Patent Application
20190142993
The present invention relates generally to strong, small-diameter sutures for controlled drug delivery, and more particularly, to fibers that can twist into high-strength sutures or coat sutures.
Development of drug-eluting sutures is of significant interest for a variety of clinical applications. Sutures are already used to close wounds or hold tissue together. Delivering active agents at the same time would promote healing and prevent complications (Casalini, T., et al., International Journal of Pharmaceutics, 429, 148-157 (2012)). The problem with many sutures for drug delivery is that incorporation of agent directly into the suture decreases strength or increases diameter if the agent is incorporated into a coating. (Weldon, C. B., et al., J Control Release, 161, 903-909 (2012)).
Eye infections such as bacterial keratitis and endophthalmitis can lead to significant negative consequences including corneal ulceration, edema, inflammation, and blindness (Lee B J, J Cataract Refract Surg, 35, 939-942 (2009)). Conventional nylon sutures used in ocular procedures can harbor bacteria and potentially facilitate infection (Katz, S., et al., Ann Surg, 194, 35-41 (1981); Leaper, D., et al., Ann R Coll Surg Engl, 92, 453-458 (2010)). This phenomenon is worsened when sutures become loose or break in situ. Almost 40% of loose or broken nylon corneal sutures are contaminated with bacteria, and Staphylococcus Epidermidis is isolated in more than 80% of cases (Heaven, C. J., Eye (Lond), 9 (Pt 1), 116-118 (1995)). It has become routine to prescribe expensive antibiotic drops off-label for prophylactic use after ophthalmic surgery; however, patients have low compliance using topical eye drops. Properly instilling eye drops is particularly difficult for pediatric patients and for those who are elderly and/or in cognitive decline (Winfield, A. J., British Journal of Ophthalmology, 74, 477-480 (1990); Burns, E., et al., Age and Ageing, 21, 168-170 (1992)). Up to 50% of patients take less than half of the prescribed doses over the course of a study on topical antibiotic eye drop compliance (Hermann, M. M., et al., Investigative Ophthalmology & Visual Science, 46, 3832-3832 (2005)). Lack of compliance may lead to re-occurrence of an infection or the development of antibiotic resistance (Bremond-Gignac, D., et al., Ophthalmol Eye Dis, 3, 29-43 (2011)).
Ophthalmic sutures are commonly used during ophthalmic surgical procedures, including trabeculectomy as well as pterygium removal, cataract surgery, strabismus correction surgery, penetrating keratoplasty, sclerectomy, and conjunctival closure. The choice of suture material can strongly impact the occurrence of complications related to infection and inflammation post ophthalmic surgery, either causing irritation and local inflammation or providing a substrate for microorganism growth. The suture materials typically employed include non-biodegradable ophthalmic suture materials such as ETHILON® nylon suture, MERSILENE® polyester fiber suture, PERMA-HAND® silk suture, PROLENE® polypropylene suture, each commercially available from Ethicon, Somerville, N.J.; and VASCUFIL® coated monofilament suture composed of a copolymer of butylene terephthalate and polyteramethylene ether glycol, MONOSOF˜DERMALON® monofilament nylon sutures composed of long-chain aliphatic polymers Nylon 6 and Nylon 6.6, NOVAFIL® monofilament sutures composed of a copolymer of butylene terephthalate and polyteramethylene ether glycol, SOFSILK® braided sutures composed of fibroin, TI-CRON-SURGIDAC® braided polyester sutures composed of polyester terephthalate, SURGILON® braided nylon sutures composed of the long-chain aliphatic polymers Nylon 6 and Nylon 6.6 and SURGIPRO II-SURGIPRO® sutures composed of polypropylene, each commercially available from U.S. Surgical, Norwalk, Conn.
Ideally, ophthalmic suture materials are biodegradable and biodegradable over the useful suture lifetime, retaining the requisite tensile strength and capable of delivering therapeutic or prophylactic agents to increase patient success. For pterygium removal, cataract surgery and strabismus correction surgery, sutures could be used to close the wound and release antibiotic and anti-inflammatory drugs. For trabeculectomy surgeries, sutures could be placed on sclera flaps providing local chemotherapeutic agents, decreasing production of scar tissue, and on conjunctival closure with antibiotic release. In penetrating keratoplasty, the sutures hold the graft, as well as release antibiotic and immunosuppressant or anti-inflammatory agents.
An alternative to frequent topical applications would be to supply antibiotics directly from the surgical suture. For this purpose, the suture needs to (i) be of suitable size, (ii) be of high-strength to resist breakage and bacterial colonization, and (iii) supply an effective amount of antibiotic.
However, clinical implementation of sutures has been limited due to the inability for drug-loaded sutures to meet United States Pharmacopeia (U.S.P.) standards for suture strength (Pruitt L A, et al., MRS Bulletin, 37, 698 (2012); Kashiwabuchi F, et al., Translational Vision Science & Technology, 6, 1 (2017)). Conventional suture manufacturing processes are not compatible with most therapeutic moieties, and drug-loaded sutures in preclinical development have demonstrated breaking strengths up to ten folds less than the strength required for clinical use (Hu W, et al., Nanotechnology, 21, 315104 (2010); Padmakumar S, et al., ACS Applied Materials & Interfaces, 8, 6925 (2016)). Attempts to develop drug-eluting sutures have been limited by lack of sufficient tensile strength (especially with the inclusion of drugs), poorly sustained drug release, or lack of scale needed for commercial viability (Wen Hu, et al., Nanotechnology, 21, 1-11 (2010); Pasternak, B., et al., Int J Colorectal Dis, 23, 271-276 (2008); Obermeier, A., et al., PLoS One, 9, e101426 (2014); Morizumi, S., et al., Journal of the American College of Cardiology, 58, 441-442 (2011); Mack, B. C., et al., J Control Release, 139, 205-211 (2009); Lee, J. E., et al., Acta Biomater, 9, 8318-8327 (2013); Joseph, J., et al., Nano Letters, 15, 5420-5426 (2015); Hu, W., et al., Nanotechnology, 21, 315104 (2010); Hu, W., et al., Society of Chemical Industry, 59, 92-99 (2010); He, C. L., et al., J Biomed Mater Res A, 89, 80-95 (2009); Catanzanoa, O., et al., Materials Science and Engineering: C, 43, 300-309 (2014); Choudhury, A. J., et al., Surgery, doi: 10.1016/j.surg.2015.07.022. Epub Aug. 29 (2015)).
Although nylon sutures are used in more than 12 million procedures per year globally to close ocular wounds and incisions, no drug-eluting sutures have been approved for ophthalmic use (Kronenthal R, P., et al., Sutures Materials in Cataract Surgery, (1984); Grinstaff, M. W., Biomaterials, 28, 5205-5214 (2007)). In 2002, Ethicon received approval to market a series of antibiotic-coated sutures. However, they are only available in sizes #1-0-#6-0, according to the United States Pharmacopeia (U.S.P.), and none were indicated for ophthalmic use. Ophthalmic sutures require sizes #8-0-#10-0 or thinner, and to date, no market offering is available for antibiotic-eluting sutures for ocular surgery (Marco F, et al., Surg Infect (Larchmt), 8(3), 359-365 (2007); Ming, X., et al., Surg Infect (Larchmt), 8, 209-214 (2007); Ming, X., et al., Surg Infect (Larchmt), 8, 201-208 (2007); Ming X, et al., Surg Infect (Larchmt), 9, 451-458 (2008)). Certain electrospun fibers were developed with a capability of drug loading (US 2013-0296933), but it is unclear how to maintain a tensile strength that satisfies the clinical strength requirement for sutures and is not compromised with drug loading. Due to this challenge drug-eluting sutures have been limited to drug-eluting coatings. While this method does not affect suture strength, it limits the amount of drug that can be included and results in rapid drug release as opposed to the sustained drug release needed for clinical applications outside of anti-infection uses. It is further desired to develop sutures capable of releasing any other type of drug for any other surgical application at any size.
Therefore, it is an object of the present invention to provide ultra-thin (small diameter), high strength multifilament composite fibers that are capable of eluting drugs in a controlled manner, without decreasing the tensile strength.
It is another object of the present invention to provide active agent-eluting sutures as substitutes for nylon (permanent) or VICRYL® (absorbable) sutures used in ophthalmic surgeries such as cataract, corneal transplant, injury, or in other specialty surgeries, supporting additional therapeutic functionality.
It is yet another object of the present invention to provide controlled coatings on existing sutures to add functionality such as elution of one or multiple drugs.
Suture materials or suture coating materials made from twisted, biocompatible polymeric fibers with high tensile strength for use in surgical repair and drug delivery have been developed. A plurality of fibers (e.g., electrospun fibers) is twisted or braided in a bundle to form a multifilament suture, or twisted or braided around a thread or suture to form a multifilament coating. The twisting increases the tensile strength of the overall multifilament composite, even in the presence of one or more active agents entrapped in the fibers. Orientation of polymer chains through molecular confinement, thus forming nanostructures, also enhances polymer crystallinity and strength. While a mono-filament fiber of a certain diameter loses its tensile strength significantly with the inclusion of therapeutic, prophylactic or diagnostic agents (e.g., 8 wt % levofloxacin), twisting of multiple (e.g., hundreds) fibers containing the drug into a multifilament composite of a similar diameter precludes the loss in tensile strength normally associated with drug loading. Additional twisting of these fibers serves to further increase the strength of the multifilament composite.
The fibers can be micro-fibers or nano-fibers. The twisted multifilament composite can have a diameter of less than 50 μm, less than 40 μm, and preferably less than 30 μm. For a multifilament composite having a diameter between 20 μm and 29 μm and optionally containing an active agent, the tensile strength of the composite should be greater than 0.24 N while satisfying the size and strength requirements for a #10-0 suture according to United States Pharmacopeia. For a multifilament composite having a diameter between 30 μm and 39 μm and optionally containing an active agent, the tensile strength of the composite should be greater than 0.49 N and the composite satisfies the size and strength requirements for a #9-0 suture according to United States Phaimacopeia. For a multifilament composite having a diameter between 40 μm and 49 μm and optionally containing an active agent, the tensile strength of the composite should be greater than 0.69 N and the composite satisfies the size and strength requirements for a #8-0 suture according to United States Pharmacopeia.
The multifilament sutures generally maintain at least about 95%, 90%, 85%, or 80% of their mechanical properties (e.g., breaking strength) even after immersion in an aqueous environment for about 1 week, 2 weeks, 30 days, 45 days, or greater. One or more therapeutic, prophylactic and/or diagnostic agents can be included, up to about 24 wt % or greater in the multifilament composite as a drug-releasing suture without compromising the tensile strength as required by United States Pharmacopeia. The multifilament composites generally have a diameter suitable for ophthalmic suturing procedures. The multifilament composites can also be twisted around a thread to provide drug-elution functionality, where the overall size of the coated suture still satisfies the needs for ophthalmic suturing procedures.
