Adaptable suture with dynamically changing wound retention properties after implantation

Abstract

The present invention relates to a length-adaptable surgical suture comprising a monofilament or a braid of multiple filaments, the suture having an initial length at implantation and a second length within the first twenty-four (24) hours after implantation that is different from the initial length to accommodate tissue swelling. The present invention also relates to constructions and combinations that achieve the length-adaptability result.

Description

Technical Field

[0001] The technical field of this invention is sutures capable of adapting to changing wound and tissue environments, including those that elongate and / or contract in the human body after implantation to respond to and adapt to increases and decreases in tissue swelling or to tighten certain body tissues or to elongate and / or contract as healing progresses. Background Technology

[0002] Surgical sutures and attached surgical needles used in a variety of routine surgical procedures are known in the art. For example, such sutures can be used to wrap around incisions or lacerations in the epidermis and underlying fascia close to tissue, to join the ends of blood vessels, to attach tissue to medical devices (such as heart valves), to repair body organs, to repair connective tissue, etc. Conventional surgical sutures can be prepared from known biocompatible materials, specifically synthetic and natural biocompatible polymeric materials, which can be absorbable or non-absorbable. Examples of synthetic non-absorbable polymeric materials that can be used to manufacture non-absorbable sutures include polyesters, polyolefins, polyvinylidene fluoride and polyamides, polypropylene, nylon, etc. Examples of synthetic absorbable polymeric materials that can be used to manufacture absorbable sutures include polymers and copolymers made from lactones such as lactide, glycolide, p-dioxane, ε-caprolactone, and trimethylene carbonate. The term "absorbable" is a general term and may also include bioabsorbability, reabsorbability, bioreabsorbability, degradability, or biodegradability.

[0003] Both absorbable and non-absorbable sutures must provide the necessary tensile strength within the body for a sufficient period to allow for effective tissue healing. Wound healing depends on the nature of the specific tissue and the healing characteristics of the individual undergoing surgery. For example, there are cases where tissues and organs swell rapidly after surgery or some mechanical injury. Closing the wound edges after the invasive procedure holds them well, but after a period of time, the swelling begins to decrease. This process may loosen conventional sutures because they cannot accommodate this new tissue contraction. Furthermore, tissue healing can lead to further tissue contraction. This, in turn, can result in undesirable leakage and wound opening.

[0004] Given the swelling of tissue after suturing, there is a strong need for adaptive / dynamic tissue-specific sutures that will close the wound in a regular manner but adapt to tissue swelling, including expansion to accommodate the swelling and / or contraction after the swelling subsides. There are other areas of tissue repair where the contraction of repair sutures over time can be used to accommodate tissue healing and the reduction of swelling. One such area is tendon fixation, where a 2%–3% to 20%–30% longitudinal contraction occurring within days / weeks may be of concern.

[0005] US Patent 4,942,089, "Rapidly Shrinking Fiber And Water-Absorbing Shrinkable Yarn And Other Materials Comprising Same," discloses a polyvinyl alcohol fiber that can rapidly shrink in the presence of water but is almost insoluble in water. The fiber is characterized by a maximum shrinkage rate of not less than 30% in water at 20°C and a time required to reach 30% of the shrinkage rate of not more than 10 seconds; a shrinkage stress of not less than 150 mg / dr measured in water at 20°C at the initial length and a time required for the shrinkage stress of 150 mg / dr to appear of not more than 10 seconds; a shrinkage stress of not less than 30 mg / dr measured in water at 20°C at the state of shrinkage of 30% relative to the initial length; and a weight loss due to dissolution when dispersed in water at 20°C of not more than 45%.

[0006] DYNACORD manufactured by DePuy Synthes TM The sutures are designed to expand radially and contract axially in the presence of loosening, allowing the sutures to minimize negative aspects of the repair, such as suture loosening, creep, and knot slippage. (DYNACORD) TM The sutures are high-strength orthopedic sutures designed to minimize repair laxity, maintain consistent tissue proximity, and improve footprint compression. (DYNACORD) TM The sutures are designed to shorten when compression is lost, thereby minimizing micromovement and gap formation.

[0007] U.S. Patent No. 8,870,915, “Joining Element,” discloses a longitudinally elongating joining element comprising: a first material including a plurality of wires, the first material elongating along the longitudinal direction; and a second material extending along the longitudinal direction through the first material, the second material having a volume capable of swelling in a direction transverse to the longitudinal direction, which in turn causes longitudinal contraction relative to the first material.

[0008] An article entitled “Degradation Behaviors of Electrospun Resorbable Polyester Nanofibers”, Yixiang Dong et al., Tissue Engineering Part B, September 2009, 15(3):333-51, discloses that amorphous PLGA or PDLLA electrospun nanofibers shrink during degradation due to thermally induced relaxation of stretched amorphous chains.

[0009] An article entitled “Analysis of Shrinkage Development of a Semicrystalline Polymer during Injection Molding” by Felice De Santis et al., Ind. Eng. Chem. Res. 2010, Vol. 49, pp. 2469-2476, discloses a study on the effects of shrinkage of a semicrystalline polymer during injection molding, as well as holding time and pressure, on the shrinkage evolution from the moment of first curing inside the die to shortly after molding.

[0010] US Patent No. 8088146, "High-Strength" "Suture" discloses a surgical suture comprising an elongated braided fiber fabric corresponding to a size ranging from USP size 5-0 to USP size 7 sutures. The fabric has 8 to 32 fibers woven together at approximately 50 wefts / inch to approximately 70 wefts / inch. More than 90% of the fibers are ultra-high molecular weight polyethylene fibers. The fiber fabric is hollow and defines an elongated, longitudinally extending, open central chamber in which no core material extends. The chamber has a cross-sectional diameter substantially larger than the fiber diameter, and the chamber allows the cross-sectional shape of the fabric to change in response to pressure experienced when the suture is knotted, allowing for the formation of an anti-slip knot. The diameter of the suture is in the range of 0.100 mm to 0.999 mm, and the suture has a substantially circular cross-sectional configuration that collapses in response to pressure experienced when the suture is knotted. The total denier of the suture is between approximately 880 denier and 3520 denier.

