weight) was used, and 0.1 grams of Tobramycin was used. Next the BoneVoid Filler ProOsteon was morselized with mortar and pestle. The quality of morselization of the ProOsteon was evaluated under a dissecting microscope to ensure consistency of the particles. The morselized ProOsteon, Polymer mixture (i.e., PCL and PEG), and Tobramycin were then weighed out according to calculations determined above (i.e., 0.7 grams morselized ProOsteon, 0.3 grams PEG/PCL combination, and 0.1 grams of tobramycin). Note that the ratio of PCL and PEG was changed according to the desired degradation properties, but typically varies between 75% PCL and 25% PEG and 90% PCL and 10% PEG in relation to the polymer component of the formulation (all ratios by weight) So, for example, if 1 gram of total mixture was desired with 90% PCL and 10% PEG, then the final mixture (i.e., that was poured into the mold) contained 700 mg of bone, 270 mg of PCL, and 30 mg of PEG.
As depicted schematically in
To generate the implants, the polymer (i.e., the mixture of PEG and PCL) was put into the slide mold (e.g., as depicted in
As an optional step, while waiting for the polymer to heat for 15-30 minutes at 80° C., a PCL solution of 60 mg/mL in acetone was prepared. To do this, 10 kD PCL is added to acetone at a concentration of 60 mg/ml using the 45° C. water bath with stirring.
Next, morselized ProOsteon and tobramycin were added to the melted polymer mixture. The resulting polymer/ProOsteon/tobramycin mixture was mixed well, especially in the corners of the slide mold.
Next, the silicone isolator (e.g., such as one depicted in
Using the spatulas, each space in the mold with the silicone isolator was filled with the polymer/ProOsteon/drug molten mixture and compress. Excess polymer was scraped away before it solidified.
After silicone isolator was filled, the mold is placed in the freezer (−20° C.) for at least 5 minutes.
Then, the mold was removed from freezer and excess polymer/ProOsteon/drug was scraped away from the silicone isolators.
The isolator was then peeled off the foil and the implants pushed out.
If the PCL solution of 60 mg/ml in acetone was made earlier, as an optional step, each implant may be dipped in this solution. This optional step may create a “sealing” coat. In other words, in this embodiment, the resulting implant with the uniform mixture of polymer/ProOsteon/drug is additionally coated with a PCL coat. To do this, each implant was dipped in the PCL/acetone solution for about 30 seconds, and then allowed to dry for approximately 2 minutes. The implant is turned over and dipped again for about 30 seconds, and then allowed to dry for about 2 minutes. The dipping/drying process was repeated three or more times.
Each implant was next tested for quality assurance.
To do this, the bone void filler was weighted (+/−5%). Generally, a 2 mm by 2 mm by 6 mm implant had a weight of 37.5+/−1.875 (5%) milligrams.
Also, the length, width, and height were measured (+/−5%). A 2 mm by 2 mm by 6 mm length should be +/−0.05 mm for width and height and 0.15 mm for length.
Also, under a dissecting microscope, smoothness is looked for. In some embodiments, the resulting implant does not have major voids. Similarly, under a dissecting microscope squareness is looked for. In some embodiments, the resulting implant has crisp 90° angles.
From every batch, an SEM (scanning electron microscopy) was taken of one implant to ensure consistency of blended material (i.e., to ensure a uniform mixture in the solid implant)
Also, from every batch, an SEM was taken of one implant after 1 day release to ensure porosity of sample. Note sample porosity is created when the PEG dissolves from the implant.
Compression and cyclic compression mechanical tests by applying pressure to each of the dimensions of the sample using an Instron testing system with BlueHill software is also performed on at least one implant of every batch to determine isotropy. It should be noted that the actual amount of antibiotic in each bone implant made using any method described herein (including, without limitation, the method described in this Example 3) can be detected by standard methods. For example, as shown in
In this example, the generation 3 fabrication generated according to the methods described in Example 4 is used in vivo in rabbits.