Exemplary suture formulation includes multi-nanofiber filaments made from polymers such as polyhydroxy acids such as poly(lactic-co-glycolic acid), polylactide, and polyglycolide, polydioxanone, polycaprolactone, or a copolymer, blend, or mixture thereof. A preferred suture formulation is made from degradable, drug-loaded polymeric multifilament twisted fibers, surpassing U.S.P. specifications for suture strengths. The suture may be of variable sizes from 2-0 to 10-0 U.S.P. specifications, based on the parameters in operating fabrication techniques (e.g., electrospinning) and the twisting and/or braiding parameters of filaments. The suture may be formulated to degrade in vivo over a time period from a few days to a few years. Suitable solvents for the polymers include chloroform, methanol, acetone, hexafluoroisopropanol, or other solvent depending on the solubility of specific polymers. In one embodiment, the suture materials and the coating materials for sutures are made from a polyhydroxy acid such as polylactide, polyglycolid, or a copolymer thereof or polycaprolactone, and optionally a polyalkylene oxide such as poly(ethylene glycol) or a polyalkylene oxide block copolymer. The sutures entrap (e.g., encapsulate) one or more therapeutic, prophylactic or diagnostic agents and provide prolonged release over a period of at least a week, preferably a month.
In some embodiments, s semi-crystalline, hydrophobic, degradable polymer, polycaprolactone (PCL) is used. Its fiber crystallinity, and therefore suture tensile strength, is maximized through the nanofiber fabrication process of low molecular weight PCL and subsequent twisting to form single sutures with additional compaction and structural reinforcement. Electro spinning may alter the molecular orientation of PCL to improve crystallinity. The molecular confinement may contribute to the increase in tensile strength of PCL nanofibers even with reduced diameter.
The degradable, multifilament sutures meet U.S.P. specifications for size and strength suitable for ophthalmic use, and surpass breaking strength specifications when loaded with a wide range of antibiotics of different physicochemical properties. Unlike micron-sized, electrospun PCL monofilament sutures which lose more than 50% of their strength with inclusion of antibiotics such as levofloxacin, the twisted bundle of a plurality of nanofibers, forming micron-wide sutures, generally do not lose strength with inclusion of an equivalent amount of levofloxacin. The multifilament sutures exhibit biocompatibility comparable to conventional nylon sutures, and are able to deliver active agent (e.g., levofloxacin) at detectable levels in eyes for at least 10, 12, 14, 16, 18, 20, 30 days, or longer. The antibiotic-eluting, multifilament sutures are generally able to prevent ocular infection and decrease bacterial load against one or multiple bacterial challenges for a period of about 1 week, 2 weeks, or longer in vivo, significantly more effective than a single post-operative antibiotic drop.
Exemplary therapeutic or prophylactic agents include, but are not limited to, anti-inflammatory agents such as dexamethasone, prednisolone, triamcinolone, and flurbiprofen, released in an effective amount to prevent post-operative inflammation resulting from the ophthalmic procedure or from the presence of the suture. Other therapeutic agents include rapamycin, neomycin, polymyxin B, bacitracin, gramicidin, gentamicin, oyxtetracycline, ciprofloxacin, ofloxacin, miconazole, itraconazole, trifluridine, and vidarabine to prevent or inhibit a disease or disorder. Sutures release anti-infective agents such as levofloxacin for a period of at least seven days, more preferably 30 days, in an effective amount to prevent or treat infection. The released drug from the multifilament suture itself may treat an infection that a common suture procedure is at high risk of developing or further developing.
In another embodiment, the drug-loaded, multifilament nanofibers may weave around and serve as a drug-eluting thin coating on existing or other commercially available sutures. The coating thickness is tunable, and the overall size of the multi-nanofiber-coated suture still meets U.S.P. size requirements.
The multifilament suture can be further coated with one or more materials for lubrication, glideability, permeation or impermeation, wettability, and/or non-fouling purposes.
The term “nanofiber” herein refers to a fiber of material with a thickness or diameter in the range of 1 nm to 1000 nm, while the length may be in the nanometer, micron, or millimeter range or greater. The term “filament” herein refers to a slender threadlike object, or in relevant sections refers to the disclosed nanofiber; whereas “multi-filament” refers to a plurality of filaments or the disclosed nanofibers, in a bundle.
The term “suture” herein refers to a thread or wire used to join together a wound or surgical incision. The disclosed multi-filament suture generally refers to a bundle of twisted nanofibers serving as the thread that may be attached or otherwise secured to a needle and be used by physicians or other medical professionals to join together wounds or incisions in surgery.
The term “electrospinning” refers to a technique that employs electric forces to elongate and decrease the diameter of a viscoelastic polymer stream, allowing for the formation of solid fibers ranging from nanometers to microns in diameter.
The term “grounded” generally refers to the status of connection to a ground. In electrical engineering, ground or earth is the reference point in an electrical circuit from which voltages are measured, a common return path for electric current, or a direct physical connection to the Earth. Therefore “grounded” as used herein in relation to electrospinning refers a collector acting as an electrode that is connected to ground or earth, as compared to a positive electrode (e.g., a charged needle tip or nozzle).
The term “collector” as used herein refers to a device where electrically charged solution, jet, melt, or gel is deposited onto in an electric field. Generally the collector is grounded, so as to provide a grounded electrode (that is apart from a positive electrode (e.g., electrically charged needle tip or nozzle)). The “collector” may also refer elements that attach or connect to the device where electrically charged solution, jet, melt, or gel is deposited onto, where the whole is electrically connected and grounded.
The term “chuck” as used herein refers to a type of clamp used to hold an object with radial symmetry (e.g., a cylinder), and herein may be mechanically and electrically connected via an adaptor to a rotator, in forming a part of a grounded collector. For examples, in drills, a chuck holds the rotating tool or workpiece.
The term “alkyl” refers to the radical of saturated aliphatic groups, including straight-chain alkyl groups, branched-chain alkyl groups, cycloalkyl (alicyclic) groups, alkyl-substituted cycloalkyl groups, and cycloalkyl-substituted alkyl groups.
In preferred embodiments, a straight chain or branched chain alkyl has 30 or fewer carbon atoms in its backbone (e.g., C1-C30 for straight chains, C3-C30 for branched chains), preferably 20 or fewer, more preferably 15 or fewer, most preferably 10 or fewer. Likewise, preferred cycloalkyls have from 3-10 carbon atoms in their ring structure, and more preferably have 5, 6 or 7 carbons in the ring structure. The term “alkyl” (or “lower alkyl”) as used throughout the specification, examples, and claims is intended to include both “unsubstituted alkyls” and “substituted alkyls”, the latter of which refers to alkyl moieties having one or more substituents replacing a hydrogen on one or more carbons of the hydrocarbon backbone. Such substituents include, but are not limited to, halogen, hydroxyl, carbonyl (such as a carboxyl, alkoxycarbonyl, formyl, or an acyl), thiocarbonyl (such as a thioester, a thioacetate, or a thioformate), alkoxyl, phosphoryl, phosphate, phosphonate, a hosphinate, amino, amido, amidine, imine, cyano, nitro, azido, sulfhydryl, alkylthio, sulfate, sulfonate, sulfamoyl, sulfonamido, sulfonyl, heterocyclyl, aralkyl, or an aromatic or heteroaromatic moiety.
Unless the number of carbons is otherwise specified, “lower alkyl” as used herein means an alkyl group, as defined above, but having from one to ten carbons, more preferably from one to six carbon atoms in its backbone structure. Likewise, “lower alkenyl” and “lower alkynyl” have similar chain lengths. Throughout the application, preferred alkyl groups are lower alkyls. In preferred embodiments, a substituent designated herein as alkyl is a lower alkyl.
It will be understood by those skilled in the art that the moieties substituted on the hydrocarbon chain can themselves be substituted, if appropriate. For instance, the substituents of a substituted alkyl may include halogen, hydroxy, nitro, thiols, amino, azido, imino, amido, phosphoryl (including phosphonate and phosphinate), sulfonyl (including sulfate, sulfonamido, sulfamoyl and sulfonate), and silyl groups, as well as ethers, alkylthios, carbonyls (including ketones, aldehydes, carboxylates, and esters), —CF3, and —CN. Cycloalkyls can be substituted in the same manner.
The term “mechanical strength”, as used herein, refers to any one of ultimate tensile strength (maximum stress bared until failure (N)), peak load, load at yield (breaking strength) (N), tenacity, initial stiffness (N/mm), or the modulus of elasticity (Young's modulus). The modulus of elasticity measures an object or substance's resistance to being deformed elastically (i.e., non-permanently) when a force is applied to it. The elastic modulus of an object is defined as the slope of its stress-strain curve in the elastic deformation region. It can be measured using the following Formula: E=Stress/Strain, where Stress is the force causing the deformation divided by the area to which the force is applied and Strain is the ratio of the change in some length parameter caused by the deformation to the original value of the length parameter. The modulus of elasticity is presented in Pascals (Pa), or megapascals (MPa). The term “attached”, as used herein, refers to the connection of elements in a system, generally via a mechanical means including, but not limited to, a clamp, a claw, a clip, an interlock, a screw, a magnetic attraction, an adhesive, or a vacuum suction. In some embodiments, “attached” can refer to elements that are already an integral piece of a whole device. It is interchangeable with “connected” as used herein.
The term “inhibit,” “inhibiting,” or “inhibition” refers to a decrease in activity, response, condition, disease, or other biological parameter. This can include but is not limited to the complete ablation of the activity, response, condition, or disease. This may also include, for example, a 10% reduction in the activity, response, condition, or disease as compared to the native or control level. Thus, the reduction can be a 10, 20, 30, 40, 50, 60, 70, 80, 90, 100%, or any amount of reduction in between as compared to native or control levels.
The term “prodrug”, as used herein, refers to compounds which, under physiological conditions, are converted into the therapeutically active agents of the present invention. A common method for making a prodrug is to include selected moieties which are hydrolyzed under physiological conditions to reveal the desired molecule. In other embodiments, the prodrug is converted by an enzymatic activity of the host animal.
The term “prevent,” “preventing,” or “prevention” does not require absolute forestalling of the condition or disease but can also include a reduction in the onset or severity of the disease or condition or inhibition of one or more symptoms of the disease or disorder.