[0011] U.S. Patent No. 4,880,002, “Stretchable Porous Sutures,” discloses a nonwoven surgical suture comprising: a suture material comprising: an elongated, flexible, nonmetallic, generally cylindrical member having multiple pores; the elongated, flexible, porous member being an elastomer and a generally cylindrical and flexible extrusion of a mixture of a polymeric material and a washable material, the extrusion having been eluted by the washable material from the polymeric material to form the pores of the elongated, flexible, elastomeric member; the polymeric material being selected from polyurethane and polycarbonate; the elongated, flexible, elastomeric member having an initial extrusion length and being axially stretchable to approximately two times or more of its initial extrusion length, thereby allowing the suture to conform to host tissue upon implantation; and the outer diameter of the elongated, flexible, elastomeric member not greater than the outer diameter of the surgical suture, and the elongated, flexible, elastomeric member having a porous surface that promotes tissue inward growth into the elongated, flexible, porous member.

[0012] U.S. Patent Application Publication No. 2006 / 0121274, “Sheathed Elastic Surgical Thread,” discloses a surgical suture having an elastic core and a non-elastic sheath. The elastic core consists of one or more biocompatible elastic threads, and the sheath consists of one or more non-stretched threads made of a biocompatible material. The non-stretched sheath threads and the elastic core are made of a non-absorbent material, and the non-stretched sheath threads are wound in a helical manner to form one or more covering helices wound in opposite directions, which restricts the maximum extension of the elastic threads.

[0013] European Patent Publication EP1501559 B1, "Surgical Thread and Surgical Implant With The Same," discloses a surgical thread comprising a core made of at least one reabsorbable material and a cover made of at least one non-reabsorbable material and / or a slowly reabsorbable material that reabsorbs more slowly than the reabsorbable material of the core. The cover comprises threads, and is characterized in that the threads of the cover are arranged in the surgical thread as a single-wrapped twist or a spun-wrapped twist, and wherein, prior to reabsorption of the core, the threads of the cover are dimensionally stable against tension of the core, and wherein, after reabsorption of the core, they lose their dimensional stability, such that, when subjected to tension, the cover can shift from its non-linear arrangement to a presumably linear arrangement. This reference discloses very large elongation achieved over a time much longer than the tissue swelling time range, which would therefore be excessive for accommodating tissue swelling.

[0014] U.S. Patent No. 6,966,918, “Reinforced Areal Implant,” discloses a reinforced areal implant comprising a mesh-like base structure with apertures ranging from 1.5 mm to 4.0 mm and a fabric reinforcement element. The fabric reinforcement element, with a support length of 20 mm, exhibits bending strength in a three-point flexibility test ranging from 0.015 N / mm to 0.4 N / mm. The fabric reinforcement element forms a mesh-like reinforcement structure with apertures ranging from 5 mm to 30 mm, the apertures being multiples of the apertures of the base structure.

[0015] U.S. Patent 6,162,962, “Areal Implant,” discloses a facial implant comprising: a flexible knitted fabric having an initial tear strength, optionally having a reabsorption time of at least 60 days, wherein the flexible knitted fabric is designed to stretch below a critical force beyond the tissue area intended to receive the implant and above the critical force less than the tissue area, the critical force being below the maximum allowable load in the tissue area; and a synthetic reabsorbable material that hardens the flexible knitted fabric, the reabsorption time being less than the reabsorption time of the flexible knitted fabric, wherein the synthetic reabsorbable material is selected from yarns, monofilaments, and combinations thereof.

[0016] Dynamic or adaptive sutures are required, which are designed to adapt to tissue swelling and / or tissue contraction immediately after surgery. These sutures are constructed to expand (elongate) and contract (contract longitudinally) after surgery to accommodate tissue response and healing. Summary of the Invention

[0017] In short, adaptive or dynamic surgical sutures are configured to dynamically alter wound retention properties (suture length) after implantation. In one embodiment, after implantation and exposure to tissue and fluids, the suture of the present invention reduces its length, or shortens, or longitudinally contracts, thereby tightening the wound or retaining tissue as tissue healing and swelling subside. In another embodiment, after implantation and exposure to tissue and fluids, the suture of the present invention initially elongates or extends longitudinally to accommodate tissue swelling, and thereafter permanently reduces its length or longitudinally contracts, thereby tightening the wound or retaining tissue as tissue healing and swelling subside.

[0018] In some embodiments, the sutures of the present invention may be monofilament or braided sutures, absorbable or non-absorbable sutures, and are used to close wounds or connect tissues during surgical repair in a manner known in the surgical field. The sutures then optionally extend or elongate longitudinally upon exposure to bodily fluids and tissues to facilitate the release of tension during postoperative tissue swelling. In some embodiments, the sutures of the present invention optionally elongate by 2%, 3%, 5%, 7%, 10%, 15%, 20%, and most preferably about 10% of their initial length within 2 hours, 4 hours, 6 hours, 12 hours, 18 hours, 24 hours, 30 hours, 36 hours, or 48 hours after suture placement, preferably within about 18 to 36 hours, and most preferably within about 24 hours, from their initial length.

[0019] Following the elongation step (which may be optional), the sutures of the present invention are shortened or contracted longitudinally to accommodate the reduction of tissue swelling and the tissue healing process. In some embodiments, the sutures of the present invention contract to about 75% to 100% of their initial length within about 36 to 96 hours post-surgery, more preferably about 80% to 98%, and most preferably about 90% to 95% of their initial length within about 48 to 72 hours.

[0020] In some implementations, rapid dissolution of certain sacrificial yarns or elements within the suture construct causes the suture to loosen and elongate, thus adapting to tissue swelling. Later crystallization of amorphous or semi-amorphous suture materials causes the suture to contract, thereby tightening as swelling subsides or healing progresses.

[0021] In one embodiment, the present invention relates to a length-adaptive surgical suture comprising a braid of monofilaments or multiple filaments having an initial length at implantation and a second length different from the initial length for the first twenty-four (24) hours after implantation to accommodate tissue swelling. The suture may be in the form of at least one filament from the braid of monofilaments or multiple filaments having a plurality of releasable loops held in place by an adhesive dissolved in the presence of physiological fluids or biocompatible solvents. Alternatively, the suture may be multiple filaments held in a contracted state at an initial length shorter than the second length by shrinkable yarns, wherein such shrinkable yarns dissolve rapidly upon exposure to bodily fluids and tissues. In another alternative, the suture may be in the form of multiple filaments arranged in a braid, wherein shrinkable yarns form the core of the braid and hold the braid in a contracted shortened state, wherein the shrinkable yarns dissolve rapidly upon exposure to bodily fluids. In yet another alternative, the suture may be in the form of multiple filaments arranged in a braid, held in a tightly shortened configuration by a biocompatible soluble adhesive. In another alternative, the suture may be in the form of at least one filament held in a wavy configuration at an initial length by a biocompatible soluble adhesive or coating.

