Implantable stent for treating sinusitis

Abstract

The present invention particularly describes materials, devices, kits, and methods that can be used for the treatment of chronic rhinosinusitis.

Description

[0001] This application is a divisional application of Chinese Patent Application No. 201680036021.6 (PCT / US2016 / 040204), with a filing date of June 29, 2016, and an invention title of "Implantable Stent for Treating Sinusitis".

[0002] Cross - Reference to Related Applications

[0003] This application claims the benefit of U.S. Provisional Application No. 62 / 186,030, filed on June 29, 2015, with an invention title of Implantable Stent for Treating Sinusitis; U.S. Provisional Application No. 62 / 289,982, filed on February 2, 2016, with an invention title of Implantable Stent for Treating Sinusitis; and U.S. Provisional Application No. 62 / 332,134, filed on May 5, 2016, with an invention title of Implantable Stent for Treating Sinusitis, the entire contents of each of which are incorporated herein by reference. Field of the Invention

[0004] The present invention particularly describes substances, devices, kits, and methods that can be used to treat chronic rhinosinusitis. Background Art

[0005] Chronic rhinosinusitis (CRS) is a common disorder defined as symptomatic inflammation of the paranasal sinuses lasting longer than 12 weeks. Up to 16% of the population is affected by this disorder. The lumens associated with CRS include the maxillary sinus, frontal sinus, ethmoid sinus, ostiomeatal complex, ethmoid infundibulum, and sphenoid sinus, as well as the middle meatus location, or combinations thereof. Common symptoms of CRS include impaired nasal congestion, facial pressure or fullness, rhinorrhea, and anosmia; these symptoms may be due to mucosal inflammation, local infection, and / or impaired mucociliary function.

[0006] Although there are no approved therapies for treating CRS, evidence - based medical management supports the use of a large number of oral or topical corticosteroid therapies for this disease. Copious daily saline irrigation via nasal sprays in combination with adjunctive administration of topical corticosteroids is commonly used as a first - line therapy. Second - line agents for acute flare - ups and worsening disease include short - course oral corticosteroids, but this approach can lead to unintended systemic side effects, including glaucoma, osteoporosis, and avascular necrosis of the hip and shoulder. It is estimated that up to 12 - 50% of CRS patients do not respond positively to the recommended medical regimen and are often candidates for functional endoscopic sinus surgery (FESS) and / or balloon sinuplasty dilation.

[0007] For patients, it is desirable to avoid surgical intervention in the treatment of CRS because these procedures carry surgery-related risks, cause postoperative pain and discomfort, and require bothersome and costly postoperative irrigation. Clinical data have confirmed that topical corticosteroids effectively reduce inflammation associated with CRS and are thus a reasonable option for controlling the condition.

[0008] An ideal treatment for CRS would provide local sustained-release anti-inflammatory drug delivery within the patient's sinuses as an alternative treatment option to sinus surgery. Such a therapy should ideally establish safe and effective sustained-release drug delivery targeted to the inflamed tissue and, in some cases, can obviate the need for surgery.

[0009] FESS involves the removal of bone and tissue to widen the sinus outflow tract, broaden the sinus openings or ostia, and enable ventilation of previously obstructed sinus cavities and restoration of mucociliary clearance. Currently, approximately 500,000 surgeries are performed annually in the United States.

[0010] By removing small pieces of bone, polyps, and / or performing tissue debridement within the sinus cavity, FESS has been demonstrated to be an effective method for improving the sinus drainage pathway. However, a significant amount of postoperative complications are frequently observed, such as inflammation, swelling, disease recurrence, the need for repeat surgery, and adhesions. Therefore, postoperative care is an important component of FESS. Approximately 10 - 20% of FESS patients become refractory, unresponsive to treatment, and may require additional surgical intervention or lifelong medical treatment.

[0011] Some form of sinus packing is typically performed after FESS. Examples of packing materials include simple dressings moistened with saline, polysaccharide gel-based foams, PEG-based materials, and middle meatal spacers. Implantable sinus stents have been invented, which are used to stabilize the sinus ostia and turbinates through tissue adhesion, reduce edema, and / or prevent obstruction. They also have the ability to incorporate therapeutic agents that can be locally delivered over time. In the postoperative setting, local delivery of therapeutic agents may be superior to topical application. In this regard, the USFDA-approved PROPEL TM system (Intersect ENT, Menlo Park, CA, USA) is a self-expanding, bioabsorbable, steroid-eluting stent designed for use in the ethmoid sinuses after FESS. Brief Description of the Invention

[0012] As used herein, the terms "sinus" and "sinus cavity" refer to the sinus cavities and nasal cavity, which include, for example, the maxillary sinus, frontal sinus, and ethmoid sinuses, the ostiomeatal complex, the ethmoid infundibulum, and the sphenoid sinus, as well as the middle meatus (sinus cavity).

[0013] The present invention describes various sinus stents having fibrous and non - fibrous based designs. These design options differ in form, size, and delivery location (i.e., maxillary sinus, frontal sinus, ethmoid sinus, sphenoid sinus, and middle meatus). Additionally, therapeutic agents can optionally be included within the stent for short - term or long - term local delivery. Thus, these stents can be used, for example, to improve sinus ostium patency in surgically - adjusted sinus spaces or sinus spaces that have not yet been surgically adjusted. Moreover, these stents can be used to deliver local therapeutic agents to such sinus spaces, including, for example, as part of a treatment procedure that is an alternative to sinus surgery (e.g., FESS), or in certain embodiments, as part of the post - operative care of FESS in other situations.

[0014] In several aspects, the present invention relates to a generally tubular stent configured for implantation into a patient's sinus cavity. As used herein, "generally tubular" includes hollow shapes with circular cross - section or non - circular (e.g., oval, etc.) cross - section and hollow shapes with constant diameter or variable diameter (e.g., gradually decreasing diameter, such as a hollow frustum). Both ends of the generally tubular stent can be open, one end can be open and the other end can be closed, or both ends can be closed. In many advantageous embodiments described herein, a generally tubular stent is used, which is in the shape of a hollow cylinder (i.e., having a circular cross - section and a constant diameter), where both ends are open). The stent can have a fibrous or non - fibrous based structure and includes a stent material and an optional conformal coating, the conformal coating including a coating material that at least partially coats the stent material.

[0015] The stent material may or may not include a therapeutic agent, such as those selected from the therapeutic agents described elsewhere herein, among other possibilities.

[0016] When the stent includes a therapeutic agent, the stent can provide various release profiles.

[0017] In certain embodiments, when the following in vitro assay is performed: where the stent is submerged in a pH 7.4 PBS buffer solution containing 2% wt% SDS at 37°C under gentle shaking on a rotary shaker, the stent can exhibit some cumulative release characteristics, where the volume of the buffer solution in which the stent is submerged is at least 10 times the volume of the buffer solution at the saturation point (sometimes referred to as sink conditions) of the amount of the therapeutic agent corresponding to the total amount of the therapeutic agent in the stent, and where the buffer solution is completely removed and quantified weekly and replaced with fresh buffer solution.

[0018] In certain embodiments, after one week in such in vitro conditions, the scaffold can exhibit a cumulative release of the therapeutic agent ranging from 1% or less to 70% or more (e.g., in the range of 1% to 2% to 5% to 10% to 15% to 20% to 25% to 30% to 35% to 40% to 45% to 50% to 55% to 60% to 65% to 70%) based on the total amount of the therapeutic agent in the scaffold (i.e., in the range between any two of the foregoing values), advantageously in the range of 2% to 50%, more advantageously in the range of 5% to 30%.

[0019] Optionally or additionally, in certain embodiments, after two weeks in such in vitro conditions, the scaffold can exhibit a cumulative release of the therapeutic agent ranging from 5% or less to 80% or more (e.g., in the range of 5% to 7% to 10% to 15% to 20% to 25% to 30% to 35% to 40% to 45% to 50% to 55% to 60% to 65% to 70% to 75% to 80%) based on the total amount of the therapeutic agent in the scaffold (i.e., in the range between any two of the foregoing values), advantageously in the range of 7% to 50%, more advantageously in the range of 10% to 30%.

[0020] Optionally or additionally, in certain embodiments, after four weeks in such in vitro conditions, the scaffold can exhibit a cumulative release of the therapeutic agent ranging from 10% or less to 90% or more (e.g., in the range of 10% to 15% to 20% to 25% to 30% to 35% to 40% to 45% to 50% to 55% to 60% to 65% to 70% to 75% to 80% to 85% to 90%) based on the total amount of the therapeutic agent in the scaffold (i.e., in the range between any two of the foregoing values), advantageously in the range of 20% to 75%, more advantageously in the range of 30% to 60%.

[0021] Optionally or additionally, in certain embodiments, after eight weeks in such in vitro conditions, the scaffold can exhibit a cumulative release of the therapeutic agent ranging from 25% or less to 100% (e.g., in the range of 20% to 25% to 30% to 35% to 40% to 45% to 50% to 55% to 60% to 65% to 70% to 75% to 80% to 85% to 90% to 95% to 100%) based on the total amount of the therapeutic agent in the scaffold (i.e., in the range between any two of the foregoing values), advantageously in the range of 30% to 90%, more advantageously in the range of 40% to 80%.

[0022] In certain embodiments, the scaffold can exhibit some cumulative in vivo release characteristics.

[0023] For example, in certain embodiments, after one week in vivo in a human or rabbit sinus, the stent can exhibit a cumulative release of the therapeutic agent ranging from 1% or less to 45% or more (e.g., a range of 1% to 1.5% to 2% to 3% to 5% to 10% to 15% to 20% to 25% to 30% to 35% to 40% to 45%) (i.e., a range between any two of the foregoing values), advantageously in the range of 1.5% to 35%, more advantageously in the range of 3% to 20%, based on the total amount of the therapeutic agent in the stent.

[0024] Optionally or additionally, in certain embodiments, after two weeks in vivo in a human or rabbit sinus, the stent can exhibit a cumulative release of the therapeutic agent ranging from 3% or less to 50% or more (e.g., a range of 3% to 5% to 7% to 10% to 15% to 20% to 25% to 30% to 35% to 40% to 45% to 50%) (i.e., a range between any two of the foregoing values), advantageously in the range of 5% to 35%, more advantageously in the range of 7% to 20%, based on the total amount of the therapeutic agent in the stent.

[0025] Optionally or additionally, in certain embodiments, after four weeks in vivo in a human or rabbit sinus, the stent can exhibit a cumulative release of the therapeutic agent ranging from 7% or less to 60% or more (e.g., a range of 7% to 10% to 15% to 20% to 25% to 30% to 35% to 40% to 45% to 50% to 55% to 60%) (i.e., a range between any two of the foregoing values), advantageously in the range of 15% to 50%, more advantageously in the range of 20% to 30%, based on the total amount of the therapeutic agent in the stent.

[0026] Optionally or additionally, in certain embodiments, after eight weeks in vivo in a human or rabbit sinus, the stent can exhibit a cumulative release of the therapeutic agent ranging from 15% or less to 100% or (e.g., a range of 15% to 20% to 25% to 30% to 35% to 40% to 45% to 50% to 55% to 60% to 65% to 70% to 75% to 80% to 85% to 90% to 95% to 100%) (i.e., a range between any two of the foregoing values), advantageously in the range of 20% to 60%, more advantageously in the range of 25% to 55%, based on the total amount of the therapeutic agent in the stent.

[0027] In embodiments where the stent comprises a fibrous infrastructure, the stent can comprise a braided structure comprising one or more chains of the stent material.

[0028] In certain embodiments that can be used with any of the above aspects and embodiments, the braided structure can comprise opposing sets of helical chains. For example, each set of helical chains can comprise from 2 to 64 members, more typically from 8 to 32 members.

[0029] In certain embodiments that can be used with any of the above aspects and embodiments, the braided structure can include a first chain of a material having a first stiffness and a second chain of a material having a second stiffness greater than the first stiffness. As a particular example, the second chain of material can have a modulus > 3 GPa, which can be at least 2 times the modulus of the first chain of material.

[0030] In certain embodiments that can be used with any of the above methods and embodiments, the braided structure can include cells of different sizes. For example, a portion of the braided structure can be removed so that cells of different sizes can be formed, or the braided structure can be braided so as to form cells of different sizes, among other possibilities. In certain cases, the cells of different sizes can include a first cell having a first region and a second cell having a second region, where the first region is at least 50% greater than the second region. The variation in cell size can occur, for example, along the longitudinal length of the stent and / or around the circumference of the stent.

[0031] In certain embodiments that can be used with any of the above aspects and embodiments, the stent can include longitudinal elastic fibers that are mechanically coupled to two or more nodes of the braided structure.

[0032] In certain embodiments that can be used with any of the above aspects and embodiments, one or more ends of one or more chains can be braided back into and connected to the braided structure.

[0033] In certain embodiments, the general tubular stent of the present invention can include a stent material that includes elongate elements wound into a helical tubular structure. In certain of these embodiments, the elongate elements are in the form of strip-like elongate elements wound into a helical tubular structure. The strip-like elongate elements can be, for example, in the form of a solid film, or can include holes (e.g., formed by holes in a solid film, formed by intersecting fibers within a braided structure, etc.).

[0034] In certain embodiments, the general tubular stent of the present invention includes a stent material that includes a plurality of parallel open rings. In certain of these embodiments, the open rings are strip-like open rings. The strip-like open rings can be, for example, in the form of a solid film, or can include holes.

[0035] In certain embodiments, the general tubular stent of the present invention includes a stent material that includes a knitted structure. In certain of these embodiments, the knitted structure can include a single chain that can be pulled to unravel and remove the stent.

