The present invention relates to the field of biological tissue repair and/or wound closure, e.g., after injury to the tissue or surgery. More particularly, the present invention relates to the use of biological or biocompatible adhesive composites for the repair of biological tissue.
Known methods of biological tissue repair include sutures, staples and clips, sealants, and adhesives. Sutures are inexpensive, reliable, readily available and can be used on many types of lacerations and incisions. However, the use of sutures has many drawbacks. Sutures are intrusive in that they require puncturing of the tissue. Also, sutures require technical skill for their application, they can result in uneven healing, and they often necessitate patient follow-up visits for their removal. In addition, placement and removal of sutures in children may require sedation or anesthesia.
Staples or clips are preferred over sutures, for example, in minimally invasive endoscopic applications. Staples and clips require less time to apply than sutures, are available in different materials to suit different applications, and generally achieve uniform results. However, staples and clips are not easily adapted to different tissue dimensions and maintaining precision of alignment of the tissue is difficult because of the relatively large force required for application. Further, none of these fasteners is capable of producing a watertight seal for the repair.
Sealants, including fibrin-, collagen-, synthetic polymer- and protein-based sealants, act as a physical barrier to fluid and air, and can be used to promote wound healing, tissue regeneration and clot formation. However, sealants are generally time-consuming to prepare and apply. Also, with fibrin-based sealants, there is a risk of blood-borne viral disease transmission. Further, sealants cannot be used in high-tension areas.
Adhesives, for example, cyanoacrylate glues, have the advantage that they are generally easy to dispense. However, application of adhesives during the procedure can be cumbersome. Because of their liquid nature, these adhesives are difficult to precisely position on tissue and thus require adept and delicate application if precise positioning is desired. Cyanoacrylates also harden rapidly; therefore, the time available to the surgeon for proper tissue alignment is limited. Further, when cyanoacrylates dry, they become brittle. Thus, they cannot be used in areas of the body that have frequent movement. In addition, the currently available adhesives are not optimal for high-tension areas.
Laser tissue solders, or “light-activated adhesives,” are a possible alternative for overcoming the problems associated with the above-mentioned techniques. Laser tissue soldering is a bonding technique in which a protein solder is applied to the surface of the tissue(s) to be joined and laser energy is used to bond the solder to the tissue surface(s).
The use of biodegradable polymer scaffolding in laser-solder tissue repairs has been shown to improve the success rate and consistency of such repairs. See, for example, McNally et al., U.S. Pat. No. 6,391,049. However, a drawback of laser-soldering techniques is the need to supply light energy to the repair site to activate the adhesive. As a result, such techniques are only suitable for a limited number of clinical applications. For example, such techniques are generally not suitable for use outside of a hospital or other laser-equipped setting. Also, with laser techniques, there is always a risk of collateral thermal damage to the surrounding tissue.
Accordingly, there is a need for an improved method of biological tissue repair; particularly, a device or surgical product, system, and/or method which is capable of replacing the conventional suture, staple and clip techniques in a wide variety of applications.
A novel biocompatible or biological adhesive composite that results from the combination of a non-light activated adhesive and a scaffold material has been invented. This composite has exhibited surprisingly good tensile strength and consistency when compared with sutures and the use of adhesives alone. It can be used effectively as an adhesive, sealing or repairing device for biological tissue. It may also be used as a depot for drugs in providing medication to a wound or repair site. The composite can be precisely positioned across, on top of, or between two materials to be joined (i.e. tissue-to-tissue or tissue-to-biocompatible implant). Proper alignment is accomplished within the time period before the adhesive sets or hardens. Thus, the composite can be applied to a repair site more quickly and easily than sutures or adhesives alone. In addition, application of the composite can provide a watertight seal at the repair site when required.
The improved ease of clinical application makes the composite of the present invention applicable to all internal and external fields of surgery, extending from emergency neurosurgical and trauma procedures to elective cosmetic surgery, as well as to ophthalmic applications. Examples of external or topical applications for the composite include, but are not limited to, wound closure from trauma or at surgical incision sites. Internal surgical applications include, but are not limited to, repair of liver, spleen, or pancreas lacerations from trauma, dural laceration/incision closure, pneumothorax repair during thoracotomy, sealing points of vascular access following endovascular procedures, vascular anastomoses, tympanoplasty, endoscopic treatment of gastrointestinal ulcers/bleeds, dental applications for mucosal ulcerations or splinting of injured teeth, ophthalmologic surgeries, tendon and ligament repair in orthopedics, episiotomy/vaginal tear repair in gynecology. Additionally, as minimally invasive techniques become more common, the application of this technology to endoscopic, laparoscopic or endovascular techniques is very promising. With appropriate single-use packaging, the invention offers the potential for quick application in the field by less skilled professionals, paraprofessionals and bystanders in emergency situations—both military and civilian—outside a hospital or clinic setting.
Various techniques for forming the composite of the present invention and/or applying it to a wound or tissue repair site may be used. Additionally, there are numerous suitable alternatives for packaging the composite depending on the desired use, environment, or applications.
In accordance with the present invention, a composition suitable for medical and surgical applications is provided. The composition includes a scaffold including at least one of a biological material, biocompatible material, and biodegradable material, and a non-light activated adhesive including at least one of a biological material, biocompatible material, and biodegradable material. The non-light activated adhesive is combined with the scaffold to form a composite that, when used to repair biological tissue, has a tensile strength of at least about 120% of the tensile strength of the adhesive alone.
Also in accordance with the present invention, a method for repairing, joining, aligning, or sealing biological tissue is provided. The method includes the steps of combining a biological, biocompatible, or biodegradable scaffold and a non-light activated biological, biocompatible, or biodegradable adhesive to form a composite having a tensile strength of at least about 120% of the tensile strength of the adhesive alone, and applying the composite to an adhesion site.
Yet further in accordance with the present invention, a product for joining, repairing, aligning or sealing biological tissue is provided. The product includes a biological, biocompatible, or biodegradable scaffold, a biological, biocompatible, or biodegradable non-light activated adhesive, and means for coupling the scaffold and the adhesive to form a composite having a tensile strength of at least about 120% of the tensile strength of the adhesive alone.
Several experimental studies have confirmed the effectiveness of the present composite, which comprises a non-light activated adhesive and a scaffold, for biological tissue repair. The attached Appendix, incorporated herein by this reference, includes data tables relating to these studies. While specific compounds have been used in these studies, it is understood that the composite of the present invention is not limited to the particular compounds used in any of the disclosed examples.
The scaffold and adhesive used to form the composite of the present invention may each be composed of either biologic or synthetic materials. Examples of biologic materials that may be used as adhesives include, but are not limited to, serum albumin, collagen, fibrin, fibrinogen, fibronectin, thrombin, barnacle glues and marine algae. Examples of synthetic materials suitable for use as adhesives include, but are not limited to, cyanoacrylate (e.g., ethyl-, propyl-, butyl- and octyl-) glues. The biologic materials are, by their very nature, biodegradable. Currently marketed synthetic adhesives such as cyanoacrylates are not in themselves biodegradable, but processes can be applied to make them biodegradable. For example, a formaldehyde-scavenging process can be applied that allows the product to degrade in the body without producing a toxic reaction.