First, a compression test was performed to look at the strength of the bone implants with or without drug. A 1% compression is typical in bone cyclical tests because it approximates the amount of strain during walking. The results of these studies are shown in
As shown in
The shelf-life of the generation 3 fabrication was next confirmed. The generation 3 fabrications were stored at four different temperatures (i.e., −20° C., 4° C., 25° C., and 55° C.), and then at specific time points, a designated number of implants were removed from storage and subjected to scanning electron microscope (SEM) imaging and mechanical testing.
To confirm bioactivity of generation 3 fabrications based on a difference in the ratio of PCL to PEG, ZOI data was collected. The fabrications tested were 95% PCL: 5% PEG ratio by weight in the polymer portion of the fabrication which is 27% of the total formulation (the other components being 10% drug and 63% bone) (blue bars in
Next, an experiment was performed to determine if bacteria could still grow in the presence of liquid released from a generation 3 fabrication bone implant. It was thought that if bacteria didn't grow in the liquid media, but a sample of that media grows on solid agar nutrients, the bacteria may have still been alive, meaning the drug was not bacteriacidal (i.e., killing bacteria) but merely stopped bacteria from growing (bacteriostatic). For these studies, each generation 3 implant was placed in phosphate buffered saline (PBS) and the PBS collected at different time points. The PBS was then added to a bacterial culture and the growth of the culture was determined using absorbance at. Subsequently, bacteria from the culture were plated on blood agar plates and the resulting colonies counted.
The results showed that molecules leaching from the polymer coating of the generation 3 fabrication bone implant seemed to have a bacteriostatic effect. Moreover, if a dosage of at least 107 CFU (colony forming unit) were added, the bacteria could grow. The implication is that if the generation 3 fabrication can block active bacteria growth (i.e., hold the bacteria to at least 105 CFU, then the rabbits' immune systems can overcome the infection. The results of this study are shown in Table 5.
Interestingly, the polymer itself was found to have bacteriostatic effects, but those effects could be overcome by using a higher initial amount of bacteria. As shown in
The effect of the generation 3 fabrication on osteoblasts was next tested. For these studies, osteoblasts grown in vitro were placed in the presence of the generation 3 fabrication itself, tobramycin-soaked generation 3 fabrication, or untreated control ProOsteon. As shown in
Additional references for this work include:
1) Fátima Varanda et al, Solubility of Antibiotics in Different Solvents. 1. Hydrochloride Forms of Tetracycline, Moxifloxacin, and Ciprofloxacin, Jul. 29, 2006
2) For solubilities: http://www.pharmacopeia.cn/usp.asp
3) Woodruff, Maria A. and Hutmacher, Dietmar W. (2010) The return of a forgotten polymer: Polycaprolactone in the 21st century. Progress in Polymer Science
For the in vivo studies, nine male, 3-4 kg New Zealand White rabbits (between 1-2 years of age) were used.
In the rabbit experiment, 9 rabbits were implanted with either the non-limiting generation 3 bone implant described herein that was not coated or the non-limiting bone implant described herein that was coated with an antibiotic-containing polymer coating. All implants were sterilized with ethylene oxide (EtO) prior to implantation
These timeline of these studies is schematically set forth in
For these studies, prior to surgery, the right forelimb of the rabbit leg was clipped, as was a small patch at the back of the head on the back, and 25 mcg Durageic (fentanyl) patch was placed on the back area. The surgical procedure was adapted from Smelter's and Koort's protocols (see Koort et al., Antimicrob Agents Chemother, vol. 49, pp. 1502-8, 2005; Koort et al., J Biomed Mater Res A, vol. 78, pp. 532-40, 2006; Smeltzer et al., J Orthop Res, vol. 15, pp. 414-21, 1997). Briefly, the radius of the right forelimb was exposed surgically and prepared by scrubbing with Povidine iodine and ethanol solution. A bone segment (approx. 6 mm by 2.7 mm by 2 mm) was drilled under saline cooling into the proximal medial metaphysis of the right tibia. Subsequently, 105 to 107 Colony Forming Unites (CFUs) of S. aureus (for the rabbits) were injected directly into the medullary canal anterior to the surgical site. The bone segment was then filled with a non-limiting bone implant containing 90% PCL:10% PEG prepared as described above in Example 5, or with a suitable control bone graft replacement (approx. 6 mm by 2 mm by 2 mm). The wound was closed with resorbable sutures and the rabbit removed from the anesthetic and observed until it was awake and mobile. The animals were dosed with 100 mg of Cefazolin SQ injection 30 minutes post-surgery. Postoperative analgesics (Duragesic/Fentanyl) were administered to all animals immediately after surgery and transdermally for 2 days at a concentration of 25 mcg/hr. Analgesics were continued with any animal that avoided use of the affected forelimb. Pain levels (e.g., lack of appetite, shivering, and postural changes) were monitored for a minimum of 3 days postoperatively. The temperature and weight of each animal were monitored weekly as well as radiographic images of the surgical site. Note that the animal was terminated if the following conditions occurred: blood borne infection, overwhelming local infection, excessive signs of distress, appetite suppression as indicated by loss of weight, limited water consumption, and/or lethargy.