[0022] The invention also relates to length-adaptive surgical sutures comprising a braid of monofilaments or multiple filaments having an initial length at implantation and a second length shorter than the initial length. Shortening suture embodiments may be in the form of a braid of monofilaments or multiple filaments. In one aspect of shortening the suture, the second length is at least 10% smaller than the initial length during the first twenty-four (24) hours after implantation.

[0023] In some implementations, the diameter of each filament increases by 5% to 10% in vivo after implantation.

[0024] In at least one embodiment, at least some of the filaments in the suture are extruded from a semi-crystalline polymer having a glass transition temperature between 40°C and 55°C and exhibiting a crystallinity level between 10% and 25%.

[0025] Each embodiment of the adaptive suture may be made from one or more filaments of an absorbable polymer, a non-absorbable polymer, or blends thereof. Additionally, each filament in the monofilament or multiple braids may possess suture tensile strength and knot strength that, after implantation, each increases in vivo by at least approximately 20% before significant degradation and hydrolysis of the suture.

[0026] This invention relates to an adaptive surgical suture comprising a semi-crystalline absorbable synthetic polymer made of a copolymer of glycolide and lactide, the synthetic polymer having an initial length, a glass transition temperature between 40°C and 45°C, and a crystallinity level of about 15%, wherein the suture has a second length that decreases by about 15% after implantation and a diameter that increases by about 10% in vivo after implantation, and wherein the tensile strength and knot strength of the suture increase by at least about 40% in vivo after implantation prior to significant degradation and hydrolysis of the suture.

[0027] The present invention relates to a length-adaptive surgical suture comprising a braid of monofilaments or multiple filaments having an initial length at implantation, a second length longer for the first twenty-four (24) hours after implantation to accommodate tissue swelling, and a third length less than the initial length to tighten and retain tissue during healing.

[0028] The suture of the present invention may have at least one elongated filament held in place by a rapidly dissolving shrinkage element to shorten or shrink the suture, or wherein the elongation of the suture is caused by the rapid dissolution of the shrinkage element, which causes the elongated filament to relax and elongate, or wherein the at least one elongated filament comprises a semi-crystalline polymer having a glass transition temperature between 40°C and 55°C and exhibiting a crystallinity level between 10% and 25%, wherein the crystallization of the semi-crystalline polymer causes the suture to subsequently shrink, thereby shortening the suture. Attached Figure Description

[0029] Figure 1 and Figure 2 A schematic diagram is shown showing the post-implantation suture length as a percentage of the initial suture length versus time expressed in hours.

[0030] Figure 3 A schematic diagram showing the suture length of the present invention after installation is shown.

[0031] Figures 4A to 4F An embodiment of the invention involving the elongation step is shown.

[0032] Figure 5 This is a flowchart of a novel method of the present invention for producing the novel suture of the present invention.

[0033] Figure 6 The data shows the percentage of suture length that has changed over time (hours) relative to the initial length.

[0034] Figure 7 The data shows the change in suture diameter over time (hours).

[0035] Figure 8 The data shows the suture strength over time (hours).

[0036] Figure 9 The data shows the change in suture knot strength over time (hours). Detailed Implementation

[0037] See now Figure 1 and 2 A diagram is drawn showing the post-implantation suture length as a percentage of the initial suture length versus time expressed in hours. The zero point of time represents the surgical implantation of the suture and its exposure to tissues and fluids. Dashed line 10 represents a routine suture with no significant change in length after implantation, i.e., a length of 100% of the initial pre-implantation length over the entire time shown. Figure 1 Line 20 illustrates one embodiment of the suture of the present invention, which begins to elongate by about 10% after installation and completes elongation within about 24 hours. Figure 1 Line 30 illustrates one embodiment of the suture of the present invention, which begins to shrink by about 10% after installation and completes the shrinkage within about 80 hours.

[0038] Figure 2 Line 40 illustrates one embodiment of the suture of the present invention, which begins to elongate by about 10% after installation, reaches about 110% of its initial length within about 24 hours, and then shrinks back to approximately the same length as 100% of its original pre-installation length within about 80 hours.

[0039] Figure 2 Line 50 illustrates one embodiment of the suture of the present invention, which begins to elongate by about 10% after installation, reaches about 110% of its initial length within about 24 hours, and then shrinks back to approximately 90% of its original pre-installation length within about 80 hours.

[0040] See now Figure 3 Several embodiments of the suture of the present invention are schematically illustrated. In the elongation-only embodiment, corresponding to Figure 1 Thread 20, the installed suture 100 is shown as elongating after installation to form an elongated suture 100a. In the shrinkage-only embodiment, corresponding to... Figure 1 Thread 30, the installed suture 100 is shown as shrinking after installation to form a shorter suture 100b. In the embodiment of elongation and return to the initial length, corresponding to Figure 2The suture 100, after installation, is shown elongating to form an elongated suture 100a, and subsequently contracting to form a suture 100c having approximately the same length as the original suture 100. In the embodiment where the suture is elongated and contracted to a length shorter than the initial length, corresponding to... Figure 2 The suture 50, the installed suture 100 is shown as elongating after installation to form an elongated suture 100a, and is subsequently shown as contracting to form a suture 100d shorter than the original suture 100.

[0041] In some embodiments, the sutures of the present invention are optionally longitudinally elongated or extended during installation to accommodate tissue swelling. As described below, the elongation of the sutures is achieved through the rapid dissolution of the components.

[0042] In one implementation scheme, such as Figure 4A As shown, for braided or monofilament sutures, the small suture loops are held in place by a biocompatible, soluble adhesive. Upon exposure to body fluids and tissues, the loops are released, causing the suture to elongate after installation.