The mechanism by which the adhesive material bonds to the tissue, and thus, the determination of whether any auxiliary equipment is necessary, is dependent at least in part on the selection of the adhesive material. Some non-light activated adhesives require an activator or initiator (other than laser energy) to cause or accelerate bonding. For example, polymerization of octyl-cyanoacrylates can be accelerated through contact with a chemical initiator such as that contained in the tip of the applicator of Ethicon's Dermabond™. Cohesion's CoStasis and Cryolife's Bioglue also rely on the addition of an activator at the time of application, namely, fibrinogen and glutaraldehyde, respectively. It is understood that all of the above-mentioned adhesives, whether or not they require an initiator or activator, are considered “non-light activated” adhesives.
The scaffold operates to ensure continuous, consistent alignment of the apposed tissue edges. The scaffold also helps ensure that the tensile strength of the apposed edges is sufficient for healing to occur without the use of sutures, staples, clips, or other mechanical closures or means of reinforcement. By keeping the tissue edges in direct apposition, the scaffold helps foster primary intention healing and direct re-apposition internally. Thus, the scaffold functions as a bridge or framework for the apposed edges of severed tissue.
As mentioned above, the scaffold is either a synthetic or biological material. A suitable biological scaffold comprises SIS (small intestine submucosa), polymerized collagen, polymerized elastin, or other similarly suitable biological materials. Examples of synthetic materials suitable for use as a scaffold include, but are not limited to, various poly(alpha ester)s such as poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(L-lactic-co-glycolic acid) (PLGA), poly(.epsilon.-caprolactone) (PGA) and poly(ethylene glycol) (PEG), as well as poly(alpha ester)s, poly(ortho ester)s and poly(anhydrides).
In alternative embodiments, the scaffold is engineered for specific applications of the composite by adjusting one or more of its properties. For example, the scaffold includes a smooth surface. Alternatively or in addition, the scaffold includes an irregular surface. Key properties of the scaffold are surface regularity or irregularity, elasticity, strength, porosity, surface area, degradation rate, and flexibility.
For purposes of this disclosure, “irregular” means that at least a portion of a surface of the scaffold is discontinuous or uneven, whether due to inherent porosity, roughness or other irregularities, or as a result of custom-engineering performed to introduce irregularities or roughness onto the surface (for example, using drilling, punching, or molding manufacturing techniques).
In further embodiments of the present invention, the scaffold is engineered to allow it to function as a depot for various dopants or biologically-active materials, such as antibiotics, anesthetics, anti-inflammatories, bacteriostatic or bacteriocidals, chemotherapeutic agents, vitamins, anti- or pro- neovascular or tissue cell growth factors, hemostatic and thrombogenic agents. This is accomplished by altering the macromolecular structure of the scaffold in order to adjust, for example, its porosity and/or degradation rate.
An ex vivo study was conducted to compare the tensile strength of tissue samples repaired using three different techniques: (i) application of a scaffold-enhanced light-activated albumin protein solder (Group I), (ii) application of a scaffold-enhanced n-butyl-cyanoacrylate (non-light activated) adhesive composite (Group II), and (iii) repair via conventional suture technique (Group III).
1.1 Preparation of the Surgical Adhesive
Porous synthetic polymer scaffolds were prepared from poly(L-lactic-co-glycolic acid) (PLGA), with a lactic:glycolic acid ratio of 85:15, using a solvent-casting and particulate leaching technique. The scaffolds were cast by dissolving 200 mg PLGA (Sigma Chemical Company, St. Louis, Mo.) in 2 mL dichloromethane (Sigma Chemical Company). Sodium chloride (salt particle size: 106-150 nm) with a 70% weight fraction was added to the polymer mix. The polymer solution was then spread to cover the bottom surface of a 60 mm diameter Petri dish that was cleaned first with dichloromethane, then ethanol, then ultra-filtered deionized water (Fisher Scientific, Pittsburgh, Pa.). The polymer was left in a fume hood for 24 hours to allow the dichloromethane to evaporate. The salt was leached out of the polymer scaffolds by immersion in filtered deionized water for 24 hours, to create the porous scaffolds. During this period the water was changed 3-4 times. The scaffolds were then air dried and stored at room temperature until required.
The PLGA scaffolds used for incision repair were cut into rectangular pieces with dimensions of 12±2 mm long by 5±1 mm wide. The scaffolds used for Group I were left to soak for a minimum of two hours before use in a protein solder mix comprised of 50% (w/v) bovine serum albumin (BSA) (Cohn Fraction V, Sigma Chemical Company) and Indocyanine Green (ICG) dye (Sigma Chemical Company) at a concentration of 0.5 mg/mL, mixed in deionized water. The thickness of the resulting scaffold-enhanced solders, determined by scanning electron microscopy and measurement with precision calipers (L. S. Starrett Co., Anthol, Mass.), was in the range of 200 to 205 μm. N-butyl-cyanoacrylate (Vetbond, 3M) was applied to the scaffolds used for Group II using a 22-G syringe immediately prior to application to the tissue.
1.2 Tissue Preparation
Porcine tissue specimens were harvested approximately 30 minutes after sacrificing the animals. Tissue specimens were stored in phosphate buffered saline for a maximum of two hours before they were prepared for experiments. Each tissue specimen was cut into small rectangular pieces with dimensions of about 2 cm long by 1 cm wide and a thickness of approximately 1.5±0.5 mm. Tissue specimens harvested included the small intestine, spleen, muscle, skin, atrium, ventricle, lung, pancreas, liver, gall bladder, kidney, ureter, sciatic nerve, carotid artery, femoral artery, splenic artery, coronary artery, pulmonary artery and aorta (both intima and adventitia). Ten repairs were performed for each tissue type and repair procedure investigated.
1.3 Incision Repair
A full thickness incision was cut through each specimen width using a scalpel, and opposing ends were placed together. All laser-assisted repairs were completed with a diode laser operating at a wavelength of 808-nm (Spectra Physics, Mountain View, Ca.). The laser light was coupled into a 660-μm diameter silica fiber bundle and focused onto the scaffold surface with an imaging hand-piece connected at the end of the fiber. The diode was operated in continuous mode with a spot size of approximately 1 mm at the surface of the scaffold-enhanced solder. An aiming beam was also incorporated into the system and was delivered through the same fiber as the 808-nm beam. The laser beam was scanned across the scaffold-enhanced solder twice, starting from the center and moving outwards in a spiral pattern with a total irradiation time of 80±2 seconds. Suture repairs were achieved using a single 4-0 nylon suture.