For these studies in rabbits, endpoint analyses included: (i) imaging of bone by X-Ray; (ii) microbiological culture of bone site and soft tissue surrounding surgical site; (iii) SEM and histological analyses of bone growth and (iv) high pressure liquid chromatography (HPLC) quantification of antibiotic excreted in the urine or still remaining in bone replacement at the conclusion of the study or termination of the animal.
The different cohorts of animals used in this study were as follows:
Cohort 1: No polymer, no drug, no infection. The implant used in this cohort 1 was a fragment of ProOsteon that was sculpted with a razor blade to be 2 mm×2 mm×6 mm in dimensions. There was no polymer and no drug used to fabricate the implant, and no infection was introduced into the surgical site. This cohort 1 was used as a control that allowed an assessment of the surgical technique and the sterility conditions.
Cohort 2: No polymer, no drug, 105 CFU S. aureus. The implant used in this cohort 2 was a fragment of ProOsteon that was sculpted with a razor blade to be 2 mm×2 mm×6 mm in dimensions. There is no polymer and no drug used to fabricate the implant. 105 CFU of S. aureus was introduced into the medullary canal anterior to the surgical site on the tibia. This cohort 2 was a control that allowed an assessment of the surgical technique and the sterility conditions.
Cohort 3: PCL-PEG coat, no drug, no infection. The implant used in this cohort 3 was morselized ProOsteon that was mixed using the generation 3 fabrication method in a ratio of 70% ProOsteon and 30% polymer. The polymer was melted at 75° C. in a ratio of 90% PCL and 10% PEG. The mixture was then packed into the silicone isolator (dimensions of 2 mm×2 mm×6 mm). A final dip of each fabricated crouton into a PCL acetone solution (10 kD PCL at 60 mg/ml in acetone) was done prior to sterilization. For this cohort 3, there was no drug used in fabricating the implant, and no infection introduced into the surgical site (i.e., the implantation site). This cohort 3 was a control that allowed an assessment of the safety of the polymer components of the generation 3 fabrication. Note that no data is shown from this control cohort 3 as it was unremarkable and looked like the results from cohort 1.
Cohort 4: PCL-PEG coat, no drug, 105 CFU S. aureus. The implant used in this cohort 4 was morselized ProOsteon that was mixed using the generation 3 fabrication method in a ratio of 70% ProOsteon and 30% polymer. The polymer was melted at 75° C. in a ratio of 90% PCL and 10% PEG. The mixture was then packed into the silicone isolator (dimensions of 2 mm×2 mm×6 mm). There was no drug used to make the implant, but 105CFU of S. aureus was injected into the medullary canal anterior to the surgical site. This cohort 4 is a control that allows an assessment of the impact of the polymer on the progression of the infection.
Cohort 5: No polymer, 10% drug soak, 105 CFU S. aureus. The implant used in this cohort 5 was a fragment of ProOsteon that was sculpted with a razor blade to be 2 mm×2 mm×6 mm in dimensions. There is no polymer in the implant, but the implant was soaked in a 10% solution of tobramycin in water for 10 minutes prior to implantation and 105 CFU of S. aureus was introduced into the medullary canal anterior to the surgical site on the tibia. This cohort 5 was a control that mimics what is currently being done in many human surgeries.