The term “treat” or “treatment” refers to the medical management of a subject with the intent to cure, ameliorate, stabilize, or prevent a disease, pathological condition, or disorder. This term includes active treatment, that is, treatment directed specifically toward the improvement of a disease, pathological condition, or disorder, and also includes causal treatment, that is, treatment directed toward removal of the cause of the associated disease, pathological condition, or disorder. In addition, this term includes palliative treatment, that is, treatment designed for the relief of symptoms rather than the curing of the disease, pathological condition, or disorder; preventative treatment, that is, treatment directed to minimizing or partially or completely inhibiting the development of the associated disease, pathological condition, or disorder; and supportive treatment, that is, treatment employed to supplement another specific therapy directed toward the improvement of the associated disease, pathological condition, or disorder.
The term “therapeutic agent” refers to an agent that can be administered to prevent or treat one or more symptoms of a disease or disorder. These may be a nucleic acid, a nucleic acid analog, a small molecule, a peptidomimetic, a protein, peptide, carbohydrate or sugar, lipid, or surfactant, or a combination thereof.
The term “diagnostic agent”, as used herein, generally refers to an agent that can be administered to reveal, pinpoint, and define the localization of a pathological process.
The term “prophylactic agent”, as used herein, generally refers to an agent that can be administered to prevent disease or to prevent certain conditions like pregnancy.
The phrase “pharmaceutically acceptable” refers to compositions, polymers and other materials and/or dosage forms which are, within the scope of sound medical judgment, suitable for use in contact with the tissues of human beings and animals without excessive toxicity, irritation, allergic response, or other problem or complication, commensurate with a reasonable benefit/risk ratio. The phrase “pharmaceutically acceptable carrier” refers to pharmaceutically acceptable materials, compositions or vehicles, such as a liquid or solid filler, diluent, solvent or encapsulating material involved in carrying or transporting any subject composition, from one organ, or portion of the body, to another organ, or portion of the body. Each carrier must be “acceptable” in the sense of being compatible with the other ingredients of a subject composition and not injurious to the patient.
The term “biodegradable” as used herein, generally refers to a material that will degrade or erode under physiologic conditions to smaller units or chemical species that are capable of being metabolized, eliminated, or excreted by the subject. The degradation time is a function of composition and morphology. Degradation times can be from hours to years.
The term “biocompatible” as used herein, generally refers to materials that are, along with any metabolites or degradation products thereof, generally non-toxic to the recipient, and do not cause any significant adverse effects to the recipient. Generally speaking, biocompatible materials are materials which do not elicit a significant inflammatory or immune response when administered to a patient.
The term “degrade”, as used herein, refers to a reduction in one or more properties of the polymer over time. The one or more properties include the molecular weight, total mass, mechanical strength, elasticity, or the density or porosity of the fibers formed from polymers. Generally, a degradable polymer is capable of being absorbed by living mammalian tissue. This can occur over a period of days, weeks, months, or years. The prevailing mechanism of degradation of hydrolytically biodegradable polymers is chemical hydrolysis of the hydrolytically unstable backbone. In a bulk eroding polymer, the polymer network is fully hydrated and chemically degraded throughout the entire polymer volume. As the polymer degrades, the molecular weight decreases. The reduction in molecular weight is followed by a decrease in mechanical properties (e.g., strength) and scaffold properties. The decrease of mechanical properties is followed by loss of mechanical integrity and then erosion or mass loss (Pistner et al., Biomaterials, 14: 291-298 (1993)). Non-degradable polymer is suitably resistant to the action of living mammalian tissue. A similar distinction between non-absorbable and absorbable sutures are used by the United States Pharmacopeia (U.S.P.).
Use of the term “about” is intended to describe values either above or below the stated value in a range of approximately +/−10%. The preceding ranges are intended to be made clear by context, and no further limitation is implied.
1. Twisting and Braiding of Fibers Using Electrospinning
Electrospinning is a versatile technique first introduced in the early 1900's, which employs electric forces to elongate and reduce the diameter of a viscoelastic polymer jet, allowing for the formation of solid fibers ranging from nanometers to microns in diameter (Bhardwaj, N., et al., Biotechnol Adv, 28, 325-347 (2010); Li, D. & Xia, Y., Advanced Materials, 16, 1151-1170 (2004)). An electrostatic charge is applied on the needle to overcome the surface tension of the solution. Usually, the concentration of the polymer solution in electrospinning is greater than a minimum concentration for any given polymer, termed the critical entanglement concentration, below which a stable jet cannot be achieved and no nanofibers will form, although nanoparticles may be achieved (electrospray) (Leach M K, et al., J Vis Exp., (47): 2494 (2011)). The multifilament composite fibers built upon electrospinning, are twisted or braided to form ultra-thin, high strength, drug-loaded sutures or to coat a commercially available suture or thread to provide additional therapeutic functionality.
A. Twisting
As shown in
In another embodiment, a suture, thread or equivalent that is conductive or non-conductive, which can be any commercially available suture, is placed between a drill chuck 370 and a parallel stand 302, where the thread end of the suture is fixed at the drill chuck 370 and the needle end is placed through the stand 302 but kept free to rotate, as shown in
The individual fibers can be so thin that they are able to align the internal polymer chains without the use of heat treatment or extrusion to provide increased strength.
In some embodiments, when the parallel collector is also connected to a motor or is a second drill chuck that is connected to a motor, the chuck can rotate in one direction, e.g., clockwise, to twist these fibers, while the other end is held stationary on the parallel collector. In other embodiments, both ends of the fiber(s) (e.g., in opposite directions, or at different speed) are rotated to twist the fibers.
The twisted fibers can optionally be further twisted in the opposite direction, e.g., counterclockwise, to ensure that the twisted fibers do not coil or snap.
When a drill chuck is rotated 360° relative to the opposing collector, one twist is done to the fiber(s). To form densely twisted fibers of sufficient strength, hundreds, thousands or tens of thousands of twists can be done to the fibers. For instance, when the distance between the drill chuck and the opposing collector is about 50 cm, 60 cm, 70 cm, 80 cm, 90 cm, or 100 cm, twists of increasing numbers can be done to the fibers, e.g., 500 twists, 1,000 twists, 1,500 twists, 2,000 twists, 2,500 twists, 3,000 twists, 3,500 twists, and 4,000 twists, respectively, or even greater. As the number of twists increases, the diameter of the overall fiber bundle generally decreases, and the strength generally increases.
Even when drug loading decreases the strength of individual fibers, twisting the fibers reduces the loss of tensile strength or even increases the strength for the multifilament composite. The number of twists needed to meet certain strength parameters will vary depending on the composition of the polymer/drug, and the size of individual fibers. For instance, with fibers made from a high molecular weight (e.g., 220 kDa) poly (L-lactic acid) (PLLA), twists ranging from 2,000 to 4,000 are generally needed to generated twisted fibers that meet the strength requirement for sutures according to United States Pharmacopeia (USP). Alternately, certain types of polycaprolactone (PCL), with or without certain drugs, can be twisted at a lower number, e.g., much below 1,575 twists, and still surpass strength requirements. One can increase the number of twists and decrease the diameter while maintaining strength. In some embodiments, including therapeutic, prophylactic or diagnostic agents up to about 5%, 10%, 15%, 20%, 25%, 30% by weight or even greater, can still meet USP requirements for strength.
Picking highly crystalline polymers and using predominantly nanofibers which have more aligned polymer chains and greater individual tensile modulus, then twisting these to make them stronger and more impervious to damage are key to producing small diameter very strong fibers. Given the high crystallinity and aligned polymer chains of a nanofiber, and potentially the hydrophobicity, the agents move to the outside/surface of each nanofiber rather than being mixed in with the polymer chains and polymer matrix, which would lead to decreased strength and is what happens to micron-sized monofilaments.
B. Braiding
A braid is an organization of three or more fibers or fiber bundles intertwined in such a way that no two fibers (or fiber bundles) are twisted around one another. In one embodiment, fibers are removed from the collector(s) and placed into braiding machines known in the art to form braids of fibers. The electrospinning system twists rather than braids. Several composite fibers can be collected and attached to the drill chuck and/or to a standstill or rotating parallel stand and the drill chuck rotated to twist the composite fibers together in the same way that individual nanofibers are twisted together to manufacture the composite fiber.
A component of a system for removing a fiber from a collection surface needs not be in the illustrated form. Any suitable component can be included to remove the fibers such as, without limitation, a blade, a wedge, a plate, or any other shaped device that can shear or cut the fiber from the collection surface.
In some embodiments, multiple 10-0 and/or 11-0 nanofiber sutures may be braided to make drug-loaded sutures that meet U.S.P. specifications for 2-0-7-0 sutures.
2. Configuration of Electrospinning Apparatus
Generally, the parallel collectors are in a lined up in a position that is perpendicular to the needle or nozzle. The needle or nozzle can be 90°, 85°, 80°, 75°, 70°, 65°, 60°, 55°, 50°, 45°, 40°, 35°, 30°, or at another non-parallel angle with respect to the collectors. The distance between the end of a needle or the tip of a nozzle and the collectors can be between about 4 cm and about 100 cm, or even greater. In some embodiments, this distance is between about 6 cm and about 25 cm.
The distance between the motor and the parallel stand (d) can be between about 2 mm up to about 200 cm, or even greater, e.g. distances between 270 and 202 in
Generally, the heights of the collectors are about the same, i.e., parallel collectors. In other instances, the heights of the opposing collectors can be of different heights, by difference of 10%, 20%, 30%, 40% or greater, of the taller collector.
A polymer solution, sol-gel, suspension or melt may be loaded into the electrospinning ejection device (e.g., needled syringe, nozzle). The needle can have be a standard needle having a diameter between 34 gauge and 7 gauge, where diameter decreases with gauge size. In some embodiments, multiple needles are used to generate multiple streams of polymer jets towards the collectors. The height of the needle or nozzle where the polymer jet starts from can be the same or different from the height(s) of the collector(s). In preferred embodiments, the height of the needle is greater than that of the parallel collectors. In one embodiment, the needle is pointed in a horizontal orientation, and in another embodiment, the needle is pointed in a vertical orientation. The angle that the needle is at relative to the horizontal level can be 0°, 10°, 20°, 30°, 40°, 50°, 60°, 70°, 80°, or 90°, preferably from a height no shorter than the height of the collectors. The needles or syringes where the needles are attached can be mounted onto a motorized platform, e.g., a stage, a dispenser, to allow for alterations in the configuration of the system or movement of the needles.
The polymer solution can be held in a syringe that is controlled by a programmable syringe pump known in the art. The gauge of the needle, the speed that the polymer solution is pushed out from the needle, and the volume of polymer to be electrospun can be tuned, according to the composition and the viscosity of the solution, the configuration of the collectors, and the desired properties of formed fibers. In some embodiments, multiple needles are used to generate multiple streams of polymer jets on the collectors.
The syringe pump can be mounted onto a base atop a motorized stage known in the art. This controls the motion of the needle in the x direction and the y direction. Moving along an x-direction may position the needle closer or farther away from the collectors, while moving along a y-direction may position the needle at a constant distance from the center-line of the parallel collectors.