1.4 Strength Testing
Tensile strength measurements were performed to test the integrity of the resultant repairs immediately following the laser procedure using a calibrated MTS Material Strength Testing Machine (858 Table Top System, MTS, Eden Prairie, Minn.). This system was interfaced with a personal computer to collect the data. Each tissue specimen was clamped to the strength testing machine using a 100N load cell with pneumatic grips. The specimens were pulled apart at a rate of 100 gf/min until the repair failed. Complete separation at the tissue edges defined failure. The maximum load in Newton's was recorded at the breaking point. The strengths of corresponding native specimens were tested and used as references. Native tissue specimens were prepared for tensile testing in an identical manner to the experimental repair group specimens, with the exception that microscissors were used to cut in from each edge with care to leave a 5±1 mm bridge of tissue in the center. This spacing matched the width of the scaffold-enhanced adhesives used on specimens from Groups I and II.
1.5 Results
The tensile strengths recorded at the breaking point of the repaired organ specimens are recorded in Table 1 and displayed in
Group I repairs formed on the ureter were the most successful followed by the small intestine, sciatic nerve, spleen, atrium, kidney, muscle, skin and ventricle. The repairs on the ureter, small intestine and sciatic nerve achieved 81-83% of the strength of native tissue while repairs on the spleen, atrium and kidney attained approximately 66-72% of the strength of native tissue. Group I repairs performed on the liver, pancreas, lung and gallbladder specimens resulted in a very weak bond between the scaffold-enhanced solder and tissue, at only approximately 24-33% of the strength of native specimens. The strongest Group I vascular repairs were achieved in the carotid arteries, aorta (adventitia) and femoral arteries where breaking strengths of approximately 83%, 78% and 77% of their native tissue specimens, respectively, were achieved.
Although, the weakest vascular repairs were made on the pulmonary artery, the repairs still achieved greater than 62% of the strength of the native tissue. The overall percentage repair strength of native tissue was equivalent between Groups I and III (Group I Organs: 58±21%; Group III Organs: 55±22%; Group I Vessels: 72±8%; Group III Vessels: 72±12%). This does not imply, however, that the strength of Group I and Group III repairs were equivalent for each tissue type (see
Group II repairs utilizing the cyanoacrylate-scaffold composite all performed extremely well. Bonds formed using the Group II composites were on average 34% stronger than Group I and III organ repairs and 24% stronger than Group I and III vascular repairs.
Group III repairs performed utilizing a single 4-0 suture revealed the high variability in tensile strength associated with this repair technique. This method is highly dependent upon operator skill and technique as indicated by the large standard deviations seen within each tissue group; as well as, tissue type. Considering organ repairs (
Traditional strabismus surgery is time-consuming and technically demanding. Specialized spatulated needles must be passed mid-depth through a curved sclera that can be as little as 0.3 mm thick. Inadvertent ocular penetration during surgery can lead to blinding complications such as retinal detachment, vitreous hemorrhage and possibly endophthalmitis. A sutureless bioadhesive would eliminate many potential complications.
2.1 Surgical Procedure
Rabbit (n=12) superior rectus muscles (n=24) were isolated, severed from their scleral insertions and recessed to a point 4.0 mm from the corneoscleral limbus. Three experimental groups based on the method of repair were designated. The ‘Suture’ group utilized standard 6-0 polyglycolic acid sutures with spatulated needles to reattach muscles. The ‘Glue’ group utilized 2-octyl-cyanoacrylate applied directly to the sclera with the spread-out tendon (superior rectus muscle) held in the desired position (
2.2 Evaluation Techniques
Half of the animals were sacrificed at 2 days and the remainder were sacrificed at 14 days after surgery (
2.3 Results
The results of the tensile strength analysis are shown below in Table 3.
As shown in Table 3, preliminary experiments utilizing a glue+scaffold composite to reattach muscles following recession are encouraging. All attachments made using the composite maintained tensile strengths above that needed in humans following recession surgery. [Collins et al., Invest. Ophthal. Vis. Sci., 20:652-64, 1981] Additionally, the technique using the composite had improved ease of application which yielded more uniform results, as is reflected in the reduced variability compared to the other repair techniques evaluated.
Histologic examination of muscle insertions at 14 days showed no significant signs of inflammation in any of the groups. Muscle-sclera attachments were histologically similar to control insertions. Clinically, all animals tolerated the surgery well with minimal clinical signs of inflammation. The ‘Composite’ group provided a more accurate placement of the muscle compared to ‘Glue’ alone. It also provided more consistent tensile strength than either ‘Suture’ or ‘Glue’ alone.
3.1 Summary
Composites comprising biodegradable scaffolds doped with a cyanoacrylate adhesive were investigated for use in wound closure as an alternative to using cyanoacrylate adhesives alone. Two different scaffold materials were investigated: (i) a biological material, small intestinal submucosa (SIS), manufactured by Cook BioTech; and (ii) a synthetic biodegradable material fabricated from poly(L-lactic-co-glycolic acid) (PLGA). Ethicon's Dermabond™, a 2-octyl-cyanoacrylate, was used as the adhesive. The tensile strengths of skin incisions repaired in vivo in a rat model were measured at seven and fourteen days postoperatively, and the time to failure was recorded. Incisions closed by suture or by cyanoacrylate alone were also tested for comparison. Finally, a histological analysis was conducted to investigate variations in wound healing associated with each technique at seven and fourteen days postoperatively. Data relating to Example 3 is shown in Tables C, D, E, and F of the Appendix, and in
3.2 Materials and Methods
3.2.1 Preparation of PLGA Scaffolds
Porous synthetic polymer scaffolds were prepared from PLGA, with a lactic:glycolic acid ratio of 50:50, using a solvent-casting and particulate leaching technique. The scaffolds were cast by dissolving 200 mg PLGA (Sigma Chemical Company, St. Louis, Mo.) in 2 ml dichloromethane (Sigma Chemical Company). Sodium chloride (salt particle size: 106-150 μm) with a 70% weight fraction was added to the polymer mix. The polymer solution was then spread to cover the bottom surface of a 60 mm diameter Petri dish that was cleaned first with dichloromethane, then ethanol, then ultra-filtered deionized water (Fisher Scientific, Pittsburgh, Pa.). The polymer was left in a fume hood for 24 hours to allow the dichloromethane to evaporate. The salt was leached out of the polymer scaffolds by immersion in filtered deionized water for 24 hours, to create the porous scaffolds. During this period the water was changed 3-4 times. The scaffolds were then air dried and stored at room temperature until required. The PLGA scaffolds were cut into rectangular pieces with dimensions of 15±0.5 mm long by 10±0.5 mm wide. The average thickness of the scaffolds, determined by scanning electron microscopy and measurement with precision calipers, was 150±5 μm. Prior to use for tissue repair, the scaffolds were soaked in saline for a period of at least 10 minutes.