Cohort 6: PCL-PEG coat, 10% drug load, no infection. The implant used in this cohort 6 was morselized ProOsteon that was mixed using the generation 3 fabrication method in a ratio of 63% ProOsteon and 27% polymer. The polymer was melted at 75° C. in a ratio of 90% PCL and 10% PEG. 10% powdered tobramycin drug was added to this molten mixture of polymer and ProOsteon. The mixture was then packed into the silicone isolator (dimensions of 2 mm×2 mm×6 mm). A final dip of each fabricated crouton into a PCL acetone solution (10 kD PCL at 60 mg/ml in acetone) was done prior to sterilization. No infection was introduced into the surgical site. This cohort 6 was a control that allowed observation of how host bone reacted to the generation formulation for safety purposes.
Cohort 7: PCL-PEG coat, 10% drug load, 105 CFU S. aureus. The implant used in this cohort 7 was morselized ProOsteon that was mixed using the generation 3 fabrication method in a ratio of 63% ProOsteon and 27% polymer. The polymer was melted at 75° C. in a ratio of 90% PCL and 10% PEG. 10% powdered tobramycin drug was added to this molten mixture of polymer and ProOsteon. The mixture was then packed into the silicone isolator (dimensions of 2 mm×2 mm×6 mm). A final dip of each fabricated crouton into a PCL acetone solution (10 kD PCL at 60 mg/ml in acetone) was done prior to sterilization. 105 CFU of S. aureus introduced into the medullary canal anterior to the surgical implantation site on the tibia.
The surgeries were successful with no systemic infection seen in any of the nine animals. The implanted grafts provided mechanical stability.
Using a high performance liquid chromatography (HPLC) protocol described above, the presence of tobramycin in rabbit urine was able to be traced following implantation of the tobramycin-soaked generation 3 fabrication. The data in
Next, photographs were taken of representative rabbits. As shown in
Radiographic analysis was also performed.
Histology analysis of the animals from cohort 6 shows that the PCL-PEG polymer coated, drug-soaked implant results in the formation of bridging calleous (“callus formation”, a typical bone healing response) as indicated by the dark bone implant being completed enveloped by red host tissue at 8 weeks (left in
Upon euthanasia and after dissecting the right forelimb from the rabbit, the bone was fixed in 10% buffered formalin and embedded in polymethylmethacrylate (PMMA). A 10 micron section of bone was then processed for histological gram stain.
Table 6 provides a summary of the findings from this study.
As Table 6 shows, bacteria were found to be present in both bone and soft tissue in cohort 4.
Interestingly, histological analysis of the implants from this study at the end of the study did not show the complete re-sorption of the bone by host (i.e., host bone did not completely take over the implant). This lack of osteoconduction (i.e., bone in-growth) is believed to be due to a lack within the implant of pores that are contiguously connected throughout the implant to allow penetration into the implant of host cells (e.g., osteoblasts and osteoclasts). More specifically, bone is porous, but finely ground bone necessarily loses some of the porosity present in bone before it is ground. As discussed above (see, e.g.,
To overcome the lack of contiguous porosity (i.e., lack of interconnected pores) within the implant, in another embodiment, an additional formulation comprising a polymer component of a PCL:PEG: poly(lactide-co-glycolide) combination; a bone component of ground bone (e.g., natural or synthetic bone such as ProOsteon); and a drug component (e.g., tobramycin) is employed. In yet another embodiment, a formulation comprising a polymer component of a PCL:PEG: poly(lactide-co-glycolide) combination including a poragen such as calcium chloride; a bone component of ground synthetic bone such as ProOsteon; and a drug component of tobramycin is employed. Of course the ground synthetic bone can be replaced with ground natural bone (or synthetic bone from other sources) and the drug can be a drug other than tobramycin. Yet another alternative is to apply an implant as a liquid paste (see, e.g., Example 7 below).
A tibia osteomyelitis model in sheep is next performed at a GLP facility. Since sheep bone is similar to human bone, sheep are studied. For these studies, eleven sheep will be used per study group, and 105-107 S. aureus will be used to infect the sheep at the implantation site.