The critical field strength required to overcome the forces due to surface tension of the solution and fotin a jet will depend on many variables of the system. These variables include not only the type of polymer and solvent, but also the solution concentration and viscosity, as well as the temperature of the system. In general, characterization of the jet formed, and hence characterization of the fibers formed, depends primarily upon solution viscosity, net charge density carried by the electrospinning jet and surface tension of the solution. The ability to form small diameter fibers depends upon the combination of all of the various parameters involved. For example, electrospinning of lower viscosity solutions will tend to form beaded fibers, rather than smooth fibers. In fact, many low viscosity solutions of low molecular weight polymers will break up into droplets or beads, rather than form fibers, when attempts are made to electrostatically spin the solution. Solutions having higher values of surface tension also tend to form beaded fibers or merely beads of polymer material, rather than smooth fibers. Thus, the preferred solvent for any particular embodiment will generally depend upon the other materials as well as the formation parameters, as is known in the art.
In some embodiments, the electrospun fibers are made under sterile conditions to avoid the need for subsequent sterilization, although most sutures can be sterilized by means such as ethylene oxide.
Additional elements can be included in the electrospinning system, such as means for temperature control, for example, a strip heater, fan, or a temperature controller.
One or two motors can be connected to one or both opposing collectors. Generally, a motor has a shaft to allow for connection with specific collectors via the adaptor. Generally, the motor is capable of providing clock-wise or counter clock-wise rotation up to a speed of 1,000 revolutions per minute (rpm), 2,000 rpm, 3,000 rpm, 4,000 rpm, 5,000 rpm, 6,000 rpm, 7,000 rpm, 8,000 rpm, 9,000 rpm, 10,000 rpm, or greater.
3. Other Methods to Fabricate Multifilament Fibers.
Other methods may be used to fabricate ultra-thin, twisted multi-nanofiber sutures with sufficient strength and drug loading capacity. In general, thin nanofibers are fabricated and twisted in a bundle to compact and strengthen the mechanical properties. Single needle single jet, single needle multi-jets, multi-needle multi-jets, or even needleless configurations for electrospinning may be used to fabricate polymeric nanofibers, which later are twisted by rotating two ends of the bundle of nanofibers in different angular speed and/or in different angular directions. Techniques other than electrospinning may also be used to fabricate polymeric nanofibers, such as meltblowing, bicomponent spinning, forcespinning, and flash-spinning.
Meltblowing
In a meltblowing process, a molten polymer is extruded through the orifice of a die. The fibers are formed by the elongation of the polymer streams coming out of the orifice by air-drag and are collected on the surface of a suitable collector in the form of a web. The average fiber diameter mainly depends on the throughput rate, melt viscosity, melt temperature, air temperature and air velocity. Nanofibers can be fabricated by special die designs with a small orifice, reducing the viscosity of the polymeric melt and suitable modification of the meltblowing setup. To reduce or prevent the sudden cooling of the fiber as it leaves the die before the formation of nanofibers, hot air flow may be provided in the same direction of the polymer around the die. The hot air stream flowing along the filaments helps in attenuating them to smaller diameter. The viscosity of polymeric melt can be lowered by increasing the temperature.
Template Melt-Extrusion
In template melt-extrusion, molten polymer is forced through the pores of a template (e.g., an anodic aluminum oxide membrane (AAOM)) and then subsequently cooled down to room temperature. A special stainless steel appliance may be designed to support the template, to bear the pressure and to restrict the molten polymer movement along the direction of the pores. The appliance containing the polymer was placed on the hot plate of a compressor (with temperature controlled functions) followed by the forcing of the polymeric melt. Isolated nanofibers may be obtained by the removal of the template (e.g., dissolution with appropriate solvent(s)).
Flash-Spinning
In the flash-spinning process, a solution of fiber forming polymer in a liquid spin agent is spun into a zone of lower temperature and substantially lower pressure to generate plexi-filamentary film-fibril strands. A spin agent is required for flash-spinning which: 1) should be a non-solvent to the polymer below its normal boiling point, 2) can form a solution with the polymer at high pressure, 3) can form a desired two-phase dispersion with the polymer when the solution pressure is reduced slightly, and 4) should vaporize when the flash is released into a substantially low pressure zone. Flash-spinning is more suitable for difficult to dissolve polymers such as polyolefins and high molecular weight polymers. The spinning temperature should be higher than the melting point of polymer and the boiling point of solvent in order to effect solvent evaporation prior to the collection of the polymer.
Bicomponent Spinning
Bicomponent spinning is a two-step process that involves spinning two polymers through the spinning die (which forms the bicomponent fiber with island-in-sea (IIS), side-by-side, sheath-core, citrus or segmented-pie structure) and the removal of one polymer.
Other Approaches
In some embodiments, the disclosed stent devices may be prepared via other methods than electrospinning, such as 3-D printing and dipping in polymer solutions and drying of a cylindrical/wire-shaped template, followed by removal of the template after the polymeric wall of the stent is formed, resulting in a stent device with a wall surrounding a lumen.
1. Properties
A. Diameter & Strength
Suture diameters are defined by the United States Pharmacopeia (U.S.P.). Modern sutures range from #2 (heavy braided suture) to #11-0 (fine monofilament suture for ophthalmics). Suitable diameters for ophthalmic use are USP size 6.0-11.0, preferably 7.0-11.0, more preferably 8.0-11.0, most preferably 9.0-11.0.
Absorbable sutures as formed satisfy the strength requirement for absorbable sutures set forth in the United States Pharmacopeia (Table 1).
For non-absorbable sutures, strength requirements for different diameter sutures are classed on the United States Pharmacopeia, e.g., based on the type of coating. For example, class I suture is composed of silk or synthetic fibers of monofilament, twisted, or braided construction where the coating, if any, does not significantly affect thickness (e.g., braided silk, polyester, or nylon; monofilament nylon or polypropylene); class II suture is composed of cotton or linen fibers or coated natural or synthetic fibers where the coating significantly affects thickness but does not contribute significantly to strength (e.g., virgin silk sutures); and class III suture is composed of monofilament or multifilament metal wire.
Non-absorbable sutures as formed satisfy the strength requirement for non-absorbable sutures set forth in the United States Pharmacopeia (Table 2).
aThe tensile strength of sizes smaller than USP size 8-0 (metric size 0.4) is measured by straight pull. The tensile strength of sizes larger than USP size 2-0 (metric size 3) of monofilament Class III (metallic) nonabsorbable surgical suture is measured by straight pull. Silver wire meets the tensile strength values of Class I sutures but is tested in the same manner as class III sutures.
bThe limits on knot-pull tensile strength apply to nonabsorbable surgical suture that has been sterilized. For nonsterile sutures of Class I and Class II, the limits are 25% higher.
indicates data missing or illegible when filed
The sutures as formed typically have a diameter between 10 μm and about 400 μm, preferably between 10 and about 50 microns, more preferably between about 20 and about 50 microns. The sutures typically have a Young's modulus of at least about 700, 800, 900, 1000, 1100, 1200, 1300, 1400, 1500, 1600, 1700, or 1800 MPa. Specifically, as fibers are twisted, the overall diameter of the multifilament “bundle” generally decreases, e.g., 500 μm, 400 μm, 300 μm, 200 μm, 150 μm, 100 μm, 50 μm, 40 μm, 30 μm, 20 μm, or 10 μm, while the “bundle” is of sufficient tensile strength according to USP standards. Essentially, the more the fibers are twisted, the more compact the composite becomes, and the more fibers that can be packed into a single suture of a specific size. This is one of the main reasons twisting increases strength.
In some embodiments, the sutures have the above diameter and a tensile modulus (e.g., Young's modulus) of at least about 600, 650, 700, 750, 800, 850, 900, 950, or 1000 MPa. In particular embodiments, the sutures have a tensile modulus (e.g., Young's modulus) from about 1000 to about 2500 MPa, preferably from about 1200 to about 2500 MPa, more preferably from about 1300 to about 2300 MPa. The sutures should retain the tensile strength for the requisite period of time.
In some embodiments of drug-loaded sutures, the sutures have a breaking strength as specified by the USP in Tables 1 and 2 for different diameters.
B. Absorbability
In some embodiments, the multifilament fibers are made with absorbable materials which are broken down in tissue after a given period of time, which depending on the material can be from ten days to one year. They can be used in many of the internal tissues of the body. In cases where three weeks is sufficient for the wound to close thinly, the suture is not needed any more, and absorbable multifilament fibers leave no foreign material inside the body and no need for the patient to have the sutures removed. Examples of absorbable polymers are listed as biodegradable polymers below.
In other embodiments, the multifilament fibers are made with non-absorbable materials which are not metabolized by the body. They can be used either on skin wound closure, where the sutures can be removed after a few weeks, or in some inner tissues in which absorbable sutures are not adequate.
Generally, any polymer can be used in electrospinning to prepare multifilament fibers. For instance, the rhythmic movement in the heart and in blood vessels may require a suture material which stays longer than three weeks, to give the wound enough time to close. Other organs, like the bladder, contain fluids may make absorbable sutures disappear too soon for the wound to heal. Hence, a non-absorbable or a mixture of absorbable and non-absorbable materials is used to prepare electrospun fibers and to twist into sutures or coatings for sutures. Examples of non-absorbable polymers are listed as nondegradable polymers below.
In an embodiment where the electrospun fibers are used to coat existing absorbable or non-absorbable sutures or suture thread, the fibers can be made with degradable or non-degradable polymer.
C. Coating Properties
The multifilament fibers can be used to coat existing sutures to provide desired surface properties. In some embodiments, the coating fibers include one or more therapeutic, prophylactic or diagnostic agents that are encapsulated in the nanofibers, which upon twisting onto non-agent-eluting suture allows for sustained release of the agent.
In other embodiments, the coating improves glide and reduces irritation and capillarity while still maintaining good knot security. By twisting the electrospun fibers around a suture, the coating fibers are still thin enough, do not compromise the strength of the sutures, and do not get easily rubbed off during manipulation. The lower the coefficient of friction, the less the thread gets stuck and injures the tissues. For instance, the glyconate, i.e., a copolymer made of glycolide (e.g., 72%), trimethylene carbonate (e.g., 14%), and caprolactone (e.g., 14%), can be used to prepare the electrospun fibers or coated onto the composite suture, combining good glideability and/or knot security.
2. Compositions
A. Polymers
In some embodiments, polymers that have been found suitable for use in biological applications can be utilized. In some embodiments, polymers that are degradable can be utilized. Non-degradable polymers can be utilized alone, in combination, or in sequence with degradable polymers.