3.2.2 Preparation of SIS Scaffolds
SIS is prepared from decellularized porcine submucosa, which essentially contains intact extracellular matrix proteins, of which collagen is the most prevalent. Sheets of SIS, with surface dimensions of 50×10 cm and an average thickness of 100 μm, were provided by Cook BioTech (Lafayette, Ind.). The sheets were cut into rectangular pieces with dimensions of 15±0.5 mm long by 10±0.5 mm wide, and rehydrated in saline for at least 10 minutes prior to being using for tissue repair.
3.2.3 Surgical Repair
Eighteen Wistar rats, weighing 450±50 g, were anesthetized with a mixture of ketamine and xylazine. Four 15 mm long incisions were then made on the dorsal skin of each rat using a #15 scalpel blade: (1) left rostral parasagital; (2) right rostral parasagital; (3) left caudal parasagital; and (4) right caudal parasagital. Each incision site was randomly assigned to a one of the four repair techniques to be investigated.
The “Suture” group utilized three, equally spaced interrupted 5-0 polyglycolic acid (Vicryl) sutures. The “Cyanoacrylate alone” group was closed in accordance with the directions provided in the packaging by Ethicon, Inc. One-half an ampoule (˜0.175 mL) was used for each closure. For the “Cyanoacrylate+PLGA” group, five drops of Dermabond (˜0.035 mL) were applied to the irregular surface of the scaffolding using a 26G syringe to create the composite. The composite was then placed across the incision and allowed to air dry (˜10-20s). Finally, for the “Cyanoacrylate+SIS” group, the hydrated SIS specimens were observed to easily fold over on themselves, and were difficult to unravel afterwards. Thus, five drops of Dermabond (˜0.035 mL) were first applied to the incision site, and a piece of hydrated SIS scaffolding was then laid across the Dermabond with its irregular surface against the tissue.
Following the surgical procedure, all animals received a post-operative analgesic dose of buprenorphine. All animals were divided into two groups. Group I (n=13) were observed for seven days after surgery and Group II (n=5) were observed for fourteen days after surgery. At the end of the observation period, all animals were euthanized with pentobarbital and the surgical sites were excised for evaluation. Ten repairs for each wound closure technique from Group I and three repairs for each wound closure technique from Group II were prepared for tensile strength testing. The remaining incision sites that did not undergo strength testing were subjected to histological examination. A summary of incision treatments is given in Table 4:
3.2.4 Tensile Strength Analysis
The integrity of the resultant repairs were determined by tensile strength measurements performed immediately following the repair procedure using a calibrated MTS Material Strength Testing Machine (858 Table Top System, MTS, Eden Prairie, Minn.). This system was interfaced with a personal computer to collect the data. Each tissue specimen was clamped to the strength-testing machine using a 100N load cell with pneumatic grips. The specimens were pulled apart at a rate of 1 gf/sec until the repair failed. Complete separation of the two pieces of tissue defined failure. The maximum load in Newton's was recorded at the breaking point, as well as the time in seconds to failure. In order to avoid variations in repair strength associated with drying, the tissue specimens were kept moist during the procedure.
3.2.5 Histological Analysis
Light microscopy was used to assess the histological characteristics of wound healing associated with each technique at seven and fourteen days postoperatively. Harvested specimens were immediately fixed in formalin and stored at 6° C. until they could be prepared for staining and mounting. Hematoxylin and Eosin (H&E) was used as the staining agent.
3.3 Results
3.3.1 Wound Healing at Seven Days Postoperatively
The tensile strengths of the repair sites using the four different repair techniques harvested at seven days postoperatively are shown in
Typical photomicrographs of rat dorsal skin 7 days after standardized full-thickness incision and repair with: (i) 5-0 Nylon suture; (ii) standard external application of cyanoacrylate (Dermabond™); and (iii) external application of PLGA scaffold combined with cyanoacrylate, are shown in
3.3.2 Wound Healing at Fourteen Days Postoperatively
The tensile strengths of the repair sites using the four different repair techniques harvested at fourteen days postoperatively are shown in
3.4 Discussion
Differences in wound healing and tensile strength observed at 7 and 14 days post-operative can likely be explained by the properties of the different techniques.
SUTURES: Wound fixation by interrupted sutures creates a physical apposition of the dermis along the entire length of the wound. However, with any applied forces (including simply the movement and stretch of the skin as the animal moves and performs activities of daily living), the force is concentrated on the individual sutures. This allows differential movement of dermis between sutures and the contact away from the sutures is constantly being stressed, lost and reestablished with the alleviation of stress. In these areas, wound healing will be different and delayed from areas where dermis is kept in constant contact. Therefore, the wound healing between the sutures—which is the majority of the wound area—falls somewhere between true primary intention and secondary intention. Secondary intention healing always results in a longer time to restoration of wound integrity. Although it is sufficient, it is not optimal and at 7 and 14 days there are large areas of the wound that have not healed as well as they would if they were in constant physical apposition and were able to move in concert with externally applied stress.
CYANOACRYLATE: Cyanoacrylate alone performed comparably to that of suture repair. Early on it had less variability than that of sutures. This is likely due to the technical simplicity with which it is effectively applied versus that of the skill required and inherent variability in suture placement. Dermabond acts as a brittle scaffold that bridges the entire wound. This theoretically keeps the wound edges in apposition at all points along the closure. However, as our ex vivo and immediate tensile strength tests have shown, the tensile strength of cyanoacrylate alone is less than for the cyanoacrylate+scaffold composite. Cyanoacrylate is brittle and tends to lose adhesion either through cracking or a separation from the epithelium as an entire sheet when external stress is applied. In this study, early cracking and loss of tight continuous apposition along the entire length of the wound was noted within 24 hours with normal rat daily living activities. Since the animal will twist and bend and stretch the wound, cyanoacrylate is not an optimum method of skin wound closure. When the glue cracks and loses adhesion in focal areas, the healing replicates that of suture healing in that sections of the dermis are separated and must heal by something between true primary and secondary intention. With time, as adhesions are significantly lost, enough native tensile strength has returned to prevent significant numbers of dehiscences, but wound stretching and less cosmetic scar formation occurs along with a decrease in potential wound tensile strength early on.
COMPOSITE: The composite acts to keep the dermis in tight apposition throughout the critical early phase of wound healing when tissue gaps are bridged by scar and granulation tissue. It has the property of being more flexible than cyanoacrylate and may allow the apposed edges to move in conjunction with each other as a unit for a longer period of time and over a greater range of stresses than cyanoacrylate alone. This permits more rapid healing and establishment of integrity since the microgaps between the dermis edges are significantly reduced. By the time the scaffolds are sloughed (by either the animal scratching them off or loss of adhesion to the epithelium) there is greater strength and healing than that produced by cyanoacrylate alone and in wounds following suture removal.