[0036] In some embodiments, the general tubular stent of the present invention may include a plurality of radially expandable inserts within a general tubular structure. In some of these embodiments, each radially expandable insert may include a center and a plurality of radially expandable arms, or may include a braided ring, among other possibilities.

[0037] In some embodiments that may be used with any of the above aspects and embodiments, the distal end of the general tubular stent may be configured to be captured by another device, and the general tubular stent may be configured to be inverted and removed by pulling the distal end into the lumen formed by the general tubular stent.

[0038] In some embodiments that may be used with any of the above aspects and embodiments, the general tubular stent may include a conformal coating, such as a conformal coating formed of an elastomeric or non-elastomeric coating material. For example, the coating material may be an elastomeric material comprising poly(L-lactide-co-ε-caprolactone) crosslinked with a urethane crosslinker, a urea crosslinker, or a urethane and urea crosslinker; the coating material may be an elastomeric material comprising a diisocyanate-cured (e.g., hexamethylene diisocyanate-cured, etc.), hydroxyl-terminated branched poly(L-lactide-co-ε-caprolactone). The coating material may or may not include a therapeutic agent, such as a therapeutic agent selected from those described elsewhere herein, among other possibilities. The coating material may cover alternating regions along the length of the general tubular stent, and / or the coating material may cover the ends of the general tubular stent without covering the regions between the ends of the general tubular stent, among other possibilities. In the case of a braided structure, the coating material may cover some of the nodes of the braided structure while leaving other nodes uncovered. The thickness of the coating material at the nodes of the braided structure may range from 1 to 100 times the thickness of the coating material between the nodes of the braided structure (e.g., any value in the range of 1 to 2 to 5 to 10 to 25 to 50 to 75 to 100 times the thickness of the coating material between the nodes of the braided structure).

[0039] In certain embodiments that can be used with any of the above aspects and embodiments, the stent can have a conformal coating, which can be further coated with a conformal coating comprising additional coating materials and therapeutic agents, for example, selected from therapeutic agents described elsewhere herein, and other possibilities. The thickness range of the additional conformal coating can be, for example, 1 μm to 25 μm (for example, a thickness range of 1 to 2 to 5 to 20 to 25 μm), and other possibilities. In certain embodiments, the additional coating material can be a biodegradable polymer such as poly (lactide-to-ε-caprolactone) or poly (lactide-to-ε-caprolactone) and an additional polymer such as a homopolymer or copolymer of lactide, such as a mixture of poly (lactide-to-glycolide). When included, the additional polymer can be present in an amount ranging from, for example, 5 to 50% by weight of the additional conformal coating. The poly(lactide-co-ε-caprolactone) can have, for example, a mole percentage range of 50 to 95% lactide and a mole percentage range of 50 to 5% caprolactone, among other possibilities. When present, the poly(lactide-co-glycolide) can have, for example, a mole percentage range of 50 to 99.9% lactide and a mole percentage range of 50 to 0.1% glycolide, among other possibilities. In certain specific embodiments, the additional conformal coating can include 50 to 99.9 wt % (e.g., 50 to 60 to 70 to 80 to 90 to 95 to 99 to 99.5 to 99.9 wt %) of one or more biodegradable polymers and 0.1 to 50 wt % (e.g., 0.1 to 0.5 to 1 to 5 to 10 to 20 to 30 to 40 to 50 wt %) mometasone furoate, among many other possibilities. Typical amounts of mometasone furoate can range, for example, from 0.1 μg / mm 2 or less to 20 μg / mm2 or more (i.e., range of 0.1 μg or less to 20 μg or more mometasone furoate / mm 2 The surface area of ​​the scaffold, wherein the surface area A of the scaffold is calculated as A=πDL, wherein D is the preparation diameter of the scaffold, and L is the preparation length of the scaffold), for example, in the range of 0.1 μg / mm 2 To 0.2 μg / mm 2 To 0.5 μg / mm 2 To 1 μg / mm 2 Up to 2 μg / mm 2 Up to 5 μg / mm 2 Up to 10 μg / mm 2 Up to 15 μg / mm 2 Up to 20 μg / mm 2 (i.e., the range is between any two of the aforementioned values), more typically the range is 1 μg / mm 2 Up to 10 μg / mm 2 , and other possible values.

[0040] In certain embodiments that can be used with any of the above aspects and embodiments, the stent can be further coated with a conformal surface layer that is disposed over an additional conformal coating that includes additional coating material and a therapeutic agent. The surface layer can be formed, for example, from a single biodegradable polymer or a mixture of biodegradable polymers selected from those described elsewhere herein. In certain embodiments, the surface layer can be formed from the same polymer that is present in the underlying additional conformal coating, but will not include the therapeutic agent. The surface layer can be used, for example, to delay and / or slowly release the therapeutic agent in the underlying additional conformal coating. The thickness of the surface layer can range, for example, between 1 μm and 30 μm, among other possibilities.

[0041] In other aspects, the invention relates to a method of treatment that includes (a) introducing a stent, such as a stent according to any of the above aspects and embodiments, into a patient's sinus cavity while in a radially constrained shape, and (b) removing the constraint that holds the stent in the constrained shape so that the stent self-expands within the sinus cavity. Examples of sinus cavities suitable for device implantation include the ethmoid sinuses, the middle meatal space, the frontal sinus ostia (also known as the frontal recess), the maxillary sinus ostia, and the sphenoid sinus ostia, among others.

[0042] In still other aspects, the invention relates to a kit that includes (a) a stent, such as a stent according to any of the above aspects and embodiments, (b) a delivery catheter, and (c) an optional loading aid. In certain embodiments, a stent according to any of the above aspects and embodiments can be loaded into a 15 French delivery catheter or smaller, a 9 French delivery catheter or smaller, a 6 French delivery catheter or smaller, or even a 4 French delivery catheter or smaller. In certain embodiments, a stent according to any of the above aspects and embodiments can be loaded into a 6.5 French to 9 French catheter.

[0043] In certain embodiments, the delivery catheter can be configured to hold the stent in a radially constrained shape and to remove the constraint that holds the stent in the radially constrained shape at the delivery location.

[0044] In certain embodiments that can be used with any of the above aspects and embodiments, the delivery catheter can include an expandable device. For example, the delivery catheter can be a balloon catheter that includes a catheter shaft having an inflatable lumen and one or more inflatable balloons disposed at or near the distal end of the catheter shaft, where the one or more inflatable balloons may or may not be at least partially coated with a coating containing a therapeutic agent.

[0045] In still other aspects, the present invention relates to a delivery system comprising (a) a stent, such as a stent according to any of the above aspects and embodiments, and (b) a delivery catheter, wherein the stent is located within the delivery catheter in a radially constrained shape. Such a delivery system can be used, for example, in a treatment method that includes (a) introducing the stent into a patient's sinus cavity while in a radially constrained shape such that the stent is in a delivery position within the sinus cavity, and (b) removing the constraint that keeps the stent in the constrained shape so that the stent self-expands within the sinus cavity.

[0046] In still other aspects, the present invention relates to a delivery system comprising (a) a stent, such as a stent according to any of the above aspects and embodiments, and (b) a delivery catheter comprising an expandable device, wherein the stent is positioned above, within, below, proximal, or distal to the expandable device. For example, the expandable device can be an expandable frame or an inflatable balloon. For example, the delivery catheter can be a balloon catheter that includes a catheter shaft having an inflatable lumen and one or more inflatable balloons disposed at or near the distal end of the catheter shaft, wherein the one or more inflatable balloons may or may not be at least partially coated with a coating containing a therapeutic agent. Such a delivery system can be used, for example, in a treatment method that includes (a) introducing the stent into a patient's sinus cavity while in a radially constrained shape such that the stent is in a delivery position within the sinus cavity, and (b) expanding the expandable device while the expandable device is within the stent lumen.

[0047] In still other aspects, the present invention relates to a method of forming a coated stent, comprising: (a) applying a first coating solution comprising a first solvent, a branched biodegradable polymer, and a diisocyanate crosslinking agent to a stent according to any of the above aspects and embodiments, and (b) curing the applied first coating solution at room temperature or at an elevated temperature.

[0048] In certain embodiments that can be used in conjunction with any of the above aspects and embodiments, the branched biodegradable polymer can be, for example, branched poly(lactide-co-s-caprolactone), such as branched hydroxy-terminated poly(lactide-co-ε-caprolactone).

[0049] In certain embodiments that can be used in conjunction with any of the above aspects and embodiments, the solution can further comprise a chain terminator. For example, the chain terminator can be an alcohol, such as a C8-C18 alcohol, such as 1-dodecanol and stearyl alcohol, among many other possibilities.

[0050] In certain embodiments that can be used in conjunction with any of the above aspects and embodiments, the diisocyanate crosslinking agent can be hexamethylene diisocyanate, among many other possibilities.

[0051] In certain embodiments that can be used with any of the above aspects and embodiments, the first solvent can include dichloromethane or ethyl acetate, among many other possibilities. In certain of these embodiments, the first solvent can further include anisole, as a co-solvent.

[0052] In certain embodiments that can be used with any of the above aspects and embodiments, the stent can have a braided structure that includes one or more stent material strands and a plurality of nodes, and the coating solution can be applied to at least the nodes of the stent.

[0053] In certain embodiments that can be used with any of the above aspects and embodiments, the method further includes, after applying the first coating solution, applying a second coating solution to the stent, the second coating solution comprising an additional biodegradable polymer (e.g., poly(lactide-co-ε-caprolactone), among many other possibilities), a second solvent (e.g., comprising ethyl formate and anisole, among many other possibilities), and a therapeutic agent. The therapeutic agent can be a steroidal anti-inflammatory drug, such as mometasone furoate, among many other possibilities.

[0054] Other aspects of the invention relate to a stent having a coating formed by the method according to any of the above aspects and embodiments.

[0055] Potential benefits of the invention include one or more of the following aspects related to adult and pediatric surgery, among others: (a) stabilizing the sinus ostium / orifice, (b) reducing adhesions and postoperative adhesions, (c) local and extended release therapeutic agent delivery for use as an alternative therapy to surgery (e.g., treating patients who have failed medical control with oral and / or topical steroids), pre- and / or postoperative care, and (d) delivering a therapeutic agent to refractory patients who are insensitive to FESS, (e) preventing stenosis of the sinus ostium / orifice after surgical dilation.

[0056] These and other aspects, embodiments, and benefits of the invention will become immediately apparent to those of ordinary skill in the art when reviewing the detailed description and the subsequent claims.

[0057] Additional aspects of the invention are set forth in the following paragraphs:

[0058] Aspect 1. A stent configured for implantation into a sinus cavity, the stent comprising a generally tubular structure having a lumen and comprising a stent material and an optional conformal coating, the conformal coating comprising a coating material that at least partially coats the stent material.

[0059] The stent of aspect 1, wherein the stent comprises a fibrous infrastructure.

[0060] Aspect 3. The stent of Aspect 1, wherein the stent comprises a braided structure comprising one or more chains of stent material.

[0061] Aspect 4. The stent of Aspect 3, wherein the braided structure comprises opposing sets of helical chains.

[0062] Aspect 5. The stent of Aspect 4, wherein each set of helical chains comprises from 2 to 64 members.

[0063] Aspect 6. The stent of Aspect 3, wherein the braided structure comprises a first chain of material having a first stiffness and a second chain of material having a second stiffness greater than the first stiffness.

[0064] Aspect 7. The stent of Aspect 6, wherein the second chain of material has a modulus > 5 GPa and is at least 2 times that of the first chain of material.

[0065] Aspect 8. The stent of any one of Aspects 3 - 8, wherein the braided structure comprises units of different sizes.

[0066] Aspect 9. The stent of Aspect 8, wherein a portion of the braided structure is removed so as to form units of different sizes, or wherein the braided structure is woven so as to form units of different sizes.

[0067] Aspect 10. The stent of Aspect 8, comprising a first unit having a first region and a second unit having a second region, wherein the first region is at least 50% greater than the second region.

[0068] Aspect 11. The stent of Aspect 8, wherein the variation in unit size occurs along the longitudinal length of the stent.

[0069] Aspect 12. The stent of Aspect 8, wherein the variation in unit size occurs around the circumference of the stent.

[0070] Aspect 13. The stent of any one of Aspects 3 - 12, further comprising longitudinal elastic fibers mechanically coupled to two or more nodes of the braided structure.

[0071] Aspect 14. The stent of any one of Aspects 3 - 13, wherein one or more ends of the one or more chains are woven back into and connected to the braided structure.

[0072] Aspect 15. The stent of any one of Aspects 3 - 14, wherein the stent comprises the conformal coating, which comprises a coating material.

[0073] Aspect 16. The stent of Aspect 15, wherein the coating material comprises an elastomer.

[0074] Aspect 17. The stent of Aspect 16, wherein the elastomer comprises a urethane crosslinker.

[0075] Aspect 18. A stent according to any one of aspects 15 - 17, wherein the coating material covers some of the nodes of the braided structure while leaving other nodes uncovered.

[0076] Aspect 19. A stent according to any one of aspects 15 - 18, wherein the coating material covers alternating regions along the length of the braided structure.

[0077] Aspect 20. A stent according to any one of aspects 15 - 18, wherein the coating material covers the ends of the braided structure without covering the regions between the ends of the braided structure.

[0078] Aspect 21. A stent according to any one of aspects 15 - 21, wherein the thickness of the coating material at the nodes of the braided structure ranges from 1 to 100 times the thickness of the coating material between the nodes of the braided structure.