The results in the sheep will show that implantation of a generation 3 fabrication bone implant that fabricated with tobramycin (i.e., the tobramycin was loaded uniformly throughout the implant during the generation 3 fabrication method as per
Taken together, these results show that the bone implant that was loaded tobramycin (i.e., the tobramycin is loaded uniformly throughout the implant) and then coated with an antibiotic-containing polymer coating is superior in stopping and preventing infection for a prolonged period of time. Table 7 provides a summary of the results described herein.
In this Example, an injectable bone paste using low molecular weight PCL and PEG, with the bone ceramic composite mixture at more than 50 weight % bone solid particles, and drug mixture is described.
This injectable bone paste will be fabricated as follows. 60% morselized ProOsteon will be mixed with 30% polymer (PCL (<3 kD) and PEG (<lkD) in ratios of 75-99% and 1-25% respectively) at 75° C. to create a molten paste. Up to 10% powdered tobramycin will be added to the molten mixture at which time it will be packed into a 1-5 ml syringe for sterilization. This paste could then be injected directly at the site of injury through a large gauge needle (e.g., 18-22 gauge). This formulation is anticipated to slightly harden in situ, but remain fairly viscous allowing it to remain at the site of injury and release its drug as the binding polymer matrix is degraded over time. With this low molecular weight polymer, this injectable paste may not have the same length of antibiotic release as the solid fabrications described herein (e.g., generated using the generation 2 or generation 3 fabrication method). Rather, the injectable paste fabrication may have an antibiotic release time of about 4-6 weeks.
It is expected that histological analysis of this implant after six weeks will reveal resorption of the implant by host (i.e., implant replaced by host bone).
In this example, the implant described herein is used in conjunction with a prosthesis.
Annual incidence of infections to orthopedic implants in the United States is substantial: 12,000 total joint infections and over 100,000 infected bone fixation implants annually. This produces a substantial cost both in terms of patient morbidity and financial coverage of these infections. All medical device-related infections are estimated to cost $1.7-4.6 billion in excess medical costs to U.S. hospitals annually. If 20% of these device-related infections could be prevented, $300-900 million in medical costs and much patient morbidity, pain and suffering could be saved.
Cement-Less Biological Bone Fixation and Implant Porosity.
Cement-less fixation represents an alternative method to popular acrylic bone cements to place and stabilize metal implants in bone. The method intends to stabilize metal implants using the patient's own direct bone-implant on-growth, on-bonding between bone and implant surface, and mechanical fixation from this interaction. The method, used in various forms since the 1980's, is intended to surpass cemented implant fixation as the method of the future—PMMA and standard thermoset cement technology will be eventually passed over in favor of cementless implant-bone bonding relying on direct bone-implant bonding. Typically, cementless fixation has been produced by host bone in-growth into carefully designed and fabricated implant pores of sufficiently large size. Porosity is critical to promote and produce this bone on-bonding fixation process with an implant. When new bone from the patient calcifies within these pores, this allows mechanical interlocking and stabilization of the bone-implant interface, eliminating the need for acrylic cements. Importantly, the ideal pore size should mimic that of native cancellous bone that ranges from 400-500 microns (dense cortical bone by comparison is only 8% porous). By contrast, most porous metallic implants (e.g., commercially pure (CP) titanium, cobalt-chrome alloys, Ti-6Al-4V alloy) have pores ranging from 100-400 microns, with 30-50% total porosity. Proper implant pores sizes and pore densities prompt enhanced bone-based fixation, achieved earlier than using fixation with allograft cortical bone, in some cases a matter of weeks.
Zimmer (Warsaw, Ind., USA) introduced their Trabecular Metal™ implants fabricated of elemental tantalum metal (a rare and highly corrosion resistant metal applied to dental implants since the 1950s) using a vapor deposition technique to create a metallic strut configuration that is similar to trabecular bone architecture. The crystalline micro- and nano-texture of a Trabecular Metal strut is conducive to direct bone apposition. Furthermore, implants fabricated from tantalum offer high porosity, allowing not only bone around implant sites to grow onto the material but also into it—a process known as osseo-incorporation or biological fixation. Zimmer's trabecular metal implants are 70-80% porous, similar to cancellous bone. Studies on dental implants containing Trabecular Metal in canine mandibular models began in 2010 and showed evidence of in-growth by maturing bone as early as two weeks after implantation. Moreover, transcortical animal implant studies have demonstrated excellent new bone in-growth of Zimmer's Trabecular Metal implants within eight weeks of surgery, promoting rapid fixation strength. According to the Zimmer company reports, human trials data are currently being collected with the first long-term results expected to be available in 2012. Zimmer has gained CE approval for another dental Trabecular Metal implant in Europe in 2011 and anticipates market approval for the USA through the Food and Drug Administration soon. Trabecular Metal has been already used fbr more than a decade in many of Zimmer's orthopedic devices.