4.1 Summary
Composites comprising biodegradable scaffolds doped with cyanoacrylate adhesive were investigated for use in wound closure as an alternative to using cyanoacrylate adhesives alone. Two different scaffold materials were investigated: (i) a biological material, small intestinal submucosa (SIS), manufactured by Cook BioTech; and (ii) a synthetic biodegradable material fabricated from poly(L-lactic-co-glycolic acid) (PLGA). Ethicon's Dermabond™, a 2-octyl-cyanoacrylate, was used as the adhesive. The tensile strengths of skin incisions repaired ex vivo in a rat model were measured, and the time to failure was recorded.
Data relating to Example 4 is shown in Tables G and H of the Appendix, and
4.2 Materials and Methods
4.2.1 Preparation of PLGA Scaffolds
Porous synthetic polymer scaffolds were prepared from PLGA, with a lactic:glycolic acid ratio of 50:50, using a solvent-casting and particulate leaching technique. The scaffolds were cast by dissolving 200 mg PLGA (Sigma Chemical Company, St. Louis, Mo.) in 2 ml dichloromethane (Sigma Chemical Company). Sodium chloride (salt particle size: 106-150 μm) with a 70% weight fraction was added to the polymer mix. The polymer solution was then spread to cover the bottom surface of a 60 mm diameter Petri dish that was cleaned first with dichloromethane, then ethanol, then ultra-filtered deionized water (Fisher Scientific, Pittsburgh, Pa.). The polymer was left in a fume hood for 24 hours to allow the dichloromethane to evaporate. The salt was leached out of the polymer scaffolds by immersion in filtered deionized water for 24 hours, to create the porous scaffolds. During this period the water was changed 3-4 times. The scaffolds were then air dried and stored at room temperature until required. The PLGA scaffolds were cut into square pieces with dimensions of 10±0.5 mm long by 10±0.5 mm wide. The average thickness of the scaffolds, determined by scanning electron microscopy and measurement with precision calipers, was 150±5 μm. Prior to use for tissue repair, the scaffolds were soaked in saline for a period of at least 10 minutes.
4.2.2 Preparation of SIS Scaffolds
SIS is prepared from decellularized porcine submucosa, which essentially contains intact extracellular matrix proteins, of which collagen is the most prevalent. Sheets of SIS, with surface dimensions of 50×10 cm and an average thickness of 100 μm, were provided by Cook BioTech (Lafayette, Ind.). The sheets were cut into square pieces with dimensions of 10±0.5 mm long by 10±0.5 mm wide, and rehydrated in saline for at least 10 minutes prior to being using for tissue repair.
4.2.3 Tissue Preparation and Incision Repair
The dorsal skin from thirteen Wistar rats was excised immediately after sacrificing the animals. Rectangular tissue specimens were cut from the skin samples with dimensions of about 20 mm long by 10 mm wide.
A full thickness incision was made with a scalpel across the width of the tissue specimen. Four drops of Dermabond™ were then applied to the irregular surface of the scaffolding using a 27G syringe and the adhesive material was placed across the incision and allowed to air dry. A sample size of ten was used for all experimental groups.
4.2.4 Tensile Strength Analysis
The integrity of the resultant repairs were determined by tensile strength measurements performed immediately following the repair procedure using a calibrated MTS Material Strength Testing Machine (858 Table Top System, MTS, Eden Prairie, Minn.). This system was interfaced with a personal computer to collect the data. Each tissue specimen was clamped to the strength-testing machine using a 100N load cell with pneumatic grips. The specimens were pulled apart at a rate of 1 gf/sec until the repair failed. Complete separation of the two pieces of tissue defined failure. The maximum load in Newton's was recorded at the breaking point, as well as the time profiles for failure of the repairs. In order to avoid variations in repair strength associated with drying, the tissue specimens were kept moist during the procedure. The strengths of corresponding specimens repaired with cyanoacrylate alone, in accordance with the directions provided by Ethicon, Inc., were tested and used as references.
4.3 Results
The tensile strength of the repairs performed in this acute wound closure study using cyanoacrylate alone and a composite including cyanoacrylate enhanced by a scaffold fabricated from either SIS or PLGA, are shown in
4.4 Discussion
Successful wound closure will occur when dermal edges are kept in physical contact (or with as little gap as possible) so that granulation and scar tissue can result in a continuous integrated matrix from edge to edge. This principle of unobstructed apposition also applies to any non-dermal tissues/surfaces where physical attachment (or reattachment) to another dermal or non-dermal surface is desired. When cyanoacrylate is applied externally to a wound and not allowed to penetrate the reticular dermal level or deeper, it provides a consistent low strength bonding of epidermal surfaces. This keeps the dermal edges in apposition so that wound healing can progress unobstructed. Failure of cyanoacrylate surface closure occurs when either the epithelium (which is loosely attached to the papillary dermis) sloughs off, or the glue loses adhesion to the epithelium for various reasons. These reasons include oil secretion and sloughing of dead surface cells.
The composite formed of either a biocompatible (i.e. PLGA) or biological (i.e. SIS) scaffold and an adhesive provided significantly enhanced tensile strength of the adhesion. This produced a consistently stronger adhesion under standardized constantly increasing tensile strength testing conditions.
The combination of either a biocompatible (i.e. PLGA) or biological (i.e. SIS) scaffold and adhesive also produced different physical characteristics of the adhesion—in a favorable manner. Under constantly increasing tensile stress, force generation curves were prolonged in reaching their peaks. This indicates that adhesions resulting from application of the composite could distribute the forces better and withstand stress for longer periods of time.
The composite including either a biocompatible (i.e. PLGA) or biological (i.e. SIS) scaffold and adhesive also produced different peak-trough behavior of the length-tension curves than the adhesive alone. With the composite, adhesions frequently displayed many mini peaks, without significant troughs, with quick recovery of functional tensile strength. Cyanoacrylate alone almost always produced a single (or infrequently a doublet) peak followed by complete failure of strength and complete physical separation of tissues.
Thus, the composite provides a stronger, more durable and consistent adhesion than the adhesive alone. This theory is also supported by several ex vivo experiments demonstrating enhanced tensile strength of irregular porous versus smooth surface scaffolds in identical tissue repairs (refer to Example 5).
5.1 Summary
An ex vivo study was conducted to determine the effect of the irregularity of the scaffold surface on the tensile strength of repairs formed using a composite comprising a scaffold and a biological adhesive. Two different scaffold materials were investigated: (i) a synthetic biodegradable material fabricated from poly(L-lactic-co-glycolic acid) (PLGA); and (ii) a biological material, small intestinal submucosa (SIS), manufactured by Cook BioTech. Ethicon's Dermabond™, a 2-octyl-cyanoacrylate, was used as the adhesive. The tensile strength of repairs performed on bovine thoracic aorta, liver, spleen, small intestine and lung, using both the smooth and irregular surfaces of the above materials were measured and the time to failure was recorded.