[0079] Aspect 22. A stent according to any one of aspects 15 - 21, wherein one or more chains of the stent material comprise poly(lactide - co - glycolide), and wherein the coating material is an elastomeric material comprising poly(L - lactide - co - caprolactone) crosslinked with a urethane crosslinker, a urethane crosslinker, or both a urethane and a urea crosslinker.

[0080] Aspect 23. A stent according to any one of aspects 15 - 21, wherein one or more chains of the stent material comprise poly(lactide - co - glycolide), and wherein the coating material is an elastomeric material comprising diisocyanate - cured, hydroxyl - terminated branched poly(L - lactide - co - caprolactone).

[0081] Aspect 24. A stent according to aspect 24, wherein the hydroxyl - terminated branched poly(L - lactide - co - caprolactone) is cured with hexamethylene diisocyanate.

[0082] Aspect 25. A stent according to any one of aspects 15 - 21, 23, and 24, wherein the stent is further coated with an additional coating material comprising 50 to 99.9 wt% poly(L - lactide - co - caprolactone) and 0.1 to 50 wt% mometasone furoate.

[0083] Aspect 26. A stent according to aspect 22, wherein the stent is further coated with an additional coating material comprising 50 to 99.9 wt% poly(L - lactide - co - caprolactone) and 0.1 to 50 wt% mometasone furoate.

[0084] Aspect 27. A stent according to aspect 3, wherein the braided structure is a strip - shaped elongate element wound into a helical tubular structure.

[0085] Aspect 28. The stent of Aspect 1, wherein the stent comprises an elongate element wound into a helical tubular structure.

[0086] Aspect 29. The stent of Aspect 1, wherein the stent comprises a plurality of parallel open rings.

[0087] Aspect 30. The stent of Aspect 29, wherein the open rings are band-shaped open rings.

[0088] Aspect 31. The stent of Aspect 30, wherein the band-shaped open rings have a plurality of holes.

[0089] Aspect 32. The stent of Aspect 31, wherein the plurality of holes produce a braided-like structure.

[0090] Aspect 33. The stent of Aspect 1, wherein the generally tubular structure is a knitted structure.

[0091] Aspect 34. The stent of Aspect 33, wherein the knitted structure comprises a single strand that can be pulled to untie and remove the stent.

[0092] Aspect 35. The stent of Aspect 1, comprising a plurality of radially expandable inserts within the generally tubular structure.

[0093] Aspect 36. The stent of Aspect 35, wherein the radially expandable inserts comprise a center and a plurality of radially expandable arms, or wherein the radially expandable inserts comprise a braided ring.

[0094] Aspect 37. The stent of Aspect 1, wherein the distal end of the stent is configured to be captured by another device, and wherein the stent is configured to be inverted and removed by pulling the distal end into the cavity.

[0095] Aspect 38. A treatment method, comprising (a) introducing a stent according to any one of Aspects 1 - 37 into a patient's sinus cavity while in a radially restricted shape, and (b) removing the restriction that maintains the stent in the restricted shape so that the stent self-expands within the sinus cavity.

[0096] Aspect 39. The method of Aspect 38, wherein the sinus cavity is the ethmoid sinus, the middle meatal space, the frontal sinus ostium, the maxillary sinus ostium, the sphenoid sinus ostium, or the frontal sinus recess.

[0097] Aspect 40. A kit, comprising (a) a stent according to any one of Aspects 1 - 37, (b) a delivery catheter, and (c) an optional loading aid.

[0098] Aspect 41. The kit of Aspect 40, wherein the delivery catheter is configured to maintain the stent in a radially restricted shape and to remove the restriction that maintains the stent in the radially restricted shape at the delivery position.

[0099] Aspect 42. A kit of aspect 40, wherein the delivery catheter comprises an expandable device.

[0100] Aspect 43. A kit of aspect 40, wherein the delivery catheter is a balloon catheter, which comprises a catheter shaft having an inflatable lumen and one or more inflatable balloons disposed at or near the distal end of the catheter shaft.

[0101] Aspect 44. A kit of aspect 43, wherein at least one of the one or more inflatable balloons is at least partially coated with a therapeutic agent-containing coating.

[0102] Aspect 45. A delivery system, comprising (a) a stent according to any one of aspects 1-37, and (b) a delivery catheter, wherein the stent is in a radially constrained shape within the delivery catheter.

[0103] Aspect 46. A treatment method using the delivery system of aspect 45, comprising: (a) introducing the stent into the sinus cavity of a patient while in a radially constrained shape such that the stent is in a delivery position within the sinus cavity, and (b) removing the constraint that keeps the stent in the constrained shape so that the stent self-expands within the sinus cavity.

[0104] Aspect 47. A delivery system, comprising (a) a stent according to any one of aspects 1-37, and (b) a delivery catheter comprising an expandable device, wherein the stent is located above, within, below, proximal or distal to the expandable device.

[0105] Aspect 48. The delivery system of aspect 47, wherein the expandable device is an inflatable balloon or an expandable frame.

[0106] Aspect 49. The delivery system of aspect 47, wherein the delivery catheter is a balloon catheter, which comprises a catheter shaft having an inflatable lumen and one or more inflatable balloons disposed at or near the distal end of the catheter shaft.

[0107] Aspect 50. The delivery system of aspect 49, wherein at least one of the one or more inflatable balloons is at least partially coated with a therapeutic agent-containing coating.

[0108] Aspect 51. A treatment method using the delivery system of aspect 47, comprising: (a) introducing the stent into the sinus cavity of a patient such that the stent is in a delivery position within the sinus cavity, and (b) inflating the expandable device while the expandable device is within the stent lumen.

[0109] Aspect 52. The method of aspect 51, wherein the expandable device is a balloon.

[0110] Aspect 53. A method of forming a coated stent, comprising: (a) applying a first coating solution comprising a first solvent, a branched biodegradable polymer, and a diisocyanate crosslinking agent to the stent, and (b) curing the applied first coating solution at an elevated temperature, wherein the stent is configured for implantation into a sinus cavity and the stent has a generally tubular structure having a lumen and comprising stent material.

[0111] Aspect 54. The method of aspect 53, wherein the branched biodegradable polymer is a branched hydroxy-terminated poly(lactide-co-caprolactone).

[0112] Aspect 55. The method of any one of aspects 53-54, wherein the stent material comprises poly(lactide-co-glycolide).

[0113] Aspect 56. The method of any one of aspects 53-54, wherein the first solution further comprises a chain terminator.

[0114] Aspect 57. The method of aspect 56, wherein the diisocyanate crosslinking agent is hexamethylene diisocyanate, wherein the chain terminator is 1-dodecanol, or a combination of both.

[0115] Aspect 58. The method of any one of aspects 53-57, wherein the first solvent comprises dichloromethane.

[0116] Aspect 59. The method of aspect 58, wherein the first solvent further comprises anisole.

[0117] Aspect 60. The method of aspect 59, wherein the stent is a braided structure comprising one or more strands of stent material and a plurality of nodes, and the coating solution is applied to at least the nodes of the stent.

[0118] Aspect 61. The method of any one of aspects 53-57, wherein the method further comprises applying a second coating solution comprising a second solvent, an additional biodegradable polymer, and a therapeutic agent to the stent after curing.

[0119] Aspect 62. The method of aspect 61, wherein the additional biodegradable polymer is poly(lactide-co-caprolactone).

[0120] Aspect 63. The method of any one of aspects 61-62, wherein the therapeutic agent is a steroidal anti-inflammatory drug.

[0121] Aspect 64. The method of any one of aspects 61-62, wherein the therapeutic agent is mometasone furoate.

[0122] Aspect 65. The method of aspect 64, wherein the second solvent comprises ethyl formate and anisole.

[0123] Aspect 66. The method of any one of aspects 61 - 65, wherein the first coating solution and the second coating solution are applied by spraying method.

[0124] Aspect 67. A stent formed by the method of any one of aspects 53 - 66.

[0125] Aspect 68. The stent of any one of aspects 1 - 37, wherein the stent material comprises a therapeutic agent.

[0126] Aspect 69. The stent of aspect 68, wherein the therapeutic agent is a steroidal anti - inflammatory drug.

[0127] Aspect 70. The stent of any one of aspects 15 - 26, wherein the coating material comprises a therapeutic agent.

[0128] Aspect 71. The stent of aspect 70, wherein the therapeutic agent is a steroidal anti - inflammatory drug.

[0129] Aspect 72. The stent of any one of aspects 15 - 26, further comprising an additional conformal coating, wherein the conformal coating comprises an additional coating material and a therapeutic agent.

[0130] Aspect 73. The stent of aspect 72, wherein the therapeutic agent is a steroidal anti - inflammatory drug.

[0131] Additional aspects and embodiments of the present invention are discussed in the detailed description set forth below. Brief Description of the Drawings

[0132] For example, non - limiting embodiments of the present invention are described with reference to the accompanying drawings, which are schematic and not necessarily drawn to scale. In the drawings, each identical or nearly identical element that is illustrated is generally represented by a single number. For clarity purposes, not every element is labeled in each figure, nor is every element of each embodiment of the present invention shown, where the examples are not necessary for a person of ordinary skill in the art to understand the present invention. In the drawings:

[0133] Figure 1A Schematic examples illustrate cross - sections of a plurality of fibers according to an embodiment of the present invention.

[0134] Figure 1B Schematic examples illustrate cross - sections of multifilament fibrils according to two embodiments of the present invention.

[0135] Figure 2 Is a schematic side view of a self - expanding stent according to an embodiment of the present invention.

[0136] Figure 3A Is a schematic side view of a self - expanding stent with a uniform braiding angle according to an embodiment of the present invention.

[0137] Figure 3B Schematic side view of a self-expanding stent with variable braiding angle according to an embodiment of the present invention.

[0138] Figure 4 Schematic side view of a self-expanding stent with an elastomeric coating according to an embodiment of the present invention.

[0139] Figure 5 Schematic side view of a self-expanding stent with elastic cross fibers according to an embodiment of the present invention.

[0140] Figure 6A Schematic side view of a self-expanding stent with filaments of different stiffnesses according to an embodiment of the present invention.

[0141] Figure 6B Schematic side view of a self-expanding stent with a removed filament section according to an embodiment of the present invention.

[0142] Figure 6C Schematic side view of a self-expanding stent with coated ends according to an embodiment of the present invention.

[0143] Figure 6D Schematic side view of a self-expanding stent with alternating coated and uncoated portions according to an embodiment of the present invention.

[0144] Figure 7 Photograph of a self-expanding stent with unequal cell sizes according to an embodiment of the present invention.

[0145] Figure 8 Schematic side view of a self-expanding stent with fold-back ends according to an embodiment of the present invention.

[0146] Figure 9 Example drawing of a knitted stent according to an embodiment of the present invention.

[0147] Figure 10 Schematic perspective view of a helical self-expanding stent according to an embodiment of the present invention.

[0148] Figure 11A Photograph of a helical self-expanding stent formed from a braided tubular stent according to an embodiment of the present invention.

[0149] Figure 11B Photograph of a helical self-expanding stent formed from two carrier braids according to an embodiment of the present invention.

[0150] Figure 12A Schematic perspective view of a self-expanding stent with strong strut rings according to an embodiment of the present invention.

[0151] Figure 12B Schematic perspective view of a self-expanding stent with strut rings in the form of two carrier weaves, according to an embodiment of the present invention.

[0152] Figure 13 Schematic side view of a conformable tube according to an embodiment of the present invention.

[0153] Figure 14A Schematic perspective view of a conformable tube with a combined three-dimensional support structure in an expanded form, according to an embodiment of the present invention.

[0154] Figure 14B Schematic end view of a conformable tube with a combined three-dimensional support structure in a crimped form, according to an embodiment of the present invention.

[0155] Figure 15 Schematic side view of a stent in the form of a single polymeric structure, according to an embodiment of the present invention.

[0156] Figure 16 Photographs of 8 mm diameter stents, 10 mm diameter stents, 20 mm diameter stents and 31 mm diameter stents, each having 16 links, according to an embodiment of the present invention.

[0157] Figure 17A Graph showing the cumulative absolute mass of released mometasone furoate (MF) as a function of time for three different drug loadings, in the presence of poly(lactic-co-caprolactone) (PLCL) as a drug carrier polymer, illustrative of an embodiment of the present invention.

[0158] Figure 17B Illustrative of an embodiment for Figure 17A Graph showing the cumulative percentage mass of released MF as a function of time, in the presence of PLCL as a drug carrier polymer.

[0159] Figure 18 Graph showing the cumulative percentage mass of released MF as a function of time, in the presence of PLCL as a drug carrier polymer, for a 400 μg MF stent without a surface layer and three 400 μg MF stents with different surface layer thicknesses, with and without a surface layer comprising PLCL and PLA, illustrative of an embodiment of the present invention.

[0160] Figure 19 Graph showing the cumulative percentage mass of released MF as a function of time, in the presence of D,L-PLGA as a drug carrier polymer, for 400 μg MF stents comprising three different types of D,L-PLGA, illustrative of an embodiment of the present invention.

[0161] Figure 20A Photograph of a 31.75 mm stent with 16 struts according to an embodiment of the present invention.

[0162] Figure 20B is a photograph of the coated nodes of a stent like Figure 20A the stent.

[0163] Figure 21 Graph of the compressive load of a stent versus compressive strain to illustrate an embodiment of the present invention.

[0164] Figure 22A - 22E Photographs to illustrate various stent designs according to embodiments of the present invention.

[0165] Figure 23A , Figure 23B , Figure 23C and Figure 23D Photograph to illustrate the deployment of a stent in a porcine nasal cavity according to an embodiment of the present invention.