[0043] In another implementation scheme, such as Figure 4B As shown, the suture of the present invention is formed of multiple yarns or filaments twisted together in a uniform construct (such as a braid), wherein shrinkable yarns or filaments hold the construct in a contracted or shortened state, and such shrinkable yarns or filaments are rapidly soluble or absorbent. Upon exposure to bodily fluids and tissues, the shrinkable yarns (solid threads) rapidly dissolve and release the remaining construct (shown as double threads), thereby allowing the suture to extend longitudinally.

[0044] In another implementation scheme, such as Figure 4C As shown, the suture of the present invention is formed from multiple yarns or filaments forming a uniform construct (such as a braid), wherein shrinkage yarns or filaments (solid threads) form the core of the braid and hold the braid in a shrunken or shortened state. Such shrinkage core filaments are rapidly soluble or absorbable. Upon exposure to bodily fluids and tissues, the shrinkage filaments rapidly dissolve and release the remaining braid construct (shown as a double thread), thereby allowing the suture to extend longitudinally as the core dissolves.

[0045] In another implementation scheme, such as Figure 4D As shown, the suture of the present invention is formed from multiple yarns or filaments forming a uniform construct (such as a braid), wherein shrinkage yarns or filaments (solid threads) form part of the braid that is not part of the core of the braid and keep the braid in a shrunken or shortened state. Such shrinkage core filaments are rapidly soluble or absorbable. Upon exposure to bodily fluids and tissues, the shrinkage filaments rapidly dissolve and release the remaining braid construct (shown as a double thread), thereby allowing the suture to extend longitudinally as the core dissolves.

[0046] In another implementation scheme, such as Figure 4E As shown, the suture of the present invention is a braid (shown as a double thread) that is tightly assembled and held together by an adhesive (shown as a translucent solid) to shorten the structure. When exposed to bodily fluids and tissues, the adhesive rapidly dissolves and releases the remaining braided construct, thereby allowing the suture to extend longitudinally.

[0047] In another implementation scheme, such as Figure 4F As shown, the suture of the present invention is a monofilament or braid, held in a wavy configuration by an adhesive (shown as a translucent solid) or adhesive coating. Upon exposure to bodily fluids and tissues, the adhesive rapidly dissolves and releases the construct (shown as a double thread), thereby allowing the suture to extend longitudinally.

[0048] In all embodiments, rapidly soluble or rapidly absorbable materials are selected to begin dissolving immediately or within hours after tissue loading and to be completely reabsorbed and / or dissolved within 16 to 48 hours, preferably within 24 hours. Materials include pre-hydrolyzed polyesters, polyethylene glycol, polyvinyl alcohol, cellulose-based materials, proteins (such as albumin), and any biocompatible, rapidly soluble, or rapidly bioresorbable materials.

[0049] Following the optional elongation step, the sutures of the present invention undergo shortening or contraction to accommodate the reduction of tissue swelling or other tissue healing processes. It should be understood that suture contraction will begin immediately upon exposure to bodily fluids and tissues, but due to the nature of the elongation process, the effects of the optional elongation will be observed earlier. Therefore, while optional suture elongation may be completed within 12 to 24 hours, suture contraction may be completed within 48 to 72 hours, or even within 120 or 240 hours.

[0050] The homopolymers and copolymers of the present invention can be melt-extruded using a variety of conventional methods. Monofilament fibers can be formed by melt extrusion followed by stretching of the extrudate, with or without annealing. Multifilament fibers can be formed using conventional methods. Methods for manufacturing monofilament and multifilament braided sutures are disclosed in U.S. Patent No. 5,133,739, entitled “Segated Copolymers of epsilon-Caprolactone and Glycolide,” and U.S. Patent No. 6,712,838, entitled “Braided Suture with Improved Knot Strength and Process to Produce Same,” the entire contents of which are incorporated herein by reference.

[0051] For monofilament sutures, dimensional stability is conventionally required. Therefore, as the molecular orientation of the polymer increases during fiber processing to enhance strength, the driving forces for shrinkage and deformation increase. This is particularly problematic for polymers with low glass transition temperatures used in the preparation of monofilament sutures. Therefore, dimensional stability is provided for all conventional sutures through rapid crystallization of the polymer during fiber formation. A common approach in fiber processing is to prevent shrinkage or deformation of highly oriented monofilament sutures by producing sufficient levels of crystallinity. However, this invention describes an innovative procedure that provides fibers with only a limited amount of crystallinity, which helps maintain the dimensional stability of the suture before use but allows it to change shape in a desired pattern upon post-implantation use.

[0052] Examples of the novel monofilament extrusion of the present invention for use with glycolide and lactide copolymers are shown in Figure 5 The flowchart is shown in a simplified manner. As mentioned above, both absorbent and non-absorbent polymer compositions can be used to practice the techniques of the present invention. A polymer resin 10 in granular or pellet form, suitable for manufacturing the novel sutures of the present invention, is fed into a nitrogen-purged hopper 22 of an extruder 20. The polymer 10 is then heated, melted, and processed by the extruder 20 to bring it into a flowable state. The molten polymer 10 is then discharged from the outlet 25 of the extruder 20 as a monofilament extrudate 30.

[0053] The fiber extrudate 30 is drawn into a quenching bath 40 containing cooling water 45; although a wide range of quenching bath temperatures can be used, a temperature of 20°C is particularly preferred. The “air gap” 27 (which is the distance between the bottom of the extruder outlet 25 and the water surface 48 of the quenching bath 40) will typically range from about 0.05 inches to about 2.0 inches. An air gap of about 0.1" to about 0.5" is particularly preferred. The fiber extrudate 30 is held in the aqueous medium 45 of the quenching bath 40 for a sufficient time to effectively provide the desired polymer morphology of the polymer chains of the fiber extrudate 30, particularly crystal nucleation. The residence time of the extrudate fiber 30 in the aqueous medium 45 of the quenching bath 40 is controlled by the linear velocity of the extrudate fiber 30 and the path length of the extrudate fiber 30 within the aqueous medium 45. The residence time of the extrudate fiber 30 in the aqueous medium 45 of the quenching bath 40 is typically from about 1 second to several minutes (e.g., about 3 minutes). The crystallinity level of fiber 30 after exiting quench bath 40 needs to be sufficiently low, i.e., sufficiently low, to allow a maximum draw ratio of at least 8.0 during the subsequent drawing step. If the crystallinity level of fiber 30 leaving bath 40 is too high, the fiber will break if a higher draw ratio is attempted, thus limiting molecular orientation and therefore the tensile strength of the fiber. The fiber extrudate 30 is then moved to the rollers of guide roller 60 and then to the rollers of guide roller 70; the linear velocity of guide roller 70 relative to guide roller 60 is greater than or equal to about 5.5. The fiber 31 being drawn between guide rollers 60 and guide roller 70 is thus drawn to a ratio greater than or equal to about 5.5. Guide rollers 60 and 70 may optionally be heated (e.g., from about 30°C to 90°C) to allow for smoother drawing.