Data relating to Example 5 is shown in Tables I-1, I-2, I-3, I-4, and I-5 of the Appendix, and
5.2 Materials and Methods
5.2.1 Preparation of PLGA Scaffolds
Porous synthetic polymer scaffolds were prepared from PLGA, with a lactic:glycolic acid ratio of 50:50, using a solvent-casting and particulate leaching technique. The scaffolds were cast by dissolving 200 mg PLGA (Sigma Chemical Company, St. Louis, Mo.) in 2 ml dichloromethane (Sigma Chemical Company). Sodium chloride (salt particle size: 106-150 nm) with a 70% weight fraction was added to the polymer mix. The polymer solution was then spread to cover the bottom surface of a 60 mm diameter Petri dish that was cleaned first with dichloromethane, then ethanol, then ultra-filtered deionized water (Fisher Scientific, Pittsburgh, Pa.). The polymer was left in a fume hood for 24 hours to allow the dichloromethane to evaporate. The salt was leached out of the polymer scaffolds by immersion in filtered deionized water for 24 hours, to create the porous scaffolds. During this period the water was changed 3-4 times. The scaffolds were then air dried and stored at room temperature until required. The PLGA scaffolds were cut into square pieces with dimensions of 10±0.5 mm long by 10±0.5 mm wide. The average thickness of the scaffolds, determined by scanning electron microscopy and measurement with precision calipers, was 150±5 mm. Prior to use for tissue repair, the scaffolds were soaked in saline for a period of at least 10 minutes.
5.2.2 Preparation of SIS Scaffolds
SIS is prepared from decellularized porcine submucosa, which essentially contains intact extracellular matrix proteins, of which collagen is the most prevalent. Sheets of SIS, with surface dimensions of 50×10 cm and an average thickness of 100 μm, were provided by Cook BioTech (Lafayette, Ind.). The sheets were cut into square pieces with dimensions of 10±0.5 mm long by 10±0.5 mm wide, and rehydrated in saline for at least 10 minutes prior to being using for tissue repair.
5.2.3 Surface Analysis using Scanning Electron Microscopy
Prior to conducting any tissue repairs, sample surfaces of all scaffolds to be investigated were viewed with a Hitachi S-3000N scanning electron microscope (SEM) to characterize the degree and nature of their smoothness or irregularity.
5.2.4 Tissue Preparation and Incision Repair
Bovine tissue specimens were harvested approximately 30 minutes after sacrificing the animals. Tissue specimens were stored in phosphate buffered saline for a maximum of two hours before they were prepared for experiments. Each tissue specimen was cut into small rectangular pieces with dimensions of about 20 mm long by 10 mm wide and a thickness of approximately 1.5±0.5 mm. Tissue specimens harvested included the thoracic aorta, liver, spleen, small intestine, and lung.
A full thickness incision was made with a scalpel across the width of the tissue specimen. Four drops of Dermabond™ were then applied to the desired surface of the scaffolding (smooth or irregular) using a 26G syringe and the adhesive material was placed across the incision and allowed to air dry. A sample size of ten was used for all experimental groups.
5.2.5 Tensile Strength Analysis
The integrity of the resultant repairs were determined by tensile strength measurements performed immediately following the repair procedure using a calibrated MTS Material Strength Testing Machine (858 Table Top System, MTS, Eden Prairie, Minn.). This system was interfaced with a personal computer to collect the data. Each tissue specimen was clamped to the strength-testing machine using a 100N load cell with pneumatic grips. The specimens were pulled apart at a rate of 1 gf/sec until the repair failed. Complete separation of the two pieces of tissue defined failure. The maximum load in Newton's was recorded at the breaking point, as well as the time in seconds to failure. In order to avoid variations in repair strength associated with drying, the tissue specimens were kept moist during the procedure. The strengths of corresponding native specimens and incisions repaired with cyanoacrylate alone were tested and used as references.
5.3 Results
Electron micrographs of both the smooth (intimal) and irregular surfaces of the SIS scaffolds are shown in
The tensile strength of repairs performed on bovine thoracic aorta, liver, spleen, small intestine and lung, by applying either the smooth or the irregular surfaces of the composites to the tissue surface, are shown in
5.4 Discussion
Several key points are immediately noted from
The irregular, rough surface of the composite provides a greater tensile strength immediately after the adhesion is initiated than does the cyanoacrylate alone, approximating the native tissue strength.
The smooth surface of the composite provides a small increase in tensile strength over cyanoacrylate alone; however, the rough surface of the composite provides a consistently high tensile strength, approximating the native tensile strength of all tissues tested. These results suggest that distributing or dispersing the adhesive forces over an increased surface area of the scaffold, either smooth or rough, can produce better results than cyanoacrylate alone. However, an irregular, rough, or porous surface can significantly increase tensile strength. This presumably occurs by distributing the forces between thousands or millions of independent “microadhesion”.
The clinical relevance of these results is significant. Surgical repairs are more likely to fail in the first hours-to-days after surgery as a result of several factors: a) wound edges are only apposed by whatever artificial means was employed to repair the incision; these methods are subject to the limitations of how they grasp the tissues and anchor them together; b) during the early surgical period, there has not been significant time enough for primary or secondary intention wound healing to provide any native tensile strength to the apposition itself; c) postoperatively edema (which contributes increased forces on the wound, greater than that seen at the time of repair) is greatest in the first 24 hours after surgery (often increasing over this period of time); and d) certain tissues will immediately be subject to high forces after repair/surgery, i.e. aortic pulsatile blood pressure, muscle/tendon contractions against insertions, etc.
All the above factors may contribute to the early postoperatively failure of suture or other methods of repair, such as adhesives or staples. If a tissue repair can achieve a tensile strength approximating the native tensile strength of the tissue in the immediate postoperatively period, the likelihood of failure is markedly diminished and it is certainly much less likely to fail than would a system characterized by more variability and lower tensile strengths.
6.1 Summary
An ex vivo study was conducted to determine the effect of varying the area of the scaffold surface in contact with the tissue on the tensile strength of repairs formed using a scaffold-enhanced biological adhesive composite. Biodegradable polymer scaffolds of controlled porosity were fabricated with poly(L-lactic-co-glycolic acid) and salt particles using a solvent-casting and particulate-leaching technique. The scaffolds were doped with Ethicon's Dermabond™, a 2-octyl-cyanoacrylate adhesive. The tensile strength of repairs performed on bovine thoracic aorta and small intestine were measured and the time to failure was recorded.