[0166] Figure 24 Photograph to illustrate a stent after deployment in a porcine nasal cavity according to an embodiment of the present invention.

[0167] Figure 25 Photograph to illustrate a 32-strut stent with a 13 mm diameter and 10 mm length after deployment in the native middle meatus of a human cadaver according to an embodiment of the present invention.

[0168] Figure 26 Photograph to illustrate a 16-strut 10 mm stent after deployment in the frontal sinus ostium of a human cadaver according to an embodiment of the present invention.

[0169] Figure 27 Photograph to illustrate a 32-strut stent with a 17.5 mm diameter and 10 mm length after deployment in the ethmoid sinus of a human cadaver after FESS according to an embodiment of the present invention.

[0170] Figure 28A - 28D Optical micrographs of a coated 8 mm stent with 16 struts with and without anisole as a co-solvent during spray coating, as follows: Figure 28A , PLGA (10:90) stent without anisole co-solvent; Figure 28B , PLGA (10:90) stent with anisole co-solvent; Figure 28C , PLGA (75:25) stent without anisole co-solvent; Figure 28D , PLGA (75:25) with anisole co-solvent.

[0171] Figure 29A - 29C Shown are optical images of scaffolds coated with and without anisole as a co-solvent during spray coating, as follows: Figure 29A , a scaffold coated with 62 wt% elastomer from a solution without anisole as a co-solvent relative to the base fabric; Figure 29B , a scaffold coated with 63 wt% elastomer from a solution containing anisole as a co-solvent; and Figure 29C , a scaffold coated with 100 wt% elastomer from a solution containing anisole as a co-solvent.

[0172] Figure 30A Examples illustrate the cumulative absolute mass of MF released from three groups of MF-coated scaffolds as a function of time.

[0173] Figure 30B Examples illustrate the cumulative percentage mass of MF released from three groups of MF-coated scaffolds as a function of time.

[0174] Figure 31 Examples illustrate the in vivo drug release profiles of MF-coated PLGA (10:90) scaffolds and MF-coated PLGA (75:25) scaffolds.

[0175] Figure 32 Examples illustrate the concentration of MF in the sinus mucosa of rabbits sacrificed after implantation as a function of time.

[0176] Figure 33 Examples illustrate total MF in vivo as a function of time (MF on the scaffold + MF in the sinus mucosa of sacrificed rabbits).

[0177] Figure 34 Examples illustrate the cumulative percentage mass of MF released from two groups of MF-coated scaffolds as a function of time.

[0178] Figure 35 Examples illustrate the cumulative percentage mass of MF released from four groups of MF-coated scaffolds as a function of time.

[0179] Figure 36A Examples illustrate the immediate recovery of the first compression amount from two groups of MF-coated scaffolds with 90 and 128 weave angles as a function of compression time.

[0180] Figure 36B Examples illustrate the immediate recovery of the second compression amount from two groups of MF-coated scaffolds with 90 and 128 weave angles as a function of compression time.

[0181] Figure 37AThe example illustrates six-hour recovery from a first compression for two groups of MF-coated stents having 90 and 128 braid angles as a function of compression time.

[0182] Figure 37B The example illustrates six hour recovery from a second compression of two sets of MF-coated stents having 90 and 128 braid angles as a function of compression time.

[0183] Figure 38 FIG. 4 is a schematic diagram of a test apparatus for performing a compression test according to an embodiment of the present invention. Detailed description

[0184] The implantable medical device of the present invention is a generally tubular device, and in multiple embodiments, the device is a self-expanding device. As used herein, "device", "scaffold", "stent", "carrier" and "implant" can be used synonymously. Moreover, "self-expanding" as used herein means a device that includes curling to reduce the delivery structure for delivery to the body, and then once released from the delivery structure, tends to expand into a larger suitable structure without the assistance of any additional expansion device or with the partial assistance of balloon assistance or similar auxiliary expansion. As used herein, "strength" and "rigidity" can be used synonymously, referring to the resistance of the medical stent of the present invention to the radial force deformation applied by the stent relative to the static adjacent material. Examples of strength and stiffness measurements such as for characterizing the medical stent of the present invention include radial resistance and chronic outward force, as further described herein.

[0185] The stent according to the present invention is a generally tubular stent, which can be a variety of sizes, including a variety of diameters and lengths, and it can be used for a variety of sinus applications. In the case of non-circular cross-section objects, "diameter" refers to width. In certain useful embodiments, the diameter range of the stent in the prepared state (or unrestricted) can be 5mm or less to 60mm or more, for example, in the range of 5mm to 10mm to 15mm to 20mm to 25mm to 30mm to 35mm to 40mm or 50mm to 60mm (that is, in the range of any two of the aforementioned values), usually in the range of 5 to 13mm or 15 to 30mm. In certain useful embodiments, the length range in the prepared state (or unrestricted) can be 5mm or less to 30mm or more, for example, in the range of 5mm to 10mm to 15mm to 20mm to 25mm or 30mm (that is, in the range of any two of the aforementioned values), usually in the range of 10mm to 20mm.

[0186] In certain beneficial embodiments, the stent mass may range from 1 to 20 mg / mm length.

[0187] Unless otherwise indicated, the stent diameters and stent lengths given herein refer to the unrestrained (as fabricated) diameters and lengths.

[0188] Many stent embodiments of the invention are self-expanding, where they are fabricated to a first diameter and then reduced or "crimped" to a second, reduced diameter for placement within a delivery catheter and self-expand to near the first diameter when extruded from the delivery catheter at the implantation site. In certain embodiments, the first diameter can be at least 10% larger than the diameter of the body cavity into which it is implanted. In certain embodiments, the stent can be designed to recover at least about 70%, at least about 80%, at least about 90%, up to about 100% of its fabricated first diameter.

[0189] The stents according to the invention provide inflation and mechanical properties suitable for enabling the stents to be effectively used for their intended purposes in the sinus cavities. Two measures of such mechanical properties used herein are "radial resistance force" ("RRF") and "chronic outward force" ("COF"). The RRF is the force exerted by the stent in reaction to the crimping force, and the COF is the force exerted by the stent relative to a static abutting surface. In certain embodiments, the stent is configured to have a relatively high RRF, capable of keeping open body cavities, lumens, and nasal features, etc., but having a relatively low COF so as to avoid exerting potentially damaging forces on the body cavity wall, optic nerve, brain, etc. For example, after being crimped, the stent of the invention preferably inflates to 70 to 100% of its as-fabricated state structure, has an RRF in the range of 50 to 300 mmHg, and / or has an acute COF (when delivered into the sinus cavity) in the range of 10 to 100 mmHg.

[0190] The stent according to the present invention can be formed from a variety of polymeric and non-polymeric materials. The stent according to the present invention can be biodegradable or non-biodegradable, or a combination of both biodegradable and non-biodegradable materials. When biodegradable, after being placed within the patient's sinus cavity, the stent can be completely absorbed, for example, in as short a time as three weeks or less up to as long as 52 weeks or more. In certain embodiments, the generally tubular structure can be completely absorbed at some time after 12 weeks of placement and before 32 weeks of placement. In other embodiments, instead of being absorbed into the nasal mucosa, the biodegradable device can also be removed by nasal irrigation. The device can also be designed such that resorption of non-continuous portions results in decomposition into predetermined small pieces (the longest dimension typically <10 mm, or more typically <5 mm), which can be cleared from the sinuses and nasal cavity via normal mucociliary activity, resulting in swallowing or expulsion from the nose. Thus, the amount of acidic resorption by-products (e.g., lactic acid, glycolic acid) in contact with the sinus or nasal surface can be reduced. This can reduce irritation or inflammation of these tissues and surrounding tissues. In certain embodiments, additives having alkaline properties can also be added to the device to neutralize the acidic by-products, which can reduce the inflammatory response associated with the acidic by-products. Moreover, in certain embodiments, a variety of substances having different rates of bioresorption can also be mixed to reduce the amount of degraded material at any given time and the resulting biological response.

[0191] In various embodiments, the implantable stent can include a generally tubular structure comprising a stent material. The stent according to the present invention can be fiber-based or non-fiber-based.

[0192] In multiple embodiments, the scaffold material can be a biodegradable scaffold material, typically a biodegradable scaffold material comprising one or more biodegradable polymers. Non-limiting examples of biodegradable polymers for forming the biodegradable scaffold material include biodegradable polyesters, polycarbonates, polyhydroxyalkanoates, polyanhydrides, and polyorthoesters, non-limiting examples of which include homopolymers of lactic acid (PLA), homopolymers of glycolic acid (PGA), homopolymers of trimethylene carbonate (PTMC), homopolymers of caprolactone (PCL), homopolymers of polypropylene fumarate, and homopolymers of dioxanone (PDO), and copolymers comprising two or more of the foregoing monomers, such as poly(lactic-co-glycolic acid) (PLGA), poly(lactic-co-caprolactone) (PLCL), and poly(glycolic-co-caprolactone) (PGCL). Preferred copolymers include PLGA having a lactic acid molar percentage range of 10 to 90% and a glycolic acid molar percentage range of 90 to 10%, more typically a lactic acid range of 10 to 75% and a glycolic acid molar percentage range of 90 to 25%; for example, in certain embodiments, PLGA 75:25 (mol / mol) or PLGA (10:90) (mol / mol) can be applied. The components of the PLGA polymer within these ranges can be optimized to meet the mechanical properties and degradation requirements of the specific application for which the scaffold is used. In certain embodiments, the biodegradable scaffold material can include prodrug-based polymers, such as polyaspirin, which can be present as a single component or subcomponent of a generally tubular structure for preparing a scaffold capable of controlled release of a therapeutic agent upon degradation.

[0193] In multiple embodiments, the scaffold material can be a non-biodegradable scaffold material, typically a non-biodegradable scaffold material comprising one or more non-biodegradable polymers. Non-limiting examples of non-biodegradable polymers for forming the non-biodegradable scaffold material include polyolefins, such as polyethylene (HDPE and LDPE) and polypropylene, halogenated polyolefins such as polyvinyl chloride (PVC), and fluoropolymers including polytetrafluoroethylene (PTFE) and perfluoroalkoxy alkane (PFA), polyarenes such as polystyrene, polyesters such as polyethylene terephthalate (PET), polyamides such as nylon, silicones, mucoadhesive materials, and biostable polyurethanes (PU).

[0194] The scaffold according to the present invention can optionally include a coating formed from a coating material, which at least partially coats the scaffold material.

[0195] Coatings can be applied for a variety of purposes, including enhancing mechanical properties, controlling degradation, and releasing and controlling therapeutic agents. The coating can cover all or part of the stent, or in a fiber-based technology, all or part of the filaments or strands that form the stent. As used herein, "strand" and "filament" can be used interchangeably and include single-fiber strands and filaments (also referred to as monofilaments) and multi-fiber strands and filaments.

[0196] If the stent to be coated is a fiber-based structure, the coating can be applied to the individual strands before the formation of the stent, or to the stent after the formation of the stent. If the stent is a non-fiber-based structure, the coating can be applied to, for example, a rigid polymer tube or sheet, before or after using a suitable cutting technique such as mechanical cutting or thermal cutting to remove material. Any suitable method can be used to produce the coating, including spraying, electrospraying, roll coating, dipping, chemical vapor deposition, electrospinning, and / or co-extrusion, among others. In certain embodiments, the coating can include additional reagents, such as therapeutic agents, as described in further detail below.

[0197] In various embodiments, the coating material can be a biodegradable or non-biodegradable coating material or a combination of both, typically a biodegradable coating material comprising one or more biodegradable polymers or a non-biodegradable coating material comprising one or more non-biodegradable polymers. Non-limiting examples of biodegradable polymers for forming biodegradable coating materials include the biodegradable polymers listed above. Non-biodegradable polymers for forming non-biodegradable coating materials include the non-biodegradable polymers listed above.

[0198] In multiple embodiments, a coating including an elastomer is formed. Potential benefits of such coatings include enhanced mechanical properties. For example, the coating can be made of an elastic polymer which, due to its elastic properties, when compressed or stretched, exerts a force on the stent that facilitates radial expansion, thereby enhancing recoverability and / or radial stiffness, as well as other properties. Potential benefits of additional elastomers may be to encapsulate the stent material (which can be a braided structure, etc.), maintain integrity and provide a smooth, soft surface that minimizes tissue irritation at the contact points while providing good conformability. In this regard, certain aspects of the designs described herein, including those resulting from combinations of composite structures and bioabsorbable filaments and elastic coatings, provide properties not achievable with other bioabsorbable stent designs. Potential benefits include higher radial resistance and / or chronic outward force with a lower amount of polymer, reduced profile (thickness of the stent wall) and / or better conformability, which is caused by the spring-like structure at each fiber intersection, enabling delivery to the target location via a smaller delivery system or guiding catheter, and / or providing good apposition and conformability to the target location with a smaller manufactured state stent diameter. Better conformability can result in more effective drug delivery to the tissue based on improved tissue contact. Also, better conformability can facilitate the surgeon in maneuvering the deployed implant to the desired location. For example, when repositioning one side of the implant, the opposite side of the implant has a tendency to reposition itself unless it is well-contoured and adheres to the tissue.

[0199] The coating thickness of the elastomeric coating can vary to a large extent, with typical coating thickness ranges being, for example, 5 to 50 μm, among other thicknesses. When a braided stent is coated, the elastomeric coating can range, for example, between 30 and 150% of the weight of the braided stent substrate.