Infection and Cementless Fixation and Porosity.
Although cementless fixation via porous metallic implants continues to provide mechanical integrity, there is an increased long-term risk of revision due to infection in hybrid and cemented implants compared to uncemented implants as evidenced by several clinical studies on total hip arthroplasties. There are considerable indications that cementing produces substantial infection risk, even acting as a nidus of infection. Hence, acute infection rates in this host post-implantation are as significant in cementless and cemented fixation in studies reported to date. Significantly, infection serves to inhibit bone generation and on-growth, limiting implant stabilization by bone growth. Limited (and clinically preliminary) evidence for more chronic infections indicates that cementless fixation infection incidence longer-term is less than for cemented. Cemented fixation notably addresses infection risk with antibiotic-containing cements, but these suffer from low fractional release and low antimicrobial capabilities long term. The presence of the cement may act as a foreign body, enhancing rates of infection after antibiotic release is exhausted after a few days post-implantation. A significant unmet clinical need and opportunity exists currently in addressing infection risk in cementless fixation with a resorbable polymer/granular bone/drug composite coating or press-fit wafer that accommodates cementless fixation while mitigating infection risk short-term. This allows bone regeneration without infection.
Porous metal fixation designs on implants represent a known clinical infection risk. The methods and compositions described herein as applied to cementless fixation implants may mitigate this risk for the following reasons. First, the antibiotic eluting implants described herein are composite polymer-bone graft-drug matrices. Also, the implants described herein can be patterned onto (e.g., coated onto) metallic implants to produce local high-resolution zones of antibiotic-release on or adjacent to cementless porous metal areas (see
Thus, the methods and compositions described herein provide an on-board controlled drug delivery antimicrobial solution to infection in cementless fixation and microporous metals used in orthopedic and dental implant applications. The methods and compositions described herein provide a versatile device formula which contains resorbable, clinically familiar polymers, synthetic or allograft granular bone graft materials and clinically approved drugs. This formula can be applied by spray, high-resolution patterned inkjet, molding, pre-fabrication or dip coating methods locally in resolved spatial locations on device surfaces. The implant formulations provided herein can also be shape-molded specifically for press fitting into defects or pre-designed groove sites on metallic implants.
In some embodiments, the non-limiting implants of the invention enable desired drug-graft material interaction including: 1. Molding of the bone graft composite material to specific dimensions and sizes, with a known, reliable drug load, 2. A capability to carve and shape the graft to fit specific defect sites, and 3. extended control over drug release for long time periods and subsequent antimicrobial protection throughout the duration of bone remodeling as evidenced by preclinical studies in a rabbit radial critical size defect infection model. Ultimately while releasing bactericidal concentrations of tobramycin, this antibiotic-loaded bone graft provides recognized beneficial osteoconductive potential, seeking to decrease orthopedic surgical infection incidence with improved filling of dead space and more reliable new bone formation.
Additional references include the following, all of which are incorporated herein by reference in their entireties.
The embodiments of the invention described above are intended to be merely exemplary; numerous variations and modifications will be apparent to those skilled in the art. All such variations and modifications are intended to be within the scope of the present invention as defined in any appended claims.
This patent application claims benefit of U.S. provisional application Ser. No. 61/616,937, filed Mar. 28, 2012 and U.S. provisional application Ser. No. 61/595,544, filed Feb. 6, 2012, the entireties of both applications are hereby incorporated by reference.
| Number | Date | Country | |
|---|---|---|---|
| 61595544 | Feb 2012 | US | |
| 61616937 | Mar 2012 | US |