[0054] like Figure 5As shown, the filament 32 is then moved to a first hot air oven 80, where it is heated to a sufficiently effective temperature (from about 30°C to 130°C) and a sufficiently effective residence time to provide sufficient crystal growth in the fiber 33. The residence time in the hot air oven 80 is controlled by the linear velocity of the fibers 32 / 33 and the path length of the fibers within the hot air oven 80. The fiber 32 exiting the guide roller 70 may undergo further stretching by employing a guide roller 90, in addition to optionally undergoing heat treatment in the hot air oven 80. The stretch ratio between the guide roller 90 and the guide roller 60 is typically about 6.5 or greater. Optionally, the roller of the guide roller 90 may be heated (from about 30°C to 90°C) to facilitate stretching. The filament 34 exiting the optional guide roller 90 is then moved to a second hot air oven 100 and reheated to a sufficiently effective temperature (from about 30°C to 130°C) and a sufficiently effective residence time to obtain the optimal polymer morphology for producing the stretched fiber 35. The fibers 34 exiting the guide roller 90, in addition to optionally undergoing heat treatment in the hot air oven 100, can also undergo further stretching or relaxation using the guide roller 110 to produce stretched fibers 36 again. The stretch ratio between the guide roller 90 and the guide roller 110 is typically from about 0.8 to about 1.2. The resulting filaments 36 are then moved to the take-up unit 120, where they are collected on roller 125. The filaments 36 collected on roller 125 can be stored in a nitrogen or vacuum chamber until further use.

[0055] During extrusion, if the polymer's glass transition temperature is above room temperature, the fiber can be stretched without inducing any measurable crystallinity. The fiber orientation achieved during stretching can be maintained by giving the fiber a sufficiently high glass transition or inducing a specific level of crystallinity to preserve its morphology during storage.

[0056] If desired, the homopolymers and copolymers of the present invention can be processed to have barbs when made into monofilament sutures. Such barbs can be placed or joined in conventional ways, including cutting, molding, preforming, shaping, attaching, etc. An example of a method for forming barbs is disclosed in U.S. Patent 8,216,497, “Tissue Holding Devices and Methods for Making the Same,” which is incorporated herein by reference. An alternative method for preparing barbed sutures is a cutting method. An example of a barb cutting method is disclosed in U.S. Patent 7,913,365, “Method of Forming Barbs on a Suit and Apparatus for Performing the Same.”

[0057] If desired, the monofilament or multifilament sutures of the present invention may contain medically usable materials. Medically usable materials may be incorporated into or onto the sutures in a variety of conventional methods, including mixing, coating, spraying, impregnation, sputtering, etc. The sutures of the present invention can be delivered to surgeons in various lengths. Preferably, conventional surgical needles are attached to one or both ends of the suture (i.e., one or both arms), but the sutures may not be equipped with any attached surgical needles.

[0058] Medically usable substances that can be incorporated into the surgical sutures of the present invention include antimicrobial agents, therapeutic agents, antibiotics, and other components or reagents.

[0059] The process equipment that can be used in the methods of this invention will be readily available commercial equipment. An example of a monofilament extruder that can be used to practice this invention is a Davis-Standard extruder, model 22450, purchased from Davis-Standard, Cedar Grove, NJ, USA. An example of a guide roller that can be used to practice this invention is a JJ Jenkins guide roller, model 9397, purchased from JJ Jenkins, Inc., Matthews, NC, USA. An example of a hot air oven that can be used to practice this invention is a JJ Jenkins oven, model 74DM1115-3, purchased from JJ Jenkins, Inc., Matthews, NC, USA. An example of a take-up unit that can be used to practice this invention is a Georg Sahm take-up unit, model 283E, manufactured by Georg Sahm GmbH, Eschwege, Germany.

[0060] The key properties of the polymer fibers prepared to support this application were measured using the different characterization methods described below.

[0061] Differential scanning calorimetry (DSC), tensile and other analytical data were used to fingerprint the produced fibers. The DSC instrument used was a TA Instruments (New Castle, DEUSA) Model Q20 equipped with a 50-position robotic automated sampler.

[0062] Mechanical properties of fibers, such as straight tensile strength and knot tensile strength (with a simple knot in the middle), were measured using an Instron testing machine before and after hydrolysis. The Instron model ID#TJ-41 was equipped with a 100-pound load cell LC-147 with pneumatic grippers, providing a gripping pressure of approximately 60 psi. The standard Instron speed was 1 inch / minute, with a gauge length of 1 inch. A 100-pound load cell was used. For zero time, a steel face was used on the Instron machine; for all other hydrolysis times, a rubber face was used to prevent slippage. Fiber diameter was measured using a Federal Gauge (Products Corp. Providence, RI) model 57B-1, identification number W-10761.

[0063] In vitro BSR measurements of the produced fibers were performed under physiologically relevant in vitro conditions: a 0.01 M (1X) phosphate buffer solution at pH 7.27 was maintained at 37°C. Two Haake water baths equipped with ThermoScientific DC10 motors (model W46, device ID: BT-029) were used. BSR assessment data are given in pounds and percentages. Tensile strength of the samples was tested at specified time points using an Instron material testing machine. Test parameters were a gauge length of one inch and a clamping speed of one inch per minute.

[0064] The following examples illustrate the principles and practice of the invention, but are not limited thereto. Many other embodiments within the scope and spirit of the invention will become apparent to those skilled in the art once this disclosure has been received.