Data relating to Example 6 is shown in Tables J-1 and J-2 of the Appendix, and in
6.2 Materials and Methods
6.2.1 Preparation of PLGA Scaffolds
Porous synthetic polymer scaffolds were prepared from PLGA, with a lactic:glycolic acid ratio of 50:50, using a solvent-casting and particulate leaching technique. The scaffolds were cast by dissolving 200 mg PLGA (Sigma Chemical Company, St. Louis, Mo.) in 2 ml dichloromethane (Sigma Chemical Company). Sodium chloride (salt particle size: 106-150 μm) with a 70% weight fraction was added to the polymer mix. The polymer solution was then spread to cover the bottom surface of a 60 mm diameter Petri dish that was cleaned first with dichloromethane, then ethanol, then ultra-filtered deionized water (Fisher Scientific, Pittsburgh, Pa.). The polymer was left in a fume hood for 24 hours to allow the dichloromethane to evaporate. The salt was leached out of the polymer scaffolds by immersion in filtered deionized water for 24 hours, to create the porous scaffolds. During this period the water was changed 3-4 times. The scaffolds were then air dried and stored at room temperature until required. The PLGA scaffolds were cut into rectangular pieces with the desired surface dimensions (length by width): (i) 10±0.5 mm by 10±0.5 mm; (ii) 10±0.5 mm by 5±0.5 mm; (iii) 5±0.5 mm by 10±0.5 mm; (iv) 15±0.5 mm by 10±0.5 mm; and (v) 15±0.5 mm by 5±0.5 mm. The average thickness of the scaffolds, determined by scanning electron microscopy and measurement with precision calipers, was 150±5 μm. Prior to use for tissue repair, the scaffolds were soaked in saline for a period of at least 10 minutes.
6.2.2 Tissue Preparation and Incision Repair
Bovine tissue specimens were harvested approximately 30 minutes after sacrificing the animal. Tissue specimens were stored in phosphate buffered saline for a maximum of two hours before they were prepared for experiments. Each tissue specimen was cut into small rectangular pieces with dimensions of about 20 mm long by 10 mm wide and a thickness of approximately 1.5±0.5 mm. Tissue specimens harvested included the thoracic aorta and small intestine.
A full thickness incision was made with a scalpel across the width of the tissue specimen. Four drops of Dermabond™ were then applied to the irregular surface of the scaffold using a 26G syringe, and the composite was placed across the incision and allowed to air dry. A sample size of ten was used for all experimental groups.
6.2.3 Tensile Strength Analysis
The integrity of the resultant repairs was determined by tensile strength measurements performed immediately following the repair procedure using a calibrated MTS Material Strength Testing Machine (858 Table Top System, MTS, Eden Prairie, Minn.). This system was interfaced with a personal computer to collect the data. Each tissue specimen was clamped to the strength-testing machine using a 100N load cell with pneumatic grips. The specimens were pulled apart at a rate of 1 gf/sec until the repair failed. Complete separation of the two pieces of tissue defined failure. The maximum load in newtons was recorded at the breaking point, as well as the time in seconds to failure. In order to avoid variations in repair strength associated with drying, the tissue specimens were kept moist during the procedure.
6.3 Results
The tensile strength of repairs performed on bovine thoracic aorta and small intestine by applying the irregular surface of the cyanoacrylate-PLGA scaffold composites to the tissue surface, are shown in
6.4 Discussion
As shown in
With the composite, a butterfly-bandage effect occurs, i.e., reinforcement of the wound by the combination of the scaffold and glue brought the edges of the incision, along its entire length, into better apposition for an extended period of time, which contributed to a more satisfactory cosmetic healing.
Geometry may not be completely unimportant (as one would expect when dealing with vector forces). However, it may be clinically insignificant. As seen in small intestine repair, less surface area (oriented differently) had a statistically significant effect (p<0.05): 10×10 mm versus 15×5 mm. This is, however, the only result like this and, depending on the size and orientation of the actual tissue in the experiment, it may be a clinically insignificant isolated result. While the rest of the time points reveal that surface area is likely proportional to the increased time to failure, as would be expected, further studies are needed to confirm these results.
7.1. Summary
An ex vivo study was conducted to determine the effect of using several different custom modified scaffold surfaces on the tensile strength of repairs formed using our scaffold-adhesive composite. Porous PLGA scaffolds were fabricated using four different manufacturing techniques: (i) a computer-controlled drilling technique; (ii) a punching technique utilizing an arbor press; (iii) a polymer molding technique, and (iv) 220 grit sandpaper.
Data relating to this Example 7 is shown in Tables K-1, K-2, K-3, K-4 and K-5 of the Appendix, and in
7.2 Materials and Methods
7.2.1 Preparation of PLGA Using Various Mechanical Manufacturing Techniques
Synthetic polymer scaffolds were prepared from PLGA, with a lactic:glycolic acid ratio of 50:50. The scaffolds were cast by dissolving 250 mg PLGA in 2.5 ml dichloromethane. The polymer solution was then spread to cover the bottom surface of a 60 mm diameter Petri dish that was cleaned first with dichloromethane, then ethanol, then ultra-filtered deionized water. The polymer was left in a fume hood for 24 hours to allow the dichloromethane to evaporate, and then allowed to soak in filtered deionized water for a period of 2 hours prior to removing from the Petri dish.
Upon drying of the polymer scaffolds, an irregularity was added to the scaffold surfaces using one of four mechanical techniques:
The PLGA scaffolds were cut into square pieces with dimensions of 10±0.5 mm long by 10±0.5 mm wide. The average thickness of the scaffolds, determined by scanning electron microscopy and measurement with precision calipers, was 150±10 μm. Prior to use for tissue repair, the scaffolds were soaked in saline for a period of at least 10 minutes.
7.2.2 Surface Analysis using Scanning Electron Microscopy
Prior to conducting any tissue repairs, the surfaces of samples of all scaffolds to be investigated were viewed with a Hitachi S-3000N scanning electron microscope (SEM) to allow characterization of their irregularity.
7.2.3 Tissue Preparation and Incision Repair
Bovine tissue specimens were harvested approximately 30 minutes after sacrificing the animal. Tissue specimens were stored in phosphate buffered saline for a maximum of two hours before they were prepared for experiments. Each tissue specimen was cut into small rectangular pieces with dimensions of about 20 mm long by 10 mm wide and a thickness of approximately 1.5±0.5 mm. Tissue specimens harvested included the thoracic aorta, liver, spleen, small intestine, and lung.
A full thickness incision was made with a scalpel across the width of the tissue specimen. Four drops of Dermabond™ were then applied to the rough surface of the scaffolding using a 26G syringe, and the adhesive material was placed across the incision and allowed to air dry. A sample size of five was used for all experimental groups.
7.2.4 Tensile Strength Analysis
The integrity of the resultant repairs was determined by tensile strength measurements performed immediately following the repair procedure using a calibrated MTS Material Strength Testing Machine (858 Table Top System, MTS, Eden Prairie, Minn.). This system was interfaced with a personal computer to collect the data. Each tissue specimen was clamped to the strength-testing machine using a 100N load cell with pneumatic grips. The specimens were pulled apart at a rate of 1 gf/sec until the repair failed. Complete separation of the two pieces of tissue defined failure. The maximum load in Newton's was recorded at the breaking point, as well as the time in seconds to failure. In order to avoid variations in repair strength associated with drying, the tissue specimens were kept moist during the procedure.