[0200] Elastomers include thermoset and thermoplastic elastomers. Thermoset or thermoplastic elastomers advantageously have a glass transition temperature (Tg) below room temperature (25 °C), and are more beneficial when below 10 °C. Thermoset elastomers can provide high elongation at break and low permanent deformation under cyclic mechanical testing. Examples of elastomers include, for example, poly(glycolide-co-ε-caprolactone) (PGCL) or poly(lactide-co-ε-caprolactone) (PLCL), including poly(L-lactide-co-ε-caprolactone) and poly(D,L-lactide-co-ε-caprolactone). In certain embodiments, PLCL can have a lactide molar percentage range of 20 to 80% and a caprolactone molar percentage range of 80 to 20%, more typically a lactide molar percentage range of 30 to 50% and a caprolactone molar percentage range of 50 to 70%.

[0201] In certain embodiments, the biodegradable coating material is a thermoset elastomer formed from a polymeric polyol, the polymeric polyol including diols, triols, tetraols, and / or higher alcohols. Such polymers can be crosslinked with a crosslinker, which is a bifunctional or polyfunctional small molecule or polymer. For example, crosslinking can be formed by reacting such a polymer with a bifunctional or polyfunctional isocyanate, which can be in the form of a small molecule or polymer.

[0202] In cases where the coating includes a thermoset elastomer polymer, the crosslink density can vary to obtain desired mechanical properties. For example, an optional chain terminator can be used in a thermoset elastomeric material such as a polyester urethane to control the crosslink density. The chemical crosslink density is adjusted using a chain terminator to control the degree of crosslinking generated during the curing of the polyester-urethane. The crosslink density of the resulting elastomer depends on the concentration of the chain terminator incorporated into the elastic network. Examples of suitable chain terminators include any suitable monofunctional compound, such as monofunctional isocyanates, alcohols, amines, acyl chlorides, and sulfonyl chlorides.

[0203] In certain embodiments, the thermoset elastomer includes a polyester polyol, a diisocyanate crosslinker, and an optional chain terminator. Such a thermoset elastomer can be prepared by a method including the steps of: at least partially dissolving the polyester polyol in a solvent to form a solution; adding the diisocyanate crosslinker to the solution; optionally adding the chain terminator to the solution; coating the solution on a scaffold material; and curing the solution. When using a solvent-based method, a weakly volatile co-solvent can be used during the coating process to improve the nodule accumulation of the thermoplastic elastomer.

[0204] Non-limiting examples of suitable polyols for forming urethane-crosslinked elastomers include, for example, branched (3 branches or more) poly(lactic acid-co-caprolactone) (PLCL) and poly(glycolide-co-caprolactone) (PGCL) polyols. When cured with an isocyanate (such as hexamethylene diisocyanate) and other suitable reagents, in addition to branched polymers, linear polymer diols can also be used to produce an elastic coating. To reduce inflammation due to material degradation, poly(trimetylene carbonate) (PTMC)-based polyols can also be used to produce an elastic coating. Various catalysts, including but not limited to Sn(Oct)2, Zn(Oct)2, dibutyltin dilaurate (DBTL), 1,4-diazabicyclo[2.2.2]octane (DABCO), and 1,8-diazabicycloundec-7-ene (DBU), can be used to facilitate the curing process.

[0205] In certain embodiments, the stent and / or coating may be made using a shape memory polymer that can change size, shape, and / or conformation to mold to the sinus anatomy. Non-limiting examples of shape memory polymers include segmented polyurethanes made from isocyanates coupled with oligomeric lactic acid, oligomeric caprolactone, oligomeric lactide-co-glycolide, oligomeric (trimethylene carbonate), or oligomeric dioxanone and various chain extenders, (co)block copolymers of lactide (glycolide) and caprolactone, dioxanone, or trimethylene carbonate, and polymer blends of polylactide and polyamide elastomers.

[0206] As previously noted, the stent according to the present invention can be fiber-based or non-fiber-based. In fiber-based embodiments, the polymeric material can first be made into fibers having a cross-sectional dimension in the range of, for example, 10 μm to 1000 μm, more typically 100 μm to 300 μm. Such fibers can be formed using a number of techniques, including, for example, extrusion or spinning techniques.

[0207] The shape of the fiber cross-section can vary to a large extent. Referring to Figure 1A , such cross-sections include fibers having a circular cross-section 10, an elliptical cross-section 12, and a polygonal cross-section (e.g., a triangular cross-section 14, a quadrilateral cross-section 16 such as shapes like a rectangle, a parallelogram, a trapezoid, etc., a pentagonal cross-section, a hexagonal cross-section 18, etc.). During fiber preparation, the fiber cross-section can be changed by selecting a die for the appropriate cross-section.

[0208] The polymeric material can also be formed into sheets, for example, via suitable casting or extrusion processes (e.g., solvent casting, melt casting, solvent-based extrusion, melt extrusion, etc.). Thereafter, the sheet can be cut into fibers (e.g., fibers having a polygonal cross-section, such as shapes like a triangle or a quadrilateral like a rectangle, a parallelogram, a trapezoid).

[0209] In certain embodiments, the strength of the fibers can be optimized, for example, by stretching at an appropriate draw ratio or annealing at an appropriate temperature.

[0210] The strength and / or flexibility of the fibers can also be optimized by braiding fibers of homogeneous or heterogeneous cross-sections into a multi-fiber strand (e.g., a fish-wire-type structure). The braided fibers can have the same composition or different compositions. Moreover, the braided fibers can have the same diameter or different diameters. Figure 1B shows two embodiments that illustrate (a) the cross-section of a multi-fiber strand 11 formed from strands of the same material and having the same diameter, and (b) the cross-section of a multi-fiber strand 13 formed from strands of different compositions and having different diameters.

[0211] Once the polymeric chains are prepared, fiber-based scaffolds can be prepared therefrom. For example, single fiber chains and / or multi-fiber chains of various shapes (e.g., Figure 1A - 1B The diameter of the strands forming the braid can vary widely, ranging from 10 to 1000 μm, as well as other possibilities. In certain embodiments, the material forming the strands can have an elastic modulus in the range of about 1 GPa to about 10 GPa, more preferably in the range of about 4-9 GPa.

[0212] To facilitate the low profile aspect of the invention (e.g., delivery of the stent into a small diameter sinus cavity), in certain beneficial embodiments, the chains used to form the stent may have a diameter ranging from 100 to 500 μm, more advantageously ranging from 125 to 250 μm. The use of small diameter chains results in a stent having minimal wall thickness and the ability to fold (i.e., curl) within a low diameter catheter delivery system. In certain embodiments, the chain diameter may be selected so that the stent can be delivered with a 15 French delivery catheter or smaller, a 9 French delivery catheter or smaller, a 6 French delivery catheter or smaller, or even a 4 French delivery catheter or smaller.

[0213] Figure 2 The embodiment of the braided stent 100 is illustrated, and the braided stent includes at least one chain (e.g., a single fiber or multi-fiber chain) braided to form a substantially tubular structure, and the tubular structure has a length 130, a width 131, and a first and a second end 132, 133 along the longitudinal dimension. For example, the tubular structure can include two groups of chains 110 and 120, each group extending along the longitudinal dimension of the stent in a relative spiral configuration. In certain embodiments, the number of spiral chains forming the stent can range from, for example, 8 to 48 chains, as well as other possibilities. The groups of chains 110 and 120 cross each other at a braiding angle 140. The range of the braiding angle 140 can be, for example, about 30 degrees or less to about 150 degrees or more, and other values, such as any value ranging from 30 degrees to 40 degrees to 50 degrees to 60 degrees to 70 degrees to 80 degrees to 90 degrees to 100 degrees to 110 degrees to 120 degrees to 130 degrees to 140 degrees to 150 degrees (i.e., ranging between any two of the aforementioned numerical values). The chains may be woven together using a variety of methods known in the art, such as various 1×1, 1×2, and 2×2 designs, and may use specific known weaving patterns such as the Regular pattern "1 thread, 2-up / 2-down", the Diamond half-load pattern "1 thread, 1-up / 1-down", or the Diamond pattern "2 threads, 1-up / 1-down".

[0214] Multiple factors contribute to the radial strength of the stent 100, including the diameter of the strands, the braiding angle 140, the strand material, and the number of strands used, etc.

[0215] Each strand can cross each other at a certain braiding angle, which can be a constant value or can vary around the circumference of the stent and / or along the longitudinal dimension of the stent. Figure 3A Embodiments are shown in which the stent 100 has a constant braiding angle, while Figure 3B Embodiments are shown in which the stent 100 has strands with a variable braiding angle. In Figure 3B a particular embodiment, a first region 100a of strands having a first braiding angle transitions to a second region 100b of strands having a second braiding angle, which is less than the first braiding angle. Various filament braiding patterns can be used to prepare such constructs.

[0216] Potential features of a design with a variable braiding angle include one or more of the following, where: (1) it allows for targeting portions with a specific density for preferential therapeutic agent delivery; (2) it allows for adjusting the radial force based on the stent location; and (3) it can be used to provide a tapered tubular design for non-cylindrical tissue structures.

[0217] Generally, the shape and diameter of the stent according to the present invention can vary along the length of the device. In some embodiments, in a cylindrical design, the diameter at the ends of the device can be greater than the diameter at the center point (e.g., dumbbell or hourglass shape). For example, the diameter at the ends of the device can be 1.5 times or more, even 2 times or more, the diameter at the center point. As another example, the shape of the device can be triangular at one end and hexagonal at the other end.

[0218] The radial stiffness of the braided stent can be adjusted by partially or fully locking a plurality of strand intersections (also referred to as "nodes"). The nodes can be partially or fully locked, for example, by welding the strands at the intersection, such as using heat (e.g., using a suitable laser such as a pico or femto laser), by using a suitable adhesive, by wrapping the intersection with a suitable filament, or by coating the intersection with a suitable material that binds the filaments together, as well as other possible techniques. In certain embodiments, an elastomer can be coated on the braid, for example, using methods such as those described in U.S. Patent Nos. 8,137,396, 8,540,765, and 8,992,601, the disclosures of which are incorporated herein by reference.

[0219] The underlying braid, with or without previously locked nodes, can be elastomer-coated. One embodiment is illustrated in Figure 4In it, a braided stent 100 completely coated with a display elastomer 100e is shown. By changing the nodes, the accumulation of the coated elastomer can simultaneously optimize the radial resistance force (RRF) and the chronic outward force (COF) for this sinus cavity.

[0220] In alternative embodiments, a stent may be provided in which the nodes of a braided structure are connected using elastic elements such as elastic filaments or chains. An example of one such embodiment is illustrated in Figure 5 In it, a stent 100 is shown, in which elastic chains 111 are connected to the braid 110 at different points along the length of the braid 110. In these embodiments, the braid itself provides a framework for supporting the sinus cavity, while elastic filaments or chains 111 are provided to enhance stent recovery during deployment. The elastic filaments or chains 111 may be woven into the stent 100, for example, during the weaving process, or introduced therein after the stent 100 is formed. In the latter case, the braid 110 forming the stent 100 may be made of a softer, non-elastic material that conforms to the sinus wall and has desired degradation characteristics. The number of elastic filaments or chains 111 used in a given stent 100 may be adjusted to obtain a suitable recovery stiffness and a radial stiffness.

[0221] In multiple embodiments, conformable stents are desirable because they can be used to improve apposition to the contacting tissue, reduce damage to the contacting tissue, and, when delivering therapeutic agents, also enhance the delivery efficacy of the therapeutic agent due to increased tissue contact.

[0222] Various strategies can be used to enhance the conformability of braided stents. For example, in some embodiments, some or all of the nodes of the braid may be partially locked or fully unlocked to allow at least some filaments to have the freedom to slide relative to each other at least somewhat. In a confined space, a stent with filaments having free movement will tend to better conform to the surrounding environmental geometry.

[0223] Optionally or additionally, the stent may be braided from filaments of different stiffnesses (e.g., a combination of higher stiffness and lower stiffness). For example, the stiffness of a given chain is determined by its intrinsic material properties, by its processing conditions, and by its dimensions. An embodiment of such a stent is schematically illustrated in Figure 6A In it, a stent 100 is shown that is formed from higher stiffness chains 112 and lower stiffness chains 113. In a particular embodiment, the higher stiffness chains 112 may have an elastic modulus greater than 3 GPa and a filament diameter greater than 100 μm, while the lower stiffness chains 113 may have an elastic modulus less than 3 GPa and a filament diameter less than 200 μm. The filaments with lower stiffness may provide a weaker point in the stent to allow the sinus cavity to deform, while the filaments with higher stiffness may maintain mechanical integrity.

[0224] Adaptability can also be improved by removing some of the strands from within the braided structure. An embodiment of such a stent is schematically illustrated in Figure 6B wherein a stent 100 is shown with variable-sized cells formed therein. In particular, a stent containing larger diamond-shaped cells 114 and smaller diamond-shaped cells 115 is shown. In some embodiments, the cut strands can be locked partially or fully at the intersection closest to the cut tip.

[0225] In related embodiments, different-sized cells are created during the braiding process, e.g., via selection of an appropriate braiding pattern. An embodiment of a stent 100 with larger braided cells 114 and smaller braided cells 115 is shown in Figure 7 .

[0226] When different-sized cells are formed, the larger cells can have an area that is 1.1 to 10 times or more times the area of the smaller cells.

[0227] A potential advantage of a stent with a combination of larger and smaller cells is that the larger cells can provide flexibility (e.g., ease of crimping and better adaptability), while the smaller cells can maintain mechanical integrity.

[0228] Another approach for producing an adaptable stent is to braid a stent using a hard rubber material such as carbon fiber-reinforced silicone, poly(acrylonitrile butadiene), and poly(styrene butadiene) rubber. This results in a fully elastic braid.