[0065] Example 1. Monofilament extrusion of 90 / 10Gly / Lac (PG-910) copolymer

[0066] Monofilament extrusion runs of the 90 / 10 Gly / Lac (PG-910) random copolymer suitable for carrying out the present invention were performed using a one-inch Davis-Standard extruder with a 24:1 barrel length (1-22-1 design) and a single-groove feed throat. The die size for each run was 40 / 1. The random 90 / 10 Gly / Lac copolymer is as described above and Figure 5 The process is as shown. A water bath capable of heating to approximately 50°C is used, along with three sets of guide rollers (with heating capabilities), one additional relaxation guide roller located at the end of the production line before the collection spool. Two annealing furnaces are located between the second and third guide rollers, and between the third guide roller and the final relaxation guide roller, for heating the fibers to enhance their polymer morphology. The fiber diameter is measured online using a Mitutoyo Lasermac (laser micrometer) located before the collection spool.

[0067] In this example, the bath temperature was set to 20°C for each copolymer run, and the guide roller speed ratios are listed in Table 1 below. The monofilament stitches have a 3-0 stitch USP size, corresponding to a fiber diameter of approximately 11 mils. The extruder temperature ranged from 205°C to 220°C, and the die temperature for all runs was maintained at 220°C.

[0068] Table 1 .

[0069] Selected extrusion conditions for 90 / 10Gly / Lac size 3-0 monofilament

[0070]

[0071] The extruded 90 / 10 Gly / Lac 3-0 monofilaments of this invention are then collected on a spool and stored in a vacuum chamber. Before use in the next stage, the monofilaments are cut into 12-inch sheets.

[0072] Example 2. Stretching and Crystallization Properties

[0073] The Instron tensile properties of the fibers produced in Example 1 are given in Table 2.

[0074] Table 2 .

[0075] Instron tensile properties of stretched 3-0 monofilament in Example 1

[0076]

[0077] *Each fiber sample is 12.0 inches in length.

[0078] To estimate the crystallinity level of the fibers given in Table 1, a first thermal DSC measurement was performed to measure the properties of the “original” fibers after extrusion. The samples were first quenched to -60°C at a rate of 60°C / min, followed by heating at a rate of 10°C / min to determine their calorimetric properties (first thermal properties); these included the glass transition temperature T0. g Crystallization temperature T c , heat of crystallization ΔH c Melting Point T m and heat of fusion ΔH m The heat of fusion is directly proportional to the level of crystallinity in the sample. If needed, the values ​​of Tg, Tm, Tc (crystallization temperature), and ΔHm can be obtained from a second thermal measurement (the resin is melted at 220°C and then quenched at below -60°C), which are independent of the previous heat treatment history.

[0079] Table 3 summarizes the DSC measurements of the fibers produced in Example 1 for their selected first thermal properties.

[0080] Table 3 .

[0081] The crystallization characteristics of the stretched 3-0 monofilament in Example 1 were estimated by DSC.

[0082]

[0083] *The fiber has a draw ratio of 7.5x.

[0084] **Stretch ratio is 5.5x**

[0085] ***The percentage of crystallinity is calculated from the heat of fusion (ΔH) of 100% crystallized PGA material. m =12 kJ / mol, equivalent to 103 J / g); [References: "Biomedical Engineering Fundamentals", Joseph D. Bronzino, Donald, and R. Peterson; "Wound Closure Biomaterials and Devices", edited by Chih-Chang Chu and J. Anthony von; "Biomaterials: Principles and Practices", edited by Joyce Y. Wong, Joseph D. Bronzino, and Donald R.; "Biotextiles as Medical Implants", edited by MW King, B. SGupta, and R. Guidoin; "The Biomedical Engineering Handbook 1", Joseph D. Bronzino; "Surfaces and Interfaces for Biomaterials", edited by P. Vadgama]

[0086] The data in Table 3 demonstrate a strong dependence of oven air temperature on the level of crystallinity formed in the extruded fibers. For example, regardless of the draw ratio used (5.5x to 7.5x), oven air temperatures of 80°C and above produced significantly higher levels of crystallinity (above 40%) in 90 / 10 Gly / Lac fibers at oven temperatures below 80°C. Interestingly, when the oven temperature was only 40°C, a very low level of crystallinity (approximately 7%) appeared to form in the 90 / 10 Gly / Lac fibers studied.

[0087] Importantly, it is also noted that for each fiber run in Table 3, the glass transition temperature was found to be in the range of 41°C to 47°C. The higher the crystallinity level in the fiber (e.g., samples 1A-1C, 2), the higher the observed glass transition temperature. This is because when the lactide portion crystallizes, the amorphous phase, which contributes only to the glass transition temperature, becomes rich in rigid lactide moieties. This, in turn, increases the overall glass transition temperature of such fibers.

[0088] Example 3. Annealing furnace treatment of fibers produced in Example 1

[0089] The 3-0 monofilament produced in Example 1 was placed in an annealing oven at a preset temperature of 37°C and held there for several hours or days. After specific time intervals, the sutures were removed from the oven and examined to observe whether any changes in length, diameter, or tensile properties occurred during the exposure in this "drying" oven. This study will contribute to establishing the stability of such sutures under environmental conditions up to 37°C.

[0090] The data after exposure to a 37°C nitrogen oven are summarized in Table 4.

[0091] Table 4 .

[0092] After exposure in a nitrogen oven at 37°C for different times, the physical and Instron properties of the stretched 3-0 monofilament of Example 1 were compared. Tensile properties

[0093]

[0094] *The first number in each column represents a 6-hour exposure time, the second number represents a 24-hour exposure time, and the third number represents a 120-hour exposure time.

[0095] As shown in Table 4, extreme shrinkage was observed in amorphous sample 3 after exposure to a nitrogen oven at 37°C for 6, 24, and 120 hours. The extrudate was found to be “rolled up and wavy,” making it very difficult to measure length and diameter, and instron testing was not possible. Sample 4 showed only minor changes (approximately 10%) in fiber length and diameter, while other fibers (samples 1A, 1B, 1C, and 2) showed almost no change in these properties.

[0096] Example 4. Breaking strength of fibers produced in Example 1 exposed to physiological conditions (37°C / pH = 7.27). Retention rate, BSR and physical properties

[0097] In this embodiment, the fibers produced in Example 1 were placed in a buffer solution at 37°C and pH 7.27 at different time intervals to simulate human physiological conditions. Before exposure to the buffer solution, the 12" fiber sample was kept under vacuum at room temperature for approximately three weeks.