7.3 Results
Electron micrographs of the PLGA polymer scaffolds given an irregular surface using one of the four mechanical techniques described above are shown in
The tensile strength of repairs performed on bovine thoracic aorta, liver, spleen, small intestine and lung, using the cyanoacrylate-scaffold composites described above, are shown in
7.4 Discussion
As can be seen in the photomicrographs, irregular scaffold surfaces can be manufactured to different specifications of irregularity and porosity, in order to suit various surgical requirements. The photomicrographs of the PLGA scaffolds produced using the punch and sandpaper techniques show the greatest areas of troughs, where the tissue would be in direct contact with the adhesive rather than the scaffold material. Repairs formed using scaffolds manufactured using the punch and sandpaper techniques were the strongest of the four custom manufactured scaffolds investigated (
Clinical relevance is less apparent here, other than as support to our theory described in Example 3. However, this finding suggests that many aspects of these scaffolds may be custom manufactured, including porosity (including pore size and distribution), roughness, non-geometric topography (irregularity), to ensure reproducibility of results and to meet the needs of specific applications.
Future studies may be directed at determining whether different surfaces actually work better with one type of adhesive versus another or with adhesives of different viscosity allowing deeper penetration into the depth of the surface irregularities.
As a result of these and other studies, it has been found that a non-light activated adhesive-scaffold composite, incorporating a biological, biocompatible, or biodegradable adhesive and a biological, biocompatible, or biodegradable scaffold, exhibits significantly enhanced tensile strength and consistently stronger adhesion under constantly increasing time periods of tensile strength testing. Also, the composite exhibits more favorable adhesion characteristics. When subjected to constantly increasing loads, the composites exhibited force generation curves that were prolonged in reaching their peaks, indicating better distribution of forces. This allowed the composites to withstand stress for longer periods of time.
Additionally, length-tension curves for the composites are remarkably different than those for bioadhesives alone (e.g., cyanoacrylate). While the bioadhesive alone frequently produced a single peak followed by a trough (indicating complete failure of strength and complete physical separation of tissues), the composite curve showed many peaks without significant troughs (indicating quick recovery of functional tensile strength and little-to-no tissue separation) (
The specifications of the composite of the present invention can be tailored to meet the specific requirements of a range of clinical applications, such as wound closure from trauma or at surgical incision sites, repair of liver, spleen, or pancreas lacerations from trauma, dural laceration/incision closure, pneumothorax repair during thoracotomy, sealing points of vascular access following endovascular procedures, vascular anastomoses, tympanoplasty, endoscopic treatment of gastrointestinal ulcers/bleeds, dental applications for mucosal ulcerations or splinting of injured teeth, ophthalmologic surgeries, tendon and ligament repair in orthopedics, and episiotomy/vaginal tear repair in gynecology. Patches prepared using the adhesive composites can be used in a non-surgical setting as a simple, quick, and effective wound closure solution, for example, in emergency situations.
The composite of the present invention may be created by a variety of methods or techniques. For example, a physician or other health care provider may place the scaffold in the desired position for tissue repair, sealing, or adhesion, then apply the adhesive to the scaffold. Alternatively, the adhesive may be applied to the scaffold and then the device containing both scaffold and adhesive placed in position. As another alternative, the adhesive may be placed at the repair site first and then the scaffold applied. Additional adhesive material may be applied to the site before or after the scaffold is positioned. It is understood that the terms “placed” and “positioned” include applying an adhesive and/or scaffold on a wound, tissue, or repair site, across edges of a wound or incision, and/or across a juncture between tissue and a biocompatible implant to be joined or adhered.
The composite of the present invention may be designed and packaged in a variety of different ways. For example, in one embodiment, the composite is packaged in an inert cellophane-like material. The inert material peels off the surface of the composite to allow immediate use. The packaged item may be made available in a variety of sizes and shapes as appropriate for various uses or applications.
In another embodiment, the composite is supported by one or two rollers made of an inert material. The rollers may be configured to be disposable or reusable. The composite is wrapped around the roller or rollers to form a scroll. The scroll is unrolled to apply the composite to a wound or repair site; for example, a curved or irregular surface. A double roller scroll is particularly advantageous in a non-sterile setting (such as an emergency setting, where surgical/sterile gloves are not available), since it avoids the need for a person to directly handle the composite. A single roller scroll is particularly suitable for sterile environments, for example, during surgery, where a gloved hand may be used to position the edge of the composite prior to unrolling.
Yet another alternative packaging technique involves positioning a thin, expendable, fracturable membrane on top of the composite in such a way that the thin membrane protects the composite until it is ready to be used. Upon application of the composite to a wound or repair site, the expendable membrane ruptures or fractures, for example, to expose the adhesive to the desired tissue site.
Further alternative embodiments involve the use of a separator, such as an inert tab made of plastic, paper, or other suitable material, to which a grip, for example a ring (similar to that used in laser printer cartridges), is attached. In one such alternative embodiment, a separator is positioned between the scaffold and the adhesive to isolate the scaffold from the adhesive until the composite is needed for application to a wound or repair site (
In another such alternative embodiment, the separator is positioned between the adhesive and an adhesive activator to isolate the adhesive from its activator until the composite is needed for use (
In yet another such alternative embodiment, two separators may be provided. A first separator may be positioned between the scaffold and the adhesive, and a second separator positioned between the adhesive and the activator. In this embodiment, one grip may be provided to remove the separator between the activator and adhesive in order to activate the adhesive, and then a second grip may be provided to remove the separator between the adhesive and scaffold, to enable contact between the adhesive and the scaffold. This design may be useful in situations where it may be necessary or desirable to activate the adhesive a certain amount of time prior to application of the composite to the wound or repair site. Alternatively, one grip may be provided, which operates to remove both separators at once.
The composite can be modified to provide biologically active materials to biological tissue. The controlled release of various dopants including hemostatic and thrombogenic agents, antibiotics, anesthetics, various growth factors, enzymes, anti-inflammatories, bacteriostatic or bacteriocidal factors, chemotherapeutic agents, anti-angiogenic agents and vitamins can be added to the composite to assist in the therapeutic goal of the procedure. The degradation rate of the composite, and consequently the drug delivery rate, can be controlled by altering the macromolecular structure of the device or a portion thereof.
Furthermore, the elasticity, strength, and flexibility of the composite can be modified to meet the demands of and enhance clinical applicability in a wide range of applications. For example, alteration of composition and pore size modifies pliability and elasticity, making it easier to process and fabricate the composite, for example, into different forms and shapes for different applications.
Although specific illustrated embodiments of the invention have been disclosed, it is understood by those skilled in the art that changes in form and details may be made without departing from the spirit and scope of the invention. The present invention is not limited to the specific details disclosed herein, but is to be defined by the appended claims.