[0229] In some embodiments, a coating can be formed over the entire stent structure or a portion thereof. By applying a relatively non-elastic material to the coating (e.g., a coating formed using a relatively hard polymer such as D,L-PLGA), the stiffness of the stent can be improved. Moreover, when an implant is formed using an inherently elastic braid and when the coating is a degradable layer, the stent strands have increased adaptability when the degradable layer degrades.

[0230] Moreover, the coating can be applied in a pattern so as to adjust the adaptability of the braided stent. In one particular example, Figure 6C an embodiment of a stent 100 is shown wherein coated regions 116 are provided at each end of the stent 100, while uncoated regions 117 are provided at the center of the stent 100. Such a design can be used to provide ends with enhanced stiffness, which can allow the device to better anchor into the sinus cavity. Moreover, leaving the middle region of the stent 100 uncoated can increase the ability of the stent to conform to the shape of the sinus cavity upon implantation. As another particular example, Figure 6DShows an embodiment of the display stent 100, where the coated area 116 and the uncoated area 117 are provided in an alternating pattern. The presence of the uncoated area 117 provides a region of increased flexibility along the length of the stent, which can provide increased adaptability to irregular surfaces such as those associated with the sinus cavity.

[0231] Regions of coated and uncoated material can be provided using a variety of techniques. For example, in certain embodiments, a mask can be used in a spraying method to produce a specific pattern of coated and uncoated areas.

[0232] By longitudinally masking a portion of the tubular braid during spraying, a U-shaped coated area (when viewed longitudinally) can be produced. In these stents, the coated area is relatively rigid, while the uncoated area is relatively soft. After deployment into the sinus cavity, the coated area will provide stent recoverability. On the other hand, the soft uncoated area can be easily deformed to adapt to the irregular surface of the sinus cavity, resulting in optimized adaptability. In a specific embodiment, in the selected case where an opening is made between the left and right paranasal sinuses, such a stent can be used to keep it open.

[0233] In certain embodiments, the stent can be longitudinally cut, allowing the stent perimeter to easily change size when deployed to match the sinus cavity, which can provide better compliance and adaptability.

[0234] To reduce potential tissue irritation or patient discomfort due to sharp stent edges, the stent edges can be coated and / or the braided stent can be prepared such that the filament ends turn back towards the center of the stent. For example, the filament ends can be woven back into the stent structure and joined, for example, at the node connection. The connection can be made using the techniques described above for joining filaments at the nodes (e.g., by welding, applying a suitable adhesive, applying an elastic coating, etc.). A schematic example of this type of stent 100 is illustrated in Figure 8 where the end filaments 118 turn back towards the center of the stent 100.

[0235] In cases where difficulties may be encountered when braiding a short stent onto a large-diameter mandrel (i.e., when forming a stent with a large diameter-to-length ratio), a zig-zag chain, including single- and multi-fiber chains, can be fabricated prior to braiding. Then, the zig-zag chain can be wound or looped around the mandrel, preferably in a braided pattern. The filament ends can be joined using, for example, the techniques described above for joining filaments at the nodes (e.g., welding, applying a suitable adhesive, applying an elastic coating, etc.) to complete the braided structure. The size and shape of the final stent can be controlled by the turns, directions, and strut lengths of the zig-zag filaments. Such a braid can also have folded ends as described above.

[0236] In some embodiments, it is beneficial to provide a sinus stent that has the ability to be easily removed, if desired. In the case of a relatively soft braided stent, a tool having one or more hooks at the end can be used to capture the distal end of the implanted stent. Optionally, the device can be removed by standard surgical instruments available to an ENT surgeon. Then, the braid can be flipped by stretching the ends so that the outer surface becomes the lumen. Thus, the stent can be removed by peeling it from the sinus wall, reducing additional abrasion, irritation, and damage to the sinus tissue.

[0237] Other stents are based on non-braided structures or hybrid braided / non-braided structures.

[0238] For example, in various embodiments, a stent formed by braided or knitted links is provided. A stent 100 in the form of a knitted tube is illustrated in Figure 9 . Such braided or knitted stents can provide the mechanical properties necessary for providing a stenting function, while also having increased compliance and adaptability (and in certain embodiments, facilitating therapeutic agent delivery). Additionally, in certain embodiments, in the case of a knitted structure, one end of the single strand used to form the tube can be stretched to unravel the stent, enabling removal of the stent.

[0239] In many other embodiments, the stent can be in a helical (e.g., spiral) form. In some of these embodiments, the helical form can be formed from a single strand (e.g., a single fiber strand or a multi-fiber strand). An example of such a stent 100 is schematically illustrated in Figure 10 .

[0240] In other embodiments of these embodiments, the helical form can be formed from a multi-strand construct. Examples of multi-strand constructs include, for example, a substantially two-dimensional structure (e.g., a ribbon-like structure), which can be formed into a helical form. Two embodiments of such helical stents are shown in Figure 11A and Figure 11B . In the embodiment shown in Figure 11A , the stent 100 in the form of a helical shape is formed from an existing tubular structure, such as, for example, a braided tubular structure (e.g., one of those described previously), which is subsequently cut into a helix. In the embodiment shown in Figure 11B , the stent 100 is formed as follows: by shaping a previously formed substantially two-dimensional braided pattern 119 into a helical structure, such as placing the substantially two-dimensional braided pattern on a mandrel and annealing it for a period of time at a temperature suitable for the two-dimensional braided pattern to form a helical shape. Examples of such braided patterns include multi-carrier braided patterns, such as a 2-carrier (shown in Figure 11B ), 3-carrier, 4-carrier, etc. braided patterns.

[0241] It is noted that scaffolds similar to the various braided structures described herein can be in the form of a single polymeric structure. Compared to using fiber-based techniques, using a single polymeric structure can provide a reduced profile, where the fiber-based techniques result in a minimum profile of the sum of the overlapping chain widths. An embodiment of such a structure is shown in Figure 15 wherein scaffold 100 is illustrated and characterized as a regular repeating pattern, such as a lattice structure. When the scaffold 100 is a single polymeric structure, it can be fabricated using various suitable techniques, such as by mechanically cutting or laser cutting a pattern into a solid polymer tube or a solid polymer strip.

[0242] In a variety of other embodiments, the scaffold can be in the form of an open cylinder. For example, as shown in Figure 12A and 12B scaffold 100 can be formed as a series of individual rings 121 that are axially aligned with each other and connected at one end. In the embodiment shown in Figure 12A the individual rings 121 are solid rings (e.g., in the form of a strip). In the embodiment shown in Figure 12B the individual rings 121 include units, particularly diamond-shaped units, similar to those formed with a 2-carrier braid. The benefits of these designs can include one or more of the following, where: the rings are directly curled to a certain size; and when delivered, the scaffold opens to an expanded diameter. Since each ring allows for expansion to a different width, the scaffold may be more conformable in a space of variable size.

[0243] In still other embodiments, the scaffold 100 can be in the form of a polymeric tube, such as that shown in Figure 13 Such a device is beneficial, for example, because it can conform to the sinus wall 200 and optionally release one or more therapeutic agents. The tube can be made of materials such as the above-described scaffold materials, and includes hot-forged PCL, or PLCL with a high caprolactone content, as well as many other possible materials.

[0244] In a related device design, a tubular conformable scaffold, such as that shown in Figure 13 can be connected to a collapsible three-dimensional support structure that can assist in expanding and supporting the sinus of the tubular scaffold. A particular example of such a device is shown in FIG. 14, which shows a tubular scaffold 110 connected to a collapsible three-dimensional structure 122. For minimally invasive delivery, the structure 122 is collapsible. The structure 122 can be provided at the end of the tubular scaffold 110 and, if desired, at one or more points along the length of the tubular scaffold 110. Examples of materials suitable for forming the collapsible three-dimensional structure include degradable or non-degradable elastic materials that can be compressed and recover from deformation. In other embodiments, a braided scaffold structure such as those discussed above can be used as the collapsible three-dimensional structure.

[0245] Supplements such as therapeutic agents and inert release control agents can be incorporated into the various devices described herein.

[0246] Examples of therapeutic agents are any suitable reagents having a desired biological effect, including small molecule reagents, biological agents, cells including stem cells, gene therapy agents, and RNAi, etc. Specific examples of therapeutic agents include: analgesics, including simple analgesics such as aspirin and paracetamol, non-steroidal anti-inflammatory drugs such as ibuprofen, diclofenac, naproxen, celecoxib, ketoprofen, piroxicam, and sulindac, and opioids such as codeine, tramadol, dextropropoxyphene, paracetamol, morphine, oxycodone, and pethidine hydrochloride; anesthetics such as lidocaine, bupivacaine, and ropivacaine; statins such as atorvastatin, cerivastatin, fluvastatin, lovastatin, mevastatin, pitavastatin, pravastatin, rosuvastatin, and simvastatin; steroidal anti-inflammatory drugs such as glucocorticoids, mometasone furoate, clobetasone dipropionate, budesonide, ciclesonide, flunisolide, fluticasone furoate, fluticasone propionate, dexamethasone, cortisone, prednisone, methylprednisolone, triamcinolone acetonide, betamethasone, dexamethasone, prednisolone, corticosterone, estrogen, sulfasalazine, rosiglitazone, mycophenolic acid, and mesalazine; antihistamines, including H1-receptor antagonists such as diphenhydramine, loratadine, fexofenadine, cyproheptadine, promethazine, desloratadine, chlorpheniramine, hydroxyzine, and pyrilamine, and H2-receptor antagonists such as cimetidine, famotidine, lafutidine, nizatidine, ranitidine, roxatidine, and tiotidine; antimicrobials such as mupirocin, gentamicin, and tobramycin; antibiotic reagents such as penicillin, cefoxitin, oxacillin, and tobramycin; endostatin, angiostatin, and thymidine kinase inhibitors, and their analogs or derivatives; anti-leukotriene reagents (e.g., montelukast, zafirlukast, zileuton, etc.); antifungal agents; and probiotics, as well as other reagents.

[0247] Further examples of therapeutic agents can be selected from antithrombotic agents such as heparin, heparin derivatives, urokinase, and PPack (dextrophenylalanine proline arginine chloromethylketone), enoxaparin, hirudin; antiproliferative agents such as angiotensin, or monoclonal antibodies capable of blocking smooth muscle cell proliferation, acetylsalicylic acid, paclitaxel, sirolimus, tacrolimus, everolimus, zotarolimus, vincristine, dasatinib (sprycel), amlodipine, and doxazosin; immunosuppressive agents such as sirolimus, tacrolimus, everolimus, zotarolimus, and dexamethasone; antineoplastic / antiproliferative / anti-mitotic agents such as paclitaxel, 5-fluorouracil, cisplatin, vinblastine, cladribine, vincristine, epothilone, methotrexate, azathioprine, halofuginone, doxorubicin, actinomycin, and mutamycin; anticoagulants such as D-Phe-Pro-Arg chloromethyl ketone, RGD peptide-containing compounds, heparin, antithrombin compounds, platelet receptor antagonists, antithrombin antibodies, anti-platelet receptor antibodies, aspirin (which is also classified as an analgesic, antipyretic, and anti-inflammatory drug), dipyridamole, hirudin, prostaglandin inhibitors, platelet inhibitors, and antiplatelet agents such as trapidil or liprostin, tick antiplatelet peptide; DNA demethylating drugs such as 5-azacytidine, which is also classified as an RNA or DNA metabolite that inhibits cell growth and induces apoptosis in certain cancer cells; vascular cell growth promoters such as growth factors, vascular endothelial growth factor (VEGF, all types including VEGF-2), growth factor receptors, transcriptional activators, and translation initiators; vascular cell growth inhibitors such as antiproliferative agents, growth factor inhibitors, growth factor receptor antagonists, transcriptional blockers, translational repressors, replication inhibitors, inhibitory antibodies, antibodies directed directly against growth factors, bifunctional molecules including growth factors and cytotoxins, bifunctional molecules including antibodies and cytotoxins; cholesterol-lowering agents; vasodilating agents; and agents that interfere with endogenous vasoactive mechanisms; antioxidants such as probucol; angiogenic substances such as acidic and basic fibroblast growth factors, estrogen, including estradiol (E2), estriol (E3), and 17-β estradiol; drugs for heart failure such as digoxin, β-blockers, angiotensin-converting enzyme (ACE) inhibitors, including captopril and enalapril, statins, and related compounds; macrolides such as sirolimus and everolimus;and reagents having a primary mechanism of action of inhibiting extracellular matrix remodeling and a secondary mechanism of action of inhibiting cell proliferation, such as 5-fluorouracil, doxycyclin, carvedilol, curcumin, and tranilast.;

[0248] Other therapeutic agents include bacteria or other microbial communities that may be beneficial for re-establishing a healthy microbiome in the nasal cavity and sinuses, as well as reagents or nutrients that can promote a healthy microbiome.

[0249] Inactive release control agents may also be included to enhance control over the release kinetics of the therapeutic agent. Examples of inactive release control agents include soluble polymers such as polyethylene glycol (PEG) (also known as poly(ethylene oxide), PEO), PEG-vinyl alcohol copolymers, polyacrylates and polymethacrylates containing cationic and anionic functionalities, polyvinylpyrrolidone, and dextran, as well as small molecule additives such as cyclodextrins or citrate esters such as acetyl tributyl citrate (ATBC) or acetyl triethyl citrate (ATEC).

[0250] In embodiments where the scaffold delivers one or more therapeutic agents at the implantation site, the therapeutic agent can be provided in the device for delivery therefrom in a variety of ways.

[0251] For example, the therapeutic agent can be directly incorporated into a polymeric construct (e.g., fibrils, sheets, rigid tubes, etc.) and subsequently used to form a generally tubular scaffold as described herein. In one embodiment, the therapeutic agent and the polymer are dissolved in a suitable solvent to obtain a homogeneous solution or suspension, or the therapeutic agent and the polymer are heated to form a polymer melt containing the therapeutic agent. Then, the solution, suspension, or melt is subjected to suitable solvent-based processing or melt-based processing, such as extrusion, wet spinning, dry spinning, melt spinning, electrospinning, or other methods, to obtain a therapeutic agent-loaded polymeric construct (e.g., fibrils, sheets, tubes, etc.) having the therapeutic agent incorporated therein. In certain embodiments, a polymeric region that does not contain the therapeutic agent (e.g., a polymeric core) is co-extruded with a therapeutic agent-loaded polymeric coating to form a therapeutic agent-loaded polymeric construct.

[0252] In certain embodiments, then, the therapeutic agent-loaded polymeric construct is then processed into an additional form and subsequently used to form a generally tubular scaffold as described herein. As a further example, a solvent-cast therapeutic agent-loaded polymeric construct in sheet form can be prepared by controlling the evaporation of a solution of the therapeutic agent and one or more carrier polymers. After removal of the solvent, the therapeutic agent-loaded polymeric sheet can be laser cut into a therapeutic agent-loaded polymeric construct in the form of flat fibrils for use in fabric preparation.

[0253] In certain embodiments, using suitable application techniques such as spraying, dip coating, roll coating, or vapor deposition, in the presence or absence of a carrier material (e.g., a polymeric coating material such as those described above), a therapeutic agent can be applied to a pre-formed construct (e.g., a filament, sheet, or tube, including a pre-formed device scaffold). The release characteristics of the therapeutic agent can be adjusted by, for example, the thickness of the coating, addition of inactive ingredients, and when a polymer is provided as a carrier, changing the carrier polymer (e.g., changing the composition and / or molecular weight of the polymer) and / or the ratio of the therapeutic agent to the polymer.

[0254] The delivery of the therapeutic agent from the device to body tissue can also be modulated by a surface layer that does not contain a therapeutic agent polymer layer. In embodiments involving a non-biodegradable surface layer, the surface layer can act as a diffusion barrier such that the delivery rate of the therapeutic agent is limited by its rate of diffusion through the surface layer. In certain embodiments involving a biodegradable surface layer, the surface layer can also act as a diffusion barrier such that the delivery rate of the therapeutic agent is limited by its rate of diffusion through the surface layer. In other embodiments involving a biodegradable surface layer, the therapeutic agent cannot diffuse through the surface layer such that its delivery is simply delayed until degradation of the surface layer is complete.

[0255] Electrospinning provides another potential method for introducing a therapeutic agent onto a pre-formed construct. In one embodiment, a fibrous or non-fibrous substrate can be covered with an electrospun fiber mesh (such as a core-sheath). During electrospinning, the therapeutic agent can be dissolved or suspended in the core polymer solution. The release characteristics of the therapeutic agent can be adjusted by modulating the therapeutic agent loading, the particle size of the therapeutic agent (when using a suspension), the types of polymers used to form the core and sheath, respectively, and the thickness of the sheath.

[0256] In other embodiments, core-sheath fibers are first fabricated by coaxial electrospinning a core polymer solution or suspension containing the therapeutic agent and a sheath polymer solution. The therapeutic agent-loaded fibers can be further woven onto a multi-fiber strand used to prepare the device. For example, a fish-line type composite strand can be formed and then fabricated into a woven scaffold as described above. In these designs, the release of the therapeutic agent can be determined by the electrospun fibers.

[0257] In certain embodiments, bioactive agents such as proteins and / or polysaccharides can be incorporated into the electrospun fibers.

[0258] In certain embodiments, the devices described herein can be used in conjunction with sinuplasty. For example, a stent such as those described herein can be deployed into a sinus cavity with the assistance of an expandable device such as an expandable frame (e.g., an expandable lattice) or a balloon, among other possibilities. In such an embodiment, a stent according to the present invention can be positioned above, within, below, proximal, or distal to the expandable device at the time of preparation or during delivery by a healthcare professional. The expandable device can be a drug-eluting device (e.g., via a drug-containing coating disposed on the expandable device) or a non-drug-eluting device. Examples of therapeutic agents that can be released by a drug-eluting device are those described above.

[0259] Thus, in the case of a balloon, the balloon can be coated or uncoated, and a stent according to the present invention can be positioned above, within, below, proximal, or distal to a balloon catheter suitable for sinuplasty at the time of preparation or during delivery by a healthcare professional. The catheter can include an inflatable balloon assembly disposed at or near the distal end of the catheter shaft, the catheter shaft including a balloon inflation lumen. In the uninflated state, the balloon assembly does not significantly increase the overall width of the distal end of the catheter. This allows the distal end of the catheter to be inserted into a patient and guided to a desired treatment site within the patient's sinuses. Once at the treatment site, the balloon assembly is inflated to position the stent against the sinus wall adjacent to the treatment site. The balloon assembly can include any number of individual balloons of many configurations, depending on the treatment site. Additionally, sinuplasty can be completed before, after, simultaneously with, or in any combination of the sequence of the perioperative process of delivering the stent.

[0260] In certain embodiments, the devices described herein can be used as an adjunctive therapy. For example, a stent such as those described herein can be deployed into a sinus cavity using a therapeutic agent-eluting delivery device such as, for example, a therapeutic agent-eluting balloon. Optionally, the stent can be deployed into the sinus cavity after treating the sinus cavity with a spray containing a therapeutic agent such as a hydrogel spray, or a flushing liquid containing one of the therapeutic agents described previously herein.

[0261] The stents of the present invention can be radiopaque so that they can be visible using conventional fluoroscopic techniques. In one embodiment, radiopaque additives are included within the polymeric material of the stent and / or its coating when present. Examples of suitable radiopaque additives include particles containing iodine, bromine, barium sulfate, platinum, iridium, tantalum, and / or palladium. In another embodiment, radiopaque groups such as iodine are incorporated into the polymer backbone. In still another embodiment, one or more biostable or biodegradable radiopaque markers, such as those containing platinum, iridium, tantalum, and / or palladium, can be fabricated in the form of tubes, coils, wires, balls, or disks and then placed at the ends of the stent or other predetermined locations thereon.

[0262] To facilitate delivery, the stent can be loaded into the delivery catheter just prior to implantation in a patient. Loading the stent in close temporal proximity to implantation avoids the possibility of the polymer of the stent relaxing within the delivery catheter during transportation, storage, etc. Thus, one aspect of the present invention includes delivering the stent of the present invention, including the step of loading the stent into the delivery catheter within a short period of time, such as within one hour prior to implanting into a body cavity. However, it should be noted that prior to implantation, the stent of the present invention does not need to be loaded into the delivery catheter.

[0263] In some embodiments, a stent suitable for implantation into the void space formed during ethmoidectomy and other uses (e.g., using a 6 mm catheter and other devices) can be provided. The diameter of such a stent can range, for example, from 10 to 30 mm, more particularly from 15 to 20 mm, and other possible values. The length of such a stent can range, for example, from 5 to 20 mm, more particularly from 8 to 12 mm, and other possible values. In certain advantageous embodiments, the stent comprises a braided stent material, which can include, for example, from 8 to 64 braided strands, more particularly from 16 to 32 braided strands, and other possible values. In certain advantageous embodiments, the braiding angle can vary within, for example, from 30 to 150 degrees, more particularly from 60 to 130 degrees, and other possible values. In certain advantageous embodiments, the diameter of each strand forming the braid can vary within the range of 50 to 500 μm, more particularly from 150 to 300 μm, and other possible values. In certain advantageous embodiments, the stent mass range can be, for example, from 1 to 20 mg / mm length, more particularly from 2 to 10 mg / mm, and other possible values. In certain advantageous embodiments, after being compressed to 30% of the unconstrained diameter for 10 minutes, the stent has at least 85% of the recovered diameter %. When the drug is released, in non-refractory patients, the drug can be released during 3 to 6 weeks and other values, while in refractory patients, the drug can be released during 8 to 26 weeks and other values.

[0264] In some embodiments, a stent suitable for implantation into the middle meatal space, among other uses, may be provided (e.g., using a 3-4 mm catheter, among other devices). The diameter of such a stent may range, for example, from 5 to 20 mm, more particularly from 10 to 15 mm, and other possible values. The length of such a stent may range, for example, from 5 to 20 mm, more particularly from 8 to 12 mm, and other possible values. In certain advantageous embodiments, the stent comprises a braided stent material, which may comprise, for example, from 8 to 64 braided strands, more particularly from 16 to 32 braided strands, and other possible values. In certain advantageous embodiments, the braiding angle may vary, for example, from 30 to 150 degrees, more particularly from 60 to 130 degrees, and other possible values. In certain advantageous embodiments, the diameter of the strands forming the braid may range from 100 to 500 μm, more particularly from 150 to 300 μm, and other possible values. In certain advantageous embodiments, the stent mass may range, for example, from 1 to 20 mg / mm in length, more particularly from 2 to 10 mg / mm in length, and other possible values. In certain advantageous embodiments, after 10 minutes of compression to 30% of the unconstrained diameter, the stent has at least 85% of its recovery diameter %. In certain advantageous embodiments, when measured in an MSI radial force tester, at a diameter less than the as-prepared diameter, the stent has an RRF in the range of 30 to 500 mmHg, and other possible values. In certain advantageous embodiments, when measured in an MSI radial force tester, at a diameter less than the as-prepared diameter, the stent has an acute COF in the range of 5 to 100 mmHg, and other possible values. When the drug is released, it may be released over a period of 8 to 26 weeks and other values.

[0265] In some embodiments, a stent may be provided that is adapted for implantation into a sinus ostium, and for other uses (frontal, maxillary, or sphenoid sinuses) or into the frontal recess. Such a stent may be delivered, for example, using a 3 - 4 mm delivery catheter, and other possible devices. The diameter of such a stent may range, for example, from 4 to 20 mm, more particularly from 6 to 10 mm, and other possible values. The length of such a stent may range, for example, from 5 to 20 mm, more particularly from 6 to 12 mm, and other possible values. In certain advantageous embodiments, the stent comprises a braided stent material, which may comprise, for example, from 8 to 64 braided strands, more particularly from 16 to 32 braided strands, and other possible values. In certain advantageous embodiments, the braiding angle may vary, for example, from 30 to 150 degrees, more particularly from 60 to 130 degrees, and other possible values. In certain advantageous embodiments, the diameter of each strand forming the braid may vary within a range of 100 to 500 μm, more particularly from 150 to 300 μm, and other possible values. In certain advantageous embodiments, the stent mass range may be, for example, from 1 to 20 mg / mm length, more particularly from 2 to 10 mg / mm length, and other possible values. In certain advantageous embodiments, after 10 minutes of compression to 30% of the unconstrained diameter, the stent has at least 85% of the recovered diameter %. In certain advantageous embodiments, when measured in an MSI radial force tester, at a diameter less than the as - fabricated diameter, the stent has an RRF ranging from 30 to 500 mmHg, and other possible values. In certain advantageous embodiments, when measured in an MSI radial force tester, at a diameter less than the as - fabricated diameter, the stent has an acute COF ranging from 5 to 100 mmHg, and other possible values. When the drug is released, it may be released over a period of 6 to 26 weeks and other values.

[0266] In some aspects, the stents described herein may be provided in a kit, the kit comprising (a) one or more stents, (b) a delivery catheter, and (c) an optional loading aid (e.g., a crimping mechanism), and other components.

[0267] Example 1

[0268] First, a uniform braided stent was prepared using a PLGA (85:15) copolymer by winding fiber - spun monofilaments around individual spools (see, for example Figure 3A)。Place each spool on a braiding machine, tighten it through rollers and eyelets, and wind it around a mandrel of the desired OD (e.g., 7, 8, or 10 mm). Set the braiding tension of the machine according to the size requirements of the monofilament. To obtain a braiding angle with optimal properties (including radial strength), set the picks per inch (pix / inch). Select a braiding pattern and braid the monofilament from the coil onto the mandrel through the braiding machine. Use a winding tape at the end of each mandrel to maintain tension on the filaments, which can be used for thermal annealing and obtaining high modulus properties. Thermally anneal the braided polymer on the mandrel, then cut it to the desired length with a blade and remove it from the mandrel.

[0269] Coat the braided PLGA scaffolds with a carrier coating made of poly(L-lactide-co-ε-caprolactone) (PLCL) cured with hexamethylene diisocyanate (HDI) in the presence of 1-dodecanol (DD) as a chain terminator and optionally a catalyst. Specifically, four-arm hydroxyl-terminated PLCL (40:60) (mol / mol), HDI, and DD were dissolved in dichloromethane to obtain a stock solution for spraying. The solution was sprayed onto the braided scaffolds. After sufficient curing at high temperature, the scaffolds were cut into various lengths for radial force and recovery tests. Figure 16 Show macroscopic images of 8 mm, 10 mm, 20 mm, and 31 mm scaffolds, each having 16 strands and a braiding angle of about 100 - 135 degrees. Some properties of these and similar scaffolds with different prepared diameters, strand numbers, and braiding angles are compiled in Table 1 (where @[D0-1]mm means the measurement result is 1 mm below the prepared diameter, i.e., 6 mm for a 7 mm scaffold, 7 mm for an 8 mm scaffold, and 9 mm for a 10 mm scaffold). All scaffolds showed excellent diameter recovery (Rec. %) after simulated use. They have variable radial stiffness (RRF and COF), depending on the design.

[0270] Table 1

[0271]

[0272] Example 2

[0273] The scaffolds prepared in Example 1 were further coated with an additional conformal coating containing a mixture of PLCL and mometasone furoate (MF) as an active agent. The PLCL in the MF-containing coating included about 70% (mol%) lactic acid, and the balance was caprolactone (PLCL 70:30). Prepare a homogeneous solution of MF and PLCL in dichloromethane (DCM). Then, spray the DCM solution onto 7 mm scaffolds with 24 strands.

[0274] The amount of MF carried by each stent is controlled by the thickness of the MF-containing coating and the loading rate. By controlling the thickness between <1 μm and 10 μm and the loading rate of MF relative to the total dry coating weight of about 1 wt% to about 40 wt%, the inventors found that for a 7 mm diameter stent, the drug loading is advantageously about 10 to 2400 μg / 10 mm stent length, more advantageously 100 to 1600 μg / 10 mm stent length. Figure 17A shows the cumulative MF released by mass for different drug loading rates (5, 10, and 20 wt% corresponding to 100, 200, and 400 μg MF / 10 mm stent length). Importantly, the inventors found that the drug release percentage curve is slightly affected by the drug-loading rate within a certain range (see Figure 17B ).

[0275] Example 3

[0276] To provide a more linear release curve, a top layer containing PLCL (70:30) and PLA was further coated on the drug-coated stent. A homogeneous solution of 0.75 wt% PLCL and 0.25 wt% PLA in DCM was prepared. Then, the DCM solution was spray-coated as a single coat with variable coating passes on a 7 mm stent with 24 struts to obtain different top layer thicknesses. As Figure 18 shown, the MF release was adjusted by changing the thickness of the top layer. The thicker the top layer, the slower the drug release. By combining different drug loading rates with this method, different daily doses with programmable release durations can be easily obtained.

[0277] Example 4

[0278] Biodegradable polymers such as D,L-PLGA were also used as drug carriers. A conformal coating containing a mixture of D,L-PLGA and mometasone furoate (MF) as the active agent was formed. The coating contained 20 wt% of MF. The D,L-PLGA in the mometasone-containing coating included D,L-PLGA having about 50% lactide and 50% ε-caprolactone (50:50) (mol%), D,L-PLGA having about 75% lactide and 25% ε-caprolactone (75:25), or D,L-PLGA having about 85% lactide and 15% ε-caprolactone (85:15). In each case, a homogeneous solution of MF and D,L-PLGA in anisole / ethyl formate (50:50 v / v) was prepared. Then, the solution was spray-coated on a 7 mm stent with 24 struts.

[0279] As Figure 19As shown, the release of drugs from the scaffolds with these polymers as coatings is significantly slower than those coated with PLCL(70:30). In the case of D,L-PLGA, without wishing to be bound by theory, drug release is likely controlled by the degradation of the carrier polymer, and drug molecules are released at a later stage after the polymer begins to degrade.

[0280] In this regard, scaffolds with a double-layer drug coating can be prepared to obtain a sustained release of MF over a long period of time. For example, a surface layer containing PLCL(70:30) and MF can be formed on a bottom layer containing DL-PLGA and MF. Without wishing to be bound by theory, in the initial stage, it is believed that drug release is dominated by diffusion-controlled release of MF from the surface layer, while in the later stage, the release of the drug in the bottom layer is related to the degradation of DL-PLGA.

[0281] Example 5

[0282] A scaffold composed of 16 monofilament strands (0.0065″ filament diameter, PLGA 85:15) was woven on a large-diameter mandrel (3.175 cm) at 25 picks per inch in a 1×1 weave pattern. Then, the scaffold was annealed at 130 °C for 24 hours, cut to the working length, and then placed on a fixture for spraying.

[0283] An elastomer solution was prepared using 5 wt% PLCL(40:60) dissolved in DCM. A crosslinking agent, hexamethylene diisocyanate (45:1 NCO:OH), and a zinc octoate catalyst (0.1 wt%) were added to the final solution. The elastomer solution was sprayed on the scaffold and cured at 100 °C for 24 hours in an open vial. A photograph of one such prepared scaffold is shown in Figure 20A and a photograph of the coating nodes of such a scaffold is shown in Figure 20B In.

[0284] A flat plate compression test was carried out to evaluate the mechanical properties of the cured scaffold. The scaffold was longitudinally compressed to at most 50% of its initial diameter. The results are shown in Figure 21 where the compressive load is per unit length of the scaffold.

[0285] Example 6

[0286] The multifilament chain was prepared by winding two 0.007″ PLGA 85:15 monofilament chains together. Then, the multifilament chain was hand-woven into various braided patterns using a fixture. An example of the fixture used to prepare the multifilament scaffold is shown in Figure 22. The same fixture was also used to prepare scaffolds using monofilament chains. After braiding, the filament ends were secured to the fixture using tape and then annealed overnight at 100 °C to set the filaments and maintain the filament intersections. Then, the scaffolds were sprayed with an elastomeric solution containing 5 wt% PLCL 40:60, HDI (45:1 NCO:OH), and zinc catalyst (0.1 wt%) in methylene chloride. All scaffolds were cured in open vials at 100 °C for 24 hours.

[0287] Table 2 contains data generated after spraying from five different braided patterns (shown in Figure 22A - 22E ). The diameter, weight, braid angle, acute recovery, and recovery after deployment were measured for all scaffolds.

[0288] Table 2

[0289]

[0290] Recovery tests were performed by using an external braided mesh sheath and curling and transferring the scaffold through a series of tubes from large to small until a curled diameter of 4 - 5 mm was achieved. Acute recovery and recovery after deployment were reported as a percentage of the initial diameter.

[0291] Example 7

[0292] The in vivo performance of the scaffolds according to the present invention was examined in a porcine cadaver. The study used scaffolds according to the present invention, approximately 7 mm in diameter and having a 32-filament braid (see Table 1, item 1) and delivered via a 7.5F catheter.

[0293] The device was implanted into the turbinate folds of a porcine cadaver. The scaffold was deployed in the porcine nasal cavity in a smooth, controlled manner by withdrawing the outer sheath of the device while keeping the middle pusher in place. Figure 23A 、 Figure 23B 、 Figure 23C and Figure 23D are photos illustrating the deployment process. The delivery catheter had a diameter of approximately 2.8 mm. The scaffold expanded to fill the space between the nasal septum and the turbinate, as Figure 24 shown.

[0294] These deployments identified potential benefits of the scaffolds of the present invention, including: (a) controlled and precise delivery, (b) improved apposition / adaptation to the nasal wall, and (c) reduced profile of the device.

[0295] Example 8

[0296] Human cadaver studies were conducted to evaluate the clinical performance of the stent according to the present invention in human anatomy. Before and after functional endoscopic sinus surgery, the device prototype and delivery system prototype were integrated to test various scenarios within representative anatomy. Endpoints included visual phenomena via endoscopy and clinical feedback.

[0297] Some small-diameter stent prototypes according to the present invention are described in Table 1, while two large-diameter stent prototypes are described in Table 3.

[0298] Table 3

[0299]

[0300] A stent formed using the method along the lines described in Example 1 was placed in the middle meatus, providing mechanical force to center the movement of the middle nasal concha and showing the potential to deliver drugs to the ethmoid sinus. Five deployments were performed: (a) 16 filaments, 8 mm stent, (b) 32 filaments, 8 mm stent, (c) 16 filaments, 10 mm stent, (d) 32 filaments, 10 mm stent, and (d) 32 filaments, 13 mm stent. Although all devices adapted relatively well to the tissue, centered the movement of the middle nasal concha (MT) and provided a lateral outward force on the uncinate process (UP) laterally, the 32-filament, 13-mm stent clearly provided the best fit for the specific space in which it was implanted. Figure 25 is a photograph illustrating the 32-filament, 13-mm stent (length 10 mm) after deployment in the middle meatus of a human cadaver. The implant adapted well to the tissue, with proper medialization of the middle nasal concha.

[0301] The device according to the present invention was also placed in the frontal recess of a human cadaver. In the first cadaver sample, the frontal recess could not be accessed before surgical intervention. The diameter of the frontal sinus ostium was approximately 1 mm and could not accommodate the delivery device. Functional endoscopic sinus surgery (FESS) was performed to remove the ethmoid cells and dilate through the frontal sinus. After this surgery, 32-filament (Table 1, item 6) and 16-filament (Table 1, group 8) implants were deployed into the frontal sinus ostium. Although both devices adapted well to the tissue, the 16-filament device clearly showed enhanced adaptability for the specific space in which the stent was implanted. Figure 26 is a photograph illustrating the 16-filament, 10-mm stent after deployment in the frontal sinus ostium.

[0302] In the second cadaver, the frontal sinus ostium was easily accessible before surgical intervention. 10-mm, 16-filament implants (n = 1, from Table 1, item 8, and n = 1, from Table 1, item 9) were deployed into the frontal sinus before and after FESS, respectively. These implants adapted well to the sinus ostium.

[0303] Sixteen 10-mm diameter strut scaffolds, four 38-mm strut scaffolds, two 38-mm strut scaffolds, and thirty-two 17.5-mm diameter strut scaffolds were placed in the ethmoid sinuses of human cadavers after functional endoscopic sinus surgery. The 10-mm diameter scaffolds were found to be undersized for the particular space in which they were implanted, the 38-mm scaffolds were found to be oversized for the particular space in which they were implanted, and the 17.5-mm scaffolds were found to provide the best fit for the particular space in which they were implanted. Figure 27 Photographs demonstrating the deployment of thirty-two strut scaffolds with a 17.5-mm diameter and 10-mm length in the ethmoid sinuses after FESS are shown for illustration.

[0304] The study used 7.5 French and 9 French catheter systems. The 7.5F system was used to access all frontal sinuses, and the 9F system was used to deploy the device into the ethmoid sinuses. Both catheter diameters were acceptable, and the device functioned appropriately during use. The 90-degree bend was appropriate for reaching the frontal sinuses. This type of catheter is described, for example, in "SINUS SCAFFOLD DELIVERY SYSTEMS" with Attorney Docket No. 81354800002, Serial No. 62 / 186,311, filed Jun. 29, 2015, which is incorporated herein by reference.

[0305] All devices were easily repositionable using standard tools after deployment. All devices were easily removable from the body.

[0306] Example 9

[0307] Uniformly woven PLGA scaffolds (10:90) or PLGA (75:25) scaffolds (diameter = 8 mm, 16 struts, with a braiding angle of 120°) were coated with a carrier coating made of poly(L-lactide-co-ε-caprolactone) (PLCL) cured with hexamethylene diisocyanate (HDI) in the presence of 1-dodecanol (DD) as a chain terminator and zinc octanoate (Zn(Oct)2) as a catalyst. More particularly, four-arm PLCL (40:60), HDI, DD, and Zn(Oct)2 were dissolved in dichloromethane (DCM) to prepare a spray stock solution. Using standard methods, this solution was sprayed onto the woven scaffolds. After drying overnight at room temperature under a nitrogen atmosphere, the scaffolds were fully cured at 60° C. and then cut into 10-mm lengths for radial force and recovery tests. To increase the accumulation of elastomer at the nodes on the scaffolds, anisole (AN) was used as a co-solvent in the spray solution. After the drying and curing treatments described above, these scaffolds were also evaluated for mechanical properties. Figure 28A - 28DOptical micrographs of 8 mm scaffolds of coatings with 16 chains, with and without anisole as a co-solvent, during spraying are as follows: Figure 28A , a PLGA (10:90) scaffold without anisole co-solvent; Figure 28B , a PLGA (10:90) scaffold with anisole co-solvent; Figure 28C , a PLGA (75:25) scaffold without anisole co-solvent; and Figure 28D , a PLGA (75:25) with anisole co-solvent. Some properties of these scaffolds are compiled in Table 4.

[0308] Table 4

[0309]

[0310] After simulated deployment, all scaffolds showed good diameter recovery. However, the scaffolds had significantly different radial stiffnesses, depending on the nodal accumulation of the elastomer. The base braided material did not significantly affect the radial stiffness of the coated scaffolds because both materials had comparable moduli. Similarly, during spraying as described above, 22 mm diameter PLGA (10:90) scaffolds were coated with the same elastomer, in the absence and presence of anisole as a co-solvent. The scaffolds had 32 chains and a braiding angle of 128 or 140. Figure 29A - 29C Optical images of scaffolds coated with and without anisole as a co-solvent during spraying are as follows: Figure 29A , a scaffold coated with 62 wt% elastomer (i.e., the ratio of the mass of the elastomer to that of the base braid) relative to the weight of the base braid from a solution without anisole as a co-solvent; Figure 29B , a scaffold coated with 63 wt% elastomer relative to the weight of the base braid from a solution containing anisole as a co-solvent; and Figure 29C , a scaffold coated with 100 wt% elastomer relative to the weight of the base braid from a solution containing anisole as a co-solvent. As described above, the presence of anisole during scaffold coating improved the nodal accumulation of the resulting elastomer on the scaffold. Additionally, more coating material would result in further nodal accumulation. To evaluate their mechanical properties, these 22 mm diameter scaffolds (length = 10 mm) were subjected to a pressure resistance test between two parallel flat aluminum plates assembled on an INSTRON device. The scaffolds were compressed to 75% of their initial diameter, and the forces during the compression and rebound phases were recorded as a function of the compression distance. Table 2 outlines the pressure (Fc) and rebound force (Fr) at 50% compression. These forces were normalized to the scaffold length.

[0311] Table 5

[0312]

[0313] It is noted that a higher braiding angle provides higher compressive force and resilience