[0098] After being placed in the buffer solution for a specific time, a portion of the fiber was removed and a series of parameters were measured to estimate the hydrolytic effect on the fiber during the first seven days of exposure: length; diameter; straight tensile strength; knot tensile strength.

[0099] Tables 5A, 5B, 5C, and 5D below summarize the buffer hydrolysis data, focusing particularly on the parameters listed above. As shown in the main Instron tensile property data (5C and 5D) below, significant degradation / hydrolysis occurs on day 7. For the purposes of this invention, particular attention is paid to the data points up to day 3, where the changes in fiber properties are attributed solely to solvation into the polymer matrix without the primary effect of chemical hydrolysis.

[0100] Table 5-A .

[0101] The fiber length of the 3-0 monofilament described in Example 1 after exposure to a buffer solution at 37°C and pH 7.27 Degree measurement results; data represent the average of eight tests conducted.

[0102]

[0103] The data in Table 5A show that the fiber length of samples 1A, 1B, 1C, and 2 remained almost unchanged. This indicates a stable morphology, or a sufficiently high level of crystallinity, to prevent the fibers from shrinking for up to 3 days under physiological conditions. On the other hand, samples 3 and 4 showed measurable shrinkage induced by buffer exposure, at 31% and 14%, respectively.

[0104] The same data is also presented as a chart in Figure 6 The data presented shows the percentage of the original length of the suture as it changes over time (hours).

[0105] Table 5-B .

[0106] The 3-0 monofilament described in Example 1 after exposure to a buffer solution at 37°C and pH 7.27 showed fiber straightness. Diameter measurement results; data represent the average of eight tests conducted.

[0107]

[0108] The fiber diameter variations presented in Table 5B show a similar trend. While the diameters of samples 1A, 1B, 1C, and 2 remained virtually the same, the diameters of samples 3 and 4 increased by 19% and 7.5%, respectively.

[0109] The same data is also presented as a chart in Figure 7 The data presented shows the suture diameter over time (hours).

[0110] Table 5-C .

[0111] The fiber strength of the 3-0 monofilament described in Example 1 after exposure to a buffer solution at 37°C and pH 7.27 Degree measurement results; data represent the average of eight tests conducted.

[0112]

[0113] The fiber strength data in Table 5C show a significant trend. For samples 1A, 1B, 1C, and 2 with higher crystallinity, almost no change was observed in the buffer solution during the first few days. On day 3, a slight decrease in strength was observed in these samples (13% to 20%), which is likely due to the earlier occurrence of hydrolysis. However, most unexpectedly, samples 3 and 4 showed a significant jump in fiber strength for the same time period, at 60% and 47%, respectively.

[0114] The same data is also presented as a chart in Figure 8 The data presented here shows the fiber strength over time (hours).

[0115] Table 5-D .

[0116] The fiber knots of the 3-0 monofilaments described in Example 1 after exposure to a buffer solution at 37°C and pH 7.27 Strength measurement results; data represent the average of eight tests conducted.

[0117]

[0118] The fiber knot strength data in Table 5D simulate the same unexpected behavior as sample 1A, which showed a slight increase in tensile knot strength of about 9% on day 3. The remaining samples with higher crystallinity, 1B, 1C, and 2, showed a significant decrease in knot strength (13% to 30%) on day 3, while samples 3 and 4 showed the opposite trend: huge increases in knot strength of 69% and 56%, respectively.

[0119] The same data is also presented as a chart in Figure 9 The data presented here show the fiber knot strength over time (hours).

[0120] To summarize the findings in Tables 5A-D, fiber samples 1A, 1B, 1C, and 2, with higher levels of crystallinity, behaved similarly to other conventional sutures. Their length, diameter, and major tensile properties appeared unaffected by the physiological buffering conditions of early hydrolysis. These can be considered samples not of this invention.

[0121] The amorphous fiber of Example 3 was significantly affected by exposure to physiologically buffered conditions. Although a significant improvement in tensile properties (straight strength and knot strength) was observed for this sample, the excessively high shrinkage rate and extensive diameter increase prevented its usability in practical suture applications. Furthermore, the "drying" oven data at 37°C given in Table 4 indicate stability and shelf-life issues, as significant fiber deformation was observed.

[0122] We found that sample 4, with its limited crystallinity, exhibited the optimal combination of properties suitable for practicing the techniques of this invention. Following exposure of the suture to a physiological buffer solution at 37°C and pH 7.27, and before significant hydrolysis began, the sample shrank to approximately 15% of the desired level, with a slight increase in diameter of 5% to 10%, while its straight strength and knot strength values ​​increased significantly by approximately 50%.

[0123] The following are examples of potential sutures currently on the market, in which practicing the techniques of this invention will not produce the desired characteristics targeted by this invention.

[0124] Example 5: Crystallization characteristics of polypropylene and polyvinyl difluoroethylene (PVDF) based resins

[0125] In this section, we will describe two polymer systems that cannot be used to practice the teachings of this invention.

[0126] Polypropylene-based sutures, such as those sold by Ethicon Inc., have a melting point of approximately 165°C. Extruder die temperatures for these polymers range from 192°C to 218°C, with approximately 200°C recommended. For the internal extrusion process of the suture (PS-0000028), the recommended temperature range for the orientation oven (oven #1) is 125°C to 145°C (257°F to 293°F) for size 3-0 fibers, while the range for the annealing / relaxation oven (oven #2) is slightly higher, from 135°C to 165°C (275°F to 329°F). To simulate the crystallization behavior of the polymer during extrusion, we employed differential scanning calorimetry (DSC) as follows.

[0127] A polypropylene resin sample was placed in a DSC apparatus and brought above its melting point to produce all crystals (200°C). The sample was then quenched to -60°C, initially rapidly cooled at a cooling rate of approximately 60°C / min. During quenching, the sample crystallized immediately at a temperature of approximately 115°C. Subsequent heating scans at 10°C / min showed no further crystallization, indicating complete crystallization during this rapid quenching. In a second experiment, the polypropylene resin sample was quenched by placing the molten sample on dry ice and holding it there (far below -60°C) for several minutes. Subsequent heating scans at 10°C / min again showed no further crystallization, indicating complete crystallization during dry ice quenching. These sets of experiments on the polypropylene resin demonstrate that the polymer crystallizes too rapidly to be used in practicing the techniques of this invention.