A polymeric solution that is loaded into an electrospinning nozzle or syringe can include any suitable solvent. Selection of solvent can be important in determining the characteristics of the solution, and hence of the characteristic properties of the nanofibers formed during the process. Examples include hexafluoroisopropanol, methanol, chloroform, dichloromethane, dimethylformamide, acetone, acetic acid, acetonitrile, m-cresole, tetrahydrofuran (THF), toluene, as well as mixtures of solvents.
Preferred polymers including polyhydroxy acids such as poly(lactic acid), poly(glycolic acid) and poly(lactic-co-glycolic acid), polycaprolactone, polydioxanone, as well as combinations of polymers (i.e., poly-1-lactic acid/polyethylene glycol) having a molecular weight between 5 kDa and 500 kDa.
Other examples of suitable biodegradable, biocompatible polymers include polyhydroxyalkanoates such as poly-3-hydroxybutyrate or poly-4-hydroxybutyrate; poly(orthoesters); polyanhydrides; poly(phosphazenes); poly(lactide-co-caprolactones); polycarbonates such as tyrosine polycarbonates; polyamides (including synthetic and natural polyamides); polyesteramides; other polyesters; poly(dioxanones); polyurethanes; polyetheresters; polymethylmethacrylates; polysiloxanes; poly(oxyethylene)/poly(oxypropylene) copolymers; polyketals; polyphosphates; polyhydroxyvalerates; as well as copolymers thereof.
In the most preferred embodiments, the biodegradable polymer is polycaprolactone or polyglycolide or a poly-(D,L-lactide-co-glycolide) such as poly-(D,L-lactide-co-glycolide) containing about 55 to about 80 mole % lactide monomer and about 45 to about 20 mole % glycolide and poly-(D, L-lactide-co-glycolide) containing about 65 to about 75 mole % lactide monomer and about 35 to about 25 mole % glycolide. The poly-(D, L-lactide-co-glycolide) can contain terminal acid groups.
The biodegradable, biocompatible polymer can be a polylactic acid polymer or copolymer containing lactide units substituted with alkyl moieties. Examples include, but are not limited to, poly(hexyl-substituted lactide) or poly(dihexyl-substituted lactide).
The molecular weight of the polymer can be varied to optimize the desired properties, such as drug release rate, tensile strength and tensile modulus, and degradation for specific applications. The one or more biodegradable, biocompatible polymers can have a molecular weight of between about 1 kDa and 500 kDa. In certain embodiments, the biodegradable, biocompatible polymers have a molecular weight of between about 15 kDa and about 300 kDa, more preferably between about 50 kDa and about 2000 kDa.
Non-degradable polymers include polyurethanes, silicones or silicon elastomers, polyesters, acrylic polymers and copolymers, vinyl halide polymers and copolymers, polyvinyl chloride, polyvinyl ethers, ethylene-methyl methacrylate copolymers, s, ABC resins and ethylene-vinyl acetate copolymers, polyamides such as nylon 66 and polycaprolactam, polyimides; polyethers, and a biocompatible polymer according to claim 1 selected from the group consisting of a combination thereof.
B. Therapeutic, Prophylactic or Diagnostic Agents
The fibers may include one or more therapeutic, prophylactic, or diagnostic agents that are blended, encapsulated, conjugated to the polymer in a solution before electro spinning, or encapsulated in/conjugated to sustained release nanoparticle/microparticle formulations that are entrapped in between or conjugated with the formed fibers. These may be proteins, peptides, nucleic acid, carbohydrate, lipid, or combinations thereof, or small molecules. Suitable small molecule active agents include organic and organometallic compounds. In some instances, the small molecule active agent has a molecular weight of less than about 2000 g/mol, preferably less than about 1500 g/mol, more preferably less than about 1200 g/mol, most preferably less than about 1000 g/mol. In other embodiments, the small molecule active agent has a molecular weight less than about 500 g/mol. The small molecule active agent can be a hydrophilic, hydrophobic, or amphiphilic compound. Biomolecules typically have a molecular weight of greater than about 2000 g/mol and may be composed of repeat units such as amino acids (peptide, proteins, enzymes, etc.) or nitrogenous base units (nucleic acids). In preferred embodiments, the active agent is an ophthalmic therapeutic, prophylactic or diagnostic agent.
Representative therapeutic agents include, but are not limited to, analgesic agents, anti-fibrotic/anti-scarring agents, anti-inflammatory drugs, including immunosuppressant agents and anti-allergenic agents, anti-infectious, and anesthetic agents. Exemplary analgesic agents include simple analgesics (e.g., paracetamol, aspirin), non-steroidal anti-inflammatory drugs (e.g., ibuprofen, diclofenac sodium, naproxen sodium), weaker opioids (e.g., combinations including codeine phosphate, tramadol hydrochloride, dextropropoxyphe hydrochloride and paracetamol), and stronger opioids (e.g., morphine sulphate, oxycodone, pethidine hydrochloride). Some examples of anti-inflammatory drugs include triamcinolone acetonide, fluocinolone acetonide, prednisolone, dexamethasone, loteprendol, fluorometholone. Immune modulating drugs such as: cyclosporine, tacrolimus and rapamycin. Non-steroidal anti-inflammatory drugs include ketorolac, nepafenac, and diclofenac. Antiinfectious agents include antiviral agents, antibacterial agents, antiparasitic agents, and anti-fungal agents. Exemplary antibiotics include moxifloxacin, ciprofloxacin, erythromycin, levofloxacin, cefazolin, vancomycin, tigecycline, gentamycin, tobramycin, ceftazidime, ofloxacin, gatifloxacin, rapamycin; antifungals: amphotericin, voriconazole, natamycin. Exemplary steroids suitable to include in the disclosed suture include, but are not limited to, testosterone, cholic acid, dexamethasone, lanosterol, progesterone, medrogestone, and β-sitosterol.
In some embodiments, levofloxacin, moxifloxacin, bacitracin, sirolimus, sunitinib, triamcinolone acetonide, cyclosporine, and dexamethasone are included individually or in combination in the formulations.
For ophthalmology applications, active agents can include anti-glaucoma agents that lower intraocular pressure (IOP), anti-angiogenesis agents, growth factors, and combinations thereof for treatment of vascular disorders or diseases. Examples of anti-glaucoma agents include mitomycin C, prostaglandin analogs such as travoprost and latanoprost, prostamides such as bimatoprost; beta-adrenergic receptor antagonists such as timolol, betaxolol, levobetaxolol, and carteolol, alpha-2 adrenergic receptor agonists such as brimonidine and apraclonidine, carbonic anhydrase inhibitors such as brinzolamide, acetazolamine, and dorzolamide, miotics (i.e., parasympathomimetics) such as pilocarpine and ecothiopate), seretonergics, muscarinics, and dopaminergic agonists.
Representative anti-angiogenesis agents include, but are not limited to, antibodies to vascular endothelial growth factor (VEGF) such as bevacizumab (AVASTIN®) and rhuFAb V2 (ranibizumab, LUCENTIS®), and other anti-VEGF compounds including aflibercept (EYLEA®); MACUGEN® (pegaptanim sodium, anti-VEGF aptamer or EYE001) (Eyetech Pharmaceuticals); pigment epithelium derived factor(s) (PEDF); COX-2 inhibitors such as celecoxib (CELEBREX®) and rofecoxib (VIOXX®); interferon alpha; interleukin-12 (IL-12); thalidomide (THALOMID®) and derivatives thereof such as lenalidomide (REVLIMID®); squalamine; endostatin; angiostatin; ribozyme inhibitors such as ANGIOZYME® (Sirna Therapeutics); multifunctional antiangiogenic agents such as NEOVASTAT® (AE-941) (Aeterna Laboratories, Quebec City, Canada); receptor tyrosine kinase (RTK) inhibitors such as sunitinib (SUTENT®); tyrosine kinase inhibitors such as sorafenib (Nexavar®) and erlotinib (Tarceva®); antibodies to the epidermal grown factor receptor such as panitumumab (VECTIBIX®) and cetuximab (ERBITUX®), as well as other anti-angiogenesis agents known in the art.
In some cases, the agent is a diagnostic agent imaging or otherwise assessing the tissue of interest. Examples of diagnostic agents include paramagnetic molecules, fluorescent compounds, magnetic molecules, and radionuclides, x-ray imaging agents, and contrast media.
The active agents may be present in their neutral form, or in the form of a pharmaceutically acceptable salt. In some cases, it may be desirable to prepare a formulation containing a salt of an active agent due to one or more of the salt's advantageous physical properties, such as enhanced stability or a desirable solubility or dissolution profile.
Generally, pharmaceutically acceptable salts can be prepared by reaction of the free acid or base forms of an active agent with a stoichiometric amount of the appropriate base or acid in water or in an organic solvent, or in a mixture of the two; generally, non-aqueous media like ether, ethyl acetate, ethanol, isopropanol, or acetonitrile are preferred. Pharmaceutically acceptable salts include salts of an active agent derived from inorganic acids, organic acids, alkali metal salts, and alkaline earth metal salts as well as salts formed by reaction of the drug with a suitable organic ligand (e.g., quaternary ammonium salts). Lists of suitable salts are found, for example, in Remington's Pharmaceutical Sciences, 20th ed., Lippincott Williams & Wilkins, Baltimore, Md., 2000, p. 704. Examples of ophthalmic drugs sometimes administered in the form of a pharmaceutically acceptable salt include timolol maleate, brimonidine tartrate, and sodium diclofenac.
The agent or agents can be directly dispersed or incorporated into the fibers as particles using common solvent with the polymer, for examples microparticles and/or nanoparticles of drug alone, or microparticles and/or nanoparticles containing a matrix, such as a polymer matrix, in which the agent or agents are encapsulated or otherwise associated with the particles.
The concentration of the drug in the finished fibers or foamed structures of fibers can vary. In some embodiments, the amount of drug is between about 0.1% and about 50% by weight, preferably between about 1% and about 20% by weight, more preferably between about 3% and about 20% by weight, most preferably between about 5% and about 20% by weight of the finished stents.
In particular embodiments, the agent is released at an effective amount to inhibit, prevent, or treat disorders or diseases in ophthalmology, cardiology, or neurology among others for at least 2 weeks, 4 weeks, 6 weeks, 8 weeks, 10 weeks, 12 weeks, 16 weeks, or 20 weeks.
The sutures can be modified by inclusion of a hydrophilic polymer such as PEG or a POLOXAMER® to provide a burst release of an active agent, such as an antimicrobial agent, followed by sustained release over an extended period of time, such as one week, two weeks, 4 weeks, 6 weeks, 8 weeks, 10 weeks, 12 weeks, 14 weeks, 16 weeks, 18 weeks, or 20 weeks.
C. Formulations
The amount of polymer or polymers in the finished fibers can vary. In some embodiments, the concentration of the polymer or polymers is from about 75% to about 85% by weight of the finished fibers. In some embodiments, the concentration of the polymer or polymers is from about 85% to about 100% by weight of the finished fibers.
Representative excipients include pH modifying agents, preservatives, antioxidants, suspending agents, wetting agents, viscosity modifiers, tonicity agents, stabilizing agents, and combinations thereof. There may be residual levels of solvent. Suitable pharmaceutically acceptable excipients are preferably selected from materials which are generally recognized as safe (GRAS), and may be administered to an individual without causing undesirable biological side effects or unwanted interactions.
D. Kit or Packaging
In some embodiments, the suture is sterilized and packaged dry or in fluid, in containers (e.g., packets) so designed that sterility is maintained until the container is opened. In some embodiments, the suture is secured to a needle, which is sterilized and packaged.
1. Sutures
The multi-filament polymeric sutures can be used simultaneously as drug delivery vehicles and sutures to close wounds in surgery including, but not limited to, ophthalmology (e.g., antibacterial operations; corneal transplant; trauma; acanthamoeba keratitis; specialty surgeries), cardiology, vascular surgery (e.g., anastomosis, or grafts), plastic surgery (e.g., keloids/hypertrophic scar removal; steroid-loaded to prevent scarring after surgery), regenerative medicine (e.g., peripheral nerve regeneration), and pediatric surgery. In the eye, applications include ones created in ophthalmologic surgery or due to injury or trauma to the eye. They can also be used to promote vascular or graft anastomosis. The fibers can also be used as sutures for skin wound or internal tissue, e.g., nerve, heart, bladder etc.
The sutures can be used in a variety of ophthalmic procedures known in the art. Examples include, but are not limited to, trabeculectomy as well as pterygium removal, cataract surgery, strabismus correction surgery, penetrating keratoplasty, sclerectomy, and conjunctival closure.
Trabeculectomy is an ophthalmic surgical procedure used in the treatment of glaucoma. Removing part of the eye's trabecular meshwork and adjacent structures allows drainage of aqueous humor from within the eye to underneath the conjunctiva to relieve intraocular pressure. The scleral flap is typically sutured loosely back in place with several sutures. Common complications include blebitis (an isolated bleb infection typically caused by microorganisms such as Staphylococcus epidermidis, Propriobacterium acnes, or Staphylococcus aureus), inflammation, and bleb-associated endophthalmitis.
Endophthalmitis is an inflammation of the ocular cavities and their adjacent structures. It is a possible complication of all intraocular surgeries, particularly cataract surgery, which can result in loss of vision or the eye itself. Endophthalmitis is usually accompanied by severe pain, loss of vision, and redness of the conjunctiva and the underlying episclera. Infectious etiology is the most common and various bacteria and fungi have been isolated as the cause of the endophthalmitis. The patient needs urgent examination by an ophthalmologist and/or vitreo-retina specialist who will usually decide for urgent intervention to provide intravitreal injection of potent antibiotics and also prepare for an urgent pars plana vitrectomy as needed. Enucleation may be required to remove a blind and painful eye.
Ideally, ophthalmic suture materials are biodegradable or absorbable and biodegradable over the useful suture lifetime, retaining the requisite tensile strength, mechanical modulus, and capable of delivering therapeutic or prophylactic agents to increase patient success. However, it is not essential the suture be biodegradable. For pterygium removal, cataract surgery and strabismus correction surgery, sutures can be used to close the wound and release antibiotic and anti-inflammatory drugs. For trabeculectomy surgeries, sutures can be placed on sclera flaps providing local chemotherapeutic agents, decreasing production of scar tissue, and on conjunctival closure with antibiotic release. In penetrating keratoplasty, the sutures hold the graft, as well as release antibiotic and immunosuppressant agents.
2. Coatings
The twisted multifilament fibers can weave around and coat existing sutures or other devices, to provide additional functionality and/or surface properties. In some embodiments, the coating fibers include one or more therapeutic, prophylactic or diagnostic agents that are encapsulated in the nanofibers. Two or more populations of fibers including different active agents can be twisted and form the coating on a suture, and depending on the sequence of coating and tightness of twisting, the release profiles for the different active agents are controlled and finely tuned.
In other embodiments, the coating improves glide and reduces irritation and capillarity while still maintaining good knot security. By twisting the electrospun fibers around a suture, the coating fibers are still thin enough, do not compromise the strength of the sutures, and do not get easily rubbed off during manipulation. Specifically, the lower the coefficient of friction, the less the thread gets stuck and injures the tissues. For instance, a glyconate, i.e., a copolymer made of glycolide (e.g., 72%), trimethylene carbonate (e.g., 14%), and caprolactone (e.g., 14%), can be used to prepare the electrospun fibers, combining good glideability and knot security.
3. Other Applications
Mixtures of materials can be electrospun to form composite fibers. For instance, a solution including one or more polymers in combination with a non-polymeric additive can be electrospun to form composite fibers. Additives are generally selected based upon the desired application of the formed fiber structures. For example, one or more polymers can be electrospun with a biologically active additive that can be polymeric or non-polymeric, as desired. By way of example, a 3D structure of fibers can include an electrospun polymer in conjunction with one or more therapeutic, prophylactic, and/or diagnostic agent. The secondary material can be incorporated in the fibers during formation as is known in the art, for example, as described in U.S. Pat. No. 6,821,479 to Smith, et al., U.S. Pat. No. 6,753,454 to Smith, et al., and U.S. Pat. No. 6,743,273 to Chung, et al.
In other embodiments, the ultra-thin, high strength multifilament fibers can be braided into membranes with defined interstices for industrial applications, e.g., water purification.
The present invention will be further understood by reference to the following non-limiting examples.
Suture breaking strength, Levo concentration, and bacterial load are presented as mean±standard error below. Statistical significance for breaking strength and bacterial load data has been determined via one-way ANOVA followed by Tukey test. Statistical significance for the Kaplan-Meier curve of long-term infection prevention was determined via the Mantel-Cox test.
Materials and Methods
Polycaprolactone (PCL), polylactic acid (PLLA), and poly(lactic-co-glycolic acid) (PLGA, 75:25) used were all at a molecular weight of 80 kDa; polyglycolide (PGA) and polydioxanone (PDO) used were the only commercially available polymers from Purac: Corbion and Sigma Aldrich, respectively.
As shown in
A 120 W regulated high voltage DC power source 200 applied a voltage to a blunt tip needle 210 on the end of a syringe 220. This allowed for the ejection, from the syringe, of electrified polymer solution 230 held in the syringe 220 and the syringe was controlled for flow rate by a NE-1000 Programmable Single Syringe Pump 240 mounted onto a plexiglass base atop a motorized stage capable of controlled x- and y-direction motion. The drill chuck 270 was connected to a mounted 120V, ⅓ hp, 300-3,450 rpm speed-control motor 250 (capable of clockwise or counter-clockwise rotation) via an adaptor 260. Solutions were electrospun via pumping at 450 μL/h through a 20 G blunt-tip needle with an applied voltage of 17 kV, at a distance of 13 cm from a set of parallel grounded collectors to form parallel nanofibers. One collector was then rotated clockwise for a specified number of rotations (twists) prior to removal of the suture from the collectors and storage at −20° C.
When a charged polymer jet was ejected from the needle, it deposited in the air gap between both collectors 204, 202. As the polymer solution continued to be ejected, hundreds of parallel fibers were formed with one end attached to the drill chuck and the other end attached to the standalone parallel stand.
Next, the drill chuck 270 was rotated about an axis defined by collectors 204, 202, to twist the parallel fibers into 28 μm-thick in diameter and having 1,575 twists.
The breaking strengths of these sutures made with different polymers were evaluated. The morphology was examined under scanning electron microscopy (SEM).
Suture diameter was determined via light microscopy using the 20× objective of an Eclipse TS100 (Nikon Instruments, Melville, N.Y.) and calibrated Spot 5.2 Basic imaging software (Spot Imaging, Sterling Heights, Mich.). Each suture was measured at three different locations at least 2 cm apart, and used in additional experimentation only if the average diameter was within ±0.5 μm of the specified diameter.
Suture morphology was observed via SEM at 1 kV using a LEO Field Emission SEM (Zeiss, Oberkochen, Germany). Prior to imaging, samples were desiccated and then sputter coated with 10 nm of Au/Pd (Desk II, Denton Vacuum, Moorestown, N.J.).
Sutures (n=3-4 for each condition) were clamped vertically and then pulled until breaking at a rate of 2.26 mm/min using a DMA 6800 (TA Instruments, Timonium, Md.).
Results
Hundreds of nano-fibers were twisted in one direction and tightly packed. As shown in
Multifilament, drug-loaded sutures were cylindrical and met U.S.P. specifications for 10-0 suture diameter (20-29 μm), making them suitable in size for ocular surgery. SEM images showed they were also comparable in both size and shape to commercially available 10-0 Ethilon® (nylon) sutures.
This study was done with monofilament sutures for comparison with the twisted multi-filament structures described herein.
Materials and Methods
As shown in
Poly (L-lactic acid) (PLLA) solutions containing levofloxacin and different amounts of polyethylene glycol (PEG) at 1%, 2%, or 4%, or 2% by weight PLURONIC® F127 were electrospun. Briefly, PLLA (221 kDa; Corbion, Amsterdam, Netherlands) at 86-89% (w/w) was mixed with levofloxacin (Sigma Aldrich, St. Louis, Mo.) at 10 wt % and either PEG (35 kDa, Sigma Aldrich) or PLURONIC® F127 (BASF, Florham Park, 73 NJ) between 1-4 wt % and dissolved in chloroform (Sigma Aldrich) at room temperature for 24 h. Levofloxacin concentration was held constant and PLLA concentration in chloroform was maintained at 15 wt % in all formulations. Sutures were produced by wet electrospinning the polymer/drug solution in a setup consisting of a high voltage power supply (Gamma High Voltage Research, Ormond Beach, Fla.), syringe pump (Fisher Scientific, Waltham, Mass.), and rotating metal collector with hexane (Sigma Aldrich) as the lending solvent. The polymer solution was ejected through a blunted 18G needle (Fisher Scientific) at 13 mL/h with 4.7 kV of applied voltage 5 cm away from the collector rotating at 40 rpm. Fibers were then collected and desiccated for two days prior to storage at −20° C.
15% PCL with 8% Levo was also used at a flow rate of 1 mL/hr and 28 μm in diameter using the same setup.
Following suture manufacture, fibers were desiccated and stored at −20° C. preceding use in additional experiments. Prior to tensile testing, sutures were allowed to fully thaw and were cut into 3 cm segments.
For suture morphology and size assessment, sutures were serially dehydrated in ethanol (Sigma Aldrich) and dried prior to sputter coating with 10 nm of Au/Pd. Samples were then imaged via scanning electron microscopy (SEM) at 1-2 kV using a LEO Field Emission SEM (Zeiss, Oberkochen, Germany) and suture diameter measured using ImageJ software (n=14 for each condition).
For tensile strength measurement, mechanical properties of the sutures were evaluated using a DMA 6800 (TA Instruments, Timonium, Md.). 3 cm long samples (n=7 for each condition) were clamped vertically and force from a 5 N load cell was applied at 0.05 N/min to stretch the sample until breaking.
For in vitro drug release, 10 mg of suture (n=3) was placed into 10 mL of 1× Dulbecco's Phosphate Buffered Saline (PBS, ATCC, Manassas, Va.) rotating at 37° C. At each time point, 2 mL aliquots were withdrawn and replaced with fresh PBS. Aliquots were frozen, lyophilized, and resuspended in ultrapure water prior to high performance liquid chromatography (HPLC; Waters Corporation, Milford, Mass.) analysis. 100 μL samples were injected into a Waters Symmetry® 300 C18 5 μm column with a mobile phase of 0.1% v/v trifluoroacetic acid (Sigma Aldrich) in water:acetonitrile (75:25 v/v, 98 Fisher Scientific) at a flow rate 1 mL/min. Elution was monitored by a 2998 photodiode array detector to detect levofloxacin with excitation at 290 nm and emission at 502 nm. Drug loading was determined by dissolving a 5 mg sample of suture into a mixture of tetrahydrofuran (Sigma Aldrich):acetonitrile (20:80) and injecting into the column under the same conditions as the release samples.
For assessment of bacterial inhibition, 1 cm of suture was placed in 1 mL of PBS and incubated at 37° C. for 1, 3, and 6 h and 1, 2, 3, 4, 5, 6, and 7 days (n=6 for each time point). S. epidermidis (ATCC) was cultured overnight at 37° C. on agar plates produced using nutrient agar (BD, Franklin Lakes, N.J.). At each time point, sutures were retrieved and placed on plated cultures in order to investigate bacterial inhibition. Bacterial inhibition zones around the sutures were measured and imaged 24 h after suture placement.
For assessment of in vivo biocompatibility, animals were cared for and experiments conducted in accordance with protocols approved by the Animal Care and Use Committee of the Johns Hopkins University. Protocols are also in accordance with the ARVO Statement for the Use of Animals in Ophthalmic and Vision Research. 1 mm of 8-0 Ethilon® (nylon), Vicryl® (poly(lactic-co-glycolic acid); PLGA) (Ethicon, Somerville, N.J.) and 4% PEG/PLLA/levofloxacin sutures (n=3) were implanted into the corneas of 6-8 weeks old, male Sprague-Dawley rats (Harlan Laboratories, Frederick, Md.). Prior to implantation, rats were intraperitoneally anesthetized with a solution of Ketamine:Xylazine (75:5 mg/kg, Sigma Aldrich) and a drop of 0.5% proparacaine hydrochloride ophthalmic solution (Bausch & Lomb Inc., Tampa, Fla.) was applied to the cornea. Following implantation, the rats were evaluated for signs of infection every day for seven days. The rats were then euthanized and eyes enucleated, fixed in formalin (Sigma Aldrich) for 24 h, embedded in paraffin, cross sectioned, and stained with hematoxylin and eosin for histological evaluation.
Results
Electro spinning of a 10 wt % polymer solution with application of 4.7 kV into a collector containing hexane and rotating at 40 rpm (as shown in
As shown in
Although thin enough to qualify as a 9-0, 8-0, or 7-0 suture, the monofilament suture was unlike multifilament sutures in Examples 1, 3, and 4, the latter ones of which satisfied the clinical strength requirement.
As shown in
Preliminary studies indicated minimal drug release from PLLA/levofloxacin sutures manufactured via electro spinning. However, the addition of small percentages of PLURONIC® F127 and PEG polymers to the formulation resulted in significant and sustained release of levofloxacin in vitro (
Bacterial inhibition zone experiments were conducted with S. epidermidis to determine whether levofloxacin released from sutures was capable of eliminating bacteria in an in vitro setting, and how long this effect might last in vivo. 4% PEG/PLLA/levofloxacin sutures were cut to 1 cm in length and incubated in 37° C. PBS from 1 h up to 7 days. After each time point, the suture was removed from solution and placed in the center of an agar plate that had been cultured with S. epidermidis for 24 h. PBS, neat drug, and 4% PEG/PLLA sutures were used as controls. Results of bacterial culture indicated PBS and 4% PEG/PLLA did not inhibit bacterial growth, while the 4% PEG/PLLA/levofloxacin suture created a 2 cm inhibition zone after 24 h of drug release in PBS.
In order to evaluate the potential clinical value of an absorbable, antibiotic-eluting suture, wet electrospun sutures were implanted into the corneal stroma of male Sprague Dawley rats. 8-0 Ethilon®, 8-0 Vicryl®, and 8-0 4% PEG/PLLA/levofloxacin sutures of approximately 1 mm in length were compared to each other and untreated controls after 7 days. Notably, 4% PEG/PLLA/levofloxacin sutures remained in the cornea and maintained integrity through the 7 day period, similar to the ETHILONEthilon® and VICRYLVicryl® sutures. Rats were monitored daily, and there were no gross signs of infection or inflammation among any of the animals for all sutures tested. Histological analysis showed that the tissue reaction to the electrospun 4% PEG/PLLA/levofloxacin suture was indistinguishable to that of the nylon suture and untreated controls. There were no obvious signs of neovascularization or inflammation in the control, nylon, or antibiotic-eluting suture conditions. However, immune cell infiltration was apparent in each of the rat eyes containing a Vicryl® suture.
A key challenge for translation of drug-loaded sutures to the clinic has been an inability to meet U.S.P. specifications for suture strength. Thus, the impacts of fiber conformation, drug concentration, drug type, and diameter on antibiotic-loaded suture breaking strength were examined.
Materials and Methods
The multifilament sutures were prepared using the setup as shown in
Strength retention test: PCL/8% Levo and PCL/16% Levo sutures (n=5) were sectioned into two halves. The breaking strength of one segment was measured as described in Example 1, while the other segment was submerged in 1× Dulbecco's Phosphate Buffered Saline 360 (ATCC, Manassas, Va.) and shaken at 225 rpm at 37° C. for 31 days. Sutures were then dried prior to measuring breaking strength.
Results
As shown in
As shown in
As shown in
So far, compared with the monofilaments in Example 3 and
As shown in
Importantly, both PCL/8% Levo and PCL/16% Levo sutures maintained their strengths and demonstrated minimal degradation in vitro over a period of 31 days in phosphate buffered saline (PBS), as shown in Table 3.
As shown in
As shown in
Materials and Methods
As shown in
A 10-0, 9-0, or 8-0 nylon suture 390 was fixed on its thread end to the drill chuck 370. The needle end of the suture was held by the parallel stand 302 where the suture was placed through the stand 302 but this end of the suture was kept free to rotate.
After the charged polymer/drug solution was released, hundreds of the charged polymer/drug jet deposited between the chuck 370 and the parallel stand 302, surrounding the suture. Due to the electric charge, the fibers are held tightly by the chuck and the parallel stand. Then, the chuck was rotated clockwise to twist these fibers, with the suture “buried” among the fiber Later the chuck was rotated counterclockwise to ensure that the suture did not coil or snap, while the coating remained tight and intact.
Results
As confirmed using SEM, an uncoated 10-0 nylon suture had a smooth surface and a diameter of approximately 25 μm. The coated 10-0 nylon suture had hundreds or thousands of nanofibers covering the surface of the suture in a compacted, spiral manner. The coating only added about 2 μm to about 5 μm to the overall thickness, continuing to meet the 10-0 suture size and strength requirement.
Overall, nanofiber-coated sutures allow for tunable drug release and loading without affecting suture breaking strength. Coating of 10-0 nylon sutures were demonstrated to add a specific amount of fiber coating thickness to the sutures: adding between 3 and 5 μm of fiber coating kept the USP size at 10-0; adding between about 10 and 15 μm of fiber coating increased the size to USP 9-0; and adding between about 20 and 20 μm of fiber coating increased the size to USP 8-0. For the rapamycin release in
Materials and Methods
The biocompatibility of antibiotic sutures was evaluated by implanting 2 mm long sutures in 6-8 week old, male Sprague-Dawley rat cornea on day 0 and enucleation and fixing on day 2 for histological analysis. No bacterial inoculation was administered in this biocompatibility study. The implanted sutures and treatment included 10-0 (i) VICRYL® suture; (ii) nylon suture; (iii) nylon suture that was coated with PCL multifilament fibers containing 8% levofloxacin (PCL/8% Levo/Nylon), (iv) multifilament composite as a suture, made from electrospun from PCL solution; (v) multifilament composite as a suture, made from electrospun from PCL solution containing 8% levofloxacin (PCL/8% Levo); and (vi) multifilament composite as a suture, made from electrospun from PCL solution containing 16% levofloxacin (PCL/16% Levo). Prior to implantation, rats were intraperitoneally anesthetized with a solution of ketamine:xylazine (75:5 mg/kg, Sigma Aldrich) and a drop of 0.5% proparacaine hydrochloride ophthalmic solution (Bausch & Lomb Inc., Tampa, Fla.) was applied to the cornea. Following implantation, the rats were evaluated daily for signs of infection, inflammation, or irritation. Two days after implantation, the rats were euthanized and eyes enucleated, fixed in formalin (Sigma Aldrich) for 24 h, embedded in paraffin, cross sectioned, and stained with H&E for histological evaluation.
Next, two models of bacterial infections were evaluated in rat cornea injury with suture implantation:
In the first, 2-day model, the cornea of Sprague-Dawley rats were scratched, followed by implantation of sutures and administration of 100 μL of Staphylococcus Aureus at 1×108 CFU/mL on day 0. Cornea without implantation or bacteria administration was used as a control (untreated). The implanted sutures and treatment included (i) 2 mm VICRYL® suture, no antibiotic; (ii) 2 mm 10-0 nylon suture, no antibiotic; (iii) 2 mm 10-0 nylon suture and a single drop of levofloxacin (10 μL of 5 mg/mL, i.e., 0.5%, levofloxacin solution), immediately following implantation of the suture; (iv) 2 mm 10-0 nylon suture and a daily levofloxacin (three 10 μL drops of 5 mg/mL, i.e., 0.5%, levofloxacin solution each day); (v) 2 mm 10-0 nylon suture that was coated with PCL multifilament fibers containing 8% levofloxacin (PCL/8% Levo/Nylon), and (vi) multifilament composite as a suture, made from electrospun from PCL solution containing 8% levofloxacin (PCL/8% Levo). At 48 hr after implantation, the following procedures were performed: either (a) enucleated the treated eye, removed and homogenized the cornea and measured bacterial concentration using a plate reader at 600 OD; (b) observed bacterial growth via the streaking method by using a sterile swab to wick the top of the rat eye and cultured on agar plates overnight at 37° C.; or (c) enucleated the treated eye, embedded in paraffin, sectioned, and stained with hematoxylin and eosin.
In the second, 7-day model, S. Aureus was re-administered on day 5 in addition to the first administration on day 0 as described above to evaluate the capacity of the 10-0 grade multifilament sutures containing either 8% levofloxacin or 16% levofloxacin, implanted on day 0, to continue to prevent ocular infection following the immediate post-operative period. On day 7, swabs were taken of each cornea followed by either histological evaluation, bacterial homogenization, or removal of sutures for examination via SEM (n=4 for each condition). For the latter experiment, sutures were removed from the cornea and fixed in formalin (Sigma-Aldrich) for 30 min prior to washing with PBS and dehydration with increasing concentrations of ethanol (Fisher Scientific). Sutures were then imaged.
All animals were cared for and experiments conducted in accordance with protocols approved by the Animal Care and Use Committee of the Johns Hopkins University. Protocols were also in accordance with the ARVO Statement for the Use of Animals in Ophthalmic and Vision Research.
Bacterial Inoculation and Evaluation in Detail:
Sprague Dawley rats were anesthetized. The operative eye was then scratched using a 20 G needle (Fisher Scientific) prior to implantation of three 2 mm long nylon (n=12), Vicryl® (n=4), or PCL/8% Levo (n 403=4) suture filaments. 100 μL of 1×108 CFU/mL of S. aureus was then administered topically over a period of 10 mins. 10 μL of 0.5% levofloxacin solution was administered topically either once post-operatively or three times daily to rat eyes with nylon sutures (n=4, each). Two days after implantation, gross images were taken of each eye, prior to swabbing the cornea with a cotton-tipped applicator (Fisher Scientific), and streaking onto tryptic soy agar (Fisher Scientific) plates. Plates were stored in an incubator at 37° C. for 24 h and then imaged. After swabbing the eye, eyes were enucleated and either prepared for histological evaluation (n=3 for each condition) or evaluated for bacterial load (n=4 for each condition). Briefly, each eye was placed in sterile tryptic soy broth (Fisher Scientific) and homogenized using a POWER GEN® 125 homogenizer (Fisher Scientific) for 4 min. Samples were then centrifuged at 300 rcf for 5 min, and optical density of the supernatant measured at a wavelength of 600 nm via spectrophotometry. Infection was confirmed by a positive swab culture and bacterial load significantly higher than a control eye.
Results
In the biocompatibility study without bacterial inoculation to the eye, histological analysis of tissues (Day 2) surrounding the sutures implanted in rat cornea showed all antibiotic-loaded sutures were biocompatible, and did not elicit an influx of innate immune cells to the site of suture implantation. There were no gross signs of irritation, inflammation, or infection among any of the treated or control groups for the duration of the study. Histological analysis further revealed that implantation of PCL or PCL/Levo sutures did not cause neovascularization, and that the tissue reaction was comparable to commercially available nylon sutures. Notably, a small ring of cells was observed surrounding implanted absorbable Vicryl® sutures.
In the first 2-day model with bacterial inoculation, hematoxylin and eosin (H&E) staining revealed substantial inflammation and cellular infiltration within the corneas of rats receiving implantation of VICRYLVicryl® or nylon sutures without post-operative administration of Levo.
Notably, the concentration of cells was greatest within the immediate vicinity of implanted sutures lacking the antibiotic loading, which was indicative that the suture itself may be the nidus of infection and location of bacterial adherence. Cells were also concentrated around nylon sutures implanted in rat eyes receiving a single post-operative dose of Levo. However, there was no sign of infection or inflammation in the corneal tissue surrounding PCL/8% Levo sutures or nylon sutures in rats receiving three daily doses of Levo, and the tissue resembled that of a healthy control. Culture of bacterial swabs on agar plates similarly confirmed the presence of infection in rats with implantation of VICRYLVicryl® or nylon sutures, or nylon sutures followed by a single dose of Levo administered topically.
As shown in
In the second, 7-day model where S. Aureus was re-inoculated on day 5,
Materials & Methods
In order to determine the duration and concentration of Levo delivery from sutures in vivo, a pharmacokinetic study was performed by implanting three 2-mm-in-length 28 μm PCL/8% Levo and PCL/16% Levo sutures into rat corneas.
PCL/8% Levo and PCL/16% Levo sutures were implanted into Sprague Dawley rat corneas as described above (n=4 for each formulation at each time point). At 15, 60, and 120 min, and at 1, 3, 7, and 14 days, aqueous humor was collected from each eye, followed by removal of implanted sutures and harvesting of the cornea. Tissue and aqueous humor samples were weighed immediately after harvesting. Corneal tissue samples were homogenized in 100 μL to 150 μL of PBS prior to extraction. The standard curve and quality control samples were prepared in PBS as a surrogate matrix for both aqueous humor and homogenized tissue. Levofloxacin was extracted from 15 μL of aqueous humor or tissue homogenate with 50 μL of acetonitrile containing 1 μg/mL of the internal standard, moxifloxacin-d4 (Toronto Research Chemicals, Canada). After centrifugation, the supernatant was then transferred into autosampler vials for LCMS/MS analysis. Separation was achieved with an AGILENT ZORBAX® Agilent Zorbax XDB-C18 (4.6×50 mm, 5 μm) column with water/acetonitrile mobile phase (40:60, v:v) containing 0.1% formic acid using isocratic flow at 0.3 mL/min for 3 minutes. The column effluent was monitored using an AB SCIEX® Sciextriple Quadrupole™ 5500 mass-spectrometric detector (Sciex, Foster City, Calif.) using electrospray ionization operating in positive mode. The spectrometer was programmed to monitor the following MRM transitions: 362.0→318.0 for levofloxacin and 406.1→108.0 for the internal standard, moxifloxacin-d4. Calibration curves for levofloxacin were computed using the area ratio peak of the analysis to the internal standard by using a quadratic equation with a 1/×2 weighting function using two different calibration ranges of 0.25 to 500 ng/mL with dilutions up to 1:10 399 (v:v) and 5-5,000 ng/mL.
Results
As shown in Table 4, analysis of Levo concentration in harvested aqueous humor and corneas revealed a burst release of antibiotic following suture implantation and for multiple hours afterwards. The Levo release profiles were similar in eyes implanted with either 8% or 16% Levo sutures. However, eyes with PCL/16% Levo sutures contained higher concentrations of Levo in both the aqueous humor and cornea at almost all time points. Sutures maintained their location and macroscopic structure throughout the course of the study, and in both 8% and 16% Levo conditions, Levo was detected in the cornea and aqueous humor 14 days after implantation. HPLC analysis of dissolved sutures revealed Levo loading of 80 and 161 μg/m, respectively, for 8% and 16% Levo sutures. A burst release of antibiotics may be important for prevention of immediate post-operative infection when wounds or surgical incisions are most vulnerable to bacterial infiltration. Herein, local antibiotic delivery from drug-loaded sutures may preclude issues of patient compliance with topical eye drops, prevent suture-related infections that lead to treatment failure and re-intervention, reduce the need for oral antibiotic use, decrease the risk of infection associated with implantable ocular devices, and serve as an alternative to the more than 12 million nylon sutures used in ocular procedures each year.
In vitro levofloxacin release assay from multifilament sutures showed drug release profile did not change as PCT/Levo suture size increased (Table 5) and sustained release was achieved with PLLA/Levo suture (Table 6).
Materials and Methods
The abdominal aorta of a rat was sectioned and interrupted suturing was performed to tie the vessels back together. The sutures used in this procedure included (i) 8-0 nylon suture, (ii) nylon suture coated with PLLA/PEG containing 20% rapamycin (8-0), or (iii) nylon suture coated with PLLA/PEG containing 40% rapamycin (8-0). Following the coating process, the size of the sutures (ii) and (iii) was within the 8-0 size range. Overall suture diameter was increased for about 20 μm with the addition of fiber coating, i.e., suture diameter increased from 10-0 to 8-0 classification. After two weeks, the aorta was harvested, sectioned and stained for histological analysis, where the neointimal hyperplasia formation was quantified.
Suture Fabrication:
Polymer solutions were made via dissolution of PLLA (221 kDa; Corbion, Amsterdam, Netherlands), PEG (35 kDa; Sigma Aldrich, St. Louis, Mo.), and rapamycin (LC Laboratories, Woburn, Mass.) in hexafluoroisopropanol (Sigma-Aldrich) by shaking overnight at room temperature. Polymer to solvent concentration was maintained at 10.8% and PEG to PLLA concentration was maintained at 3.9% for all formulations. Rapamycin concentration was 20%, 40%, or 80% in regard to PLLA for the 20% Rap/PLLA/PEG, 40% Rap/PLLA/PEG, and 80% Rap/PLLA/PEG formulations, respectively. Prior to electrospinning, the non-needled end of a 10-0 nylon suture (Ethicon, Somerville, N.J. or AROSurgical Instruments, Newport, Calif.) was placed into the rotational collector. The needled end was driven through the hole in the opposing collector and allowed to hang loosely. Rap/PLLA/PEG solutions were then pumped through a 20 G blunt-tip needle at 1 mL/h with an applied voltage of 15 kV at a distance of 17 cm from the parallel, grounded collectors. The collector containing the non-needled end of the suture was then rotated clockwise for five minutes at 150 rpm and counter-clockwise for 30 s at an identical speed. The suture was then removed from the collectors and stored at −20° C.
Suture diameter was determined via light microscopy (Eclipse TS100, Nikon Instruments, Melville, N.Y.) and calibrated imaging software (Spot 5.2 Basic, Spot Imaging, Sterling Heights, Mich.). Each suture was measured at four different locations at least 2 cm apart, and used in further experimentation only if the average diameter was between 46 and 49 μm, qualifying as an 8-0 suture.
Both quantitatively and qualitatively, increasing rapamycin concentration significantly decreased the formation of neointimal hyperplasia and the potential occlusion of the vessel. A degradable, drug loaded polymer fiber coating on a traditional nylon suture could reduce neointimal hyperplasia formation in vascular anastomosis surgeries.
This application claims benefit of and priority to U.S. Provisional Application Nos. 62/307,230 and 62/307,096, both filed on Mar. 11, 2016, which are hereby incorporated herein by reference in their entirety.
This invention was made with government support under grant number BGE-1232825, awarded by the National Science Foundation. The government has certain rights in the invention.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2017/022093 | 3/13/2017 | WO | 00 |
| Number | Date | Country | |
|---|---|---|---|
| 62307096 | Mar 2016 | US | |
| 62307230 | Mar 2016 | US |