[0128] The resin formulation for poly(hexafluoropropylene-VDF) sutures, also supplied by Ethicon Inc., is a blend of approximately 50% polyvinylidene fluoride (PVDF) homopolymer and 50% copolymer (95% PVDF and 5% hexafluoropropylene). Another composition, the 80 / 20 resin for PRONOVA sutures, consists of approximately 80% polyvinylidene fluoride (PVDF) homopolymer and 20% copolymer (95% PVDF and 5% hexafluoropropylene). For the purposes of this invention, we examined the crystallization characteristics of the first composition, the 50 / 50 resin, expecting it to exhibit a slower crystallization rate than the 80 / 20 formulation primarily used for smaller suture sizes. We repeated the DSC melting and quenching steps as described in the preceding paragraphs for polypropylene. After examining the second heat scan of the 50 / 50 resin using DSC, we observed no additional crystallization in the thermal recorder. During quenching (approximately 130°C), the sample crystallized extremely rapidly again, leaving no amorphous portion for the second heat scan.

[0129] In addition, we conducted melt index testing, during which we produced fiber extruders from 50 / 50 resin and instantaneously stretched them by two, four, or seven times their initial length. As a control, we also tested unstretched samples. Both stretched and unstretched fiber sheets were immediately placed on dry ice to quench the samples at -60°C for five minutes. Subsequent DSC heating scans at 10°C / min showed no crystallization in any of the tested samples. The glass transition temperature was approximately -33°C, the melting point was approximately 172°C, and the heat of fusion for all samples was approximately 55 J / g, a value very close to that of the 50 / 50 resin in the first heating scan. These measurements indicate that the crystallization rate of this polymer is too high for the resin to be used in the technology of this invention.

[0130] However, non-absorbable sutures made of nylon 6 polymer and nylon 6,6 polymer Extra polyester sutures and The nylon sutures meet the requirements and are excellent candidate sutures for practicing the techniques of this invention. They have fiber glass transition temperatures in the range of 40°C to 50°C and exhibit excellent water absorption after implantation. The nylon 6 polymer exhibits lower crystallinity and a slower crystallization rate than its nylon 6,6 counterpart.

[0131] The field of this invention is semi-crystalline synthetic sutures made of absorbent or non-absorbent polymers having a glass transition temperature between 40°C and 55°C and exhibiting a limited crystallinity level between 10% and 25%. Ideally, the polymer needs to exhibit sufficiently slow crystallization kinetics during monofilament or multifilament extrusion to allow for precise control over the crystallinity level of the resulting fibers. Data from this study indicate that 90 / 10Gly / Lac monofilaments exhibit a 14% reduction in fiber length and an 8% increase in fiber diameter after exposure to human physiological conditions. The primary 90 / 10Gly / Lac composition of this invention exhibits a crystallinity level of approximately 15% and, after exposure to human physiological conditions, shows a significant increase in straight tensile strength and knot strength of approximately 50% before substantial hydrolysis occurs.

[0132] Advantageously, the sutures of the present invention exhibit surprising properties, including: shrinkage of the suture in vivo by about 10% to 15% after implantation to reduce swelling of surrounding tissues by adjustment; an increase in the diameter of the suture in vivo by about 5% to 10% after implantation, thereby allowing closure of the gap between the hole formed by the difference in diameter between the needle and the suture; a significant increase in the main tensile properties of the suture (such as straight strength and knot strength) in vivo by about 50% before significant hydrolysis occurs after implantation; and suture crystallization / morphology is controlled by extrusion processing parameters without changes in chemical composition.

[0133] Although the invention has been shown and described with respect to detailed embodiments thereof, those skilled in the art will understand that various changes in form and detail may be made to the invention without departing from the spirit and scope of the claimed invention.

Claims

1. A length-adaptive surgical suture comprising a braid of monofilaments or multiple filaments, said suture having an initial length at implantation, a second length different from said initial length for the first twenty-four (24) hours after implantation to accommodate tissue swelling, and a third length thereafter smaller than said initial length to tighten the tissue during healing, wherein, The suture comprises at least one elongated filament held in place by a rapidly dissolving shrinkage element to shorten or shrink the suture, wherein the elongation of the suture is caused by the rapid dissolution of the shrinkage element, which causes the elongated filament to relax and elongate, wherein the at least one elongated filament comprises a semi-crystalline polymer having a glass transition temperature between 40°C and 55°C and exhibiting a crystallinity level between 10% and 25%, wherein the crystallization of the semi-crystalline polymer causes the suture to subsequently shrink, thereby shortening the suture.

2. The suture according to claim 1, wherein, At least one of the filaments in the braid of the monofilament or the plurality of filaments has a plurality of releasable loops, which are held in place by an adhesive dissolved in the presence of a physiological fluid or a biocompatible solvent.

3. The suture thread according to claim 1, comprising a plurality of filaments held in a contracted state by shrinkage yarn at an initial length shorter than the second length, wherein, These types of shrink yarns dissolve rapidly when exposed to bodily fluids and tissues.

4. The suture according to claim 1, comprising a plurality of filaments arranged in a braided pattern, wherein, Shrink yarns form the core of the braid and keep the braid in a shortened state, wherein the shrink yarns dissolve rapidly when exposed to bodily fluids.

5. The suture of claim 1, comprising a plurality of filaments arranged in a braid, the plurality of filaments being held in a tightly shortened configuration by a biocompatible soluble adhesive.

6. The suture of claim 1, comprising at least one filament held in a wavy configuration at an initial length by a biocompatible soluble adhesive or coating.

7. The suture according to claim 1, wherein, After implantation, the diameter of each filament increases by 5% to 10% in the body.

8. The suture according to claim 1, wherein, Each of the monofilaments or the plurality of braids has suture tensile strength and knot strength, which, after implantation, increase in vivo by at least about 20% before significant degradation and hydrolysis of the suture.

9. The suture according to claim 1, wherein, The suture has an initial length at implantation and a longer second length for the first twenty-four (24) hours after implantation to accommodate tissue swelling.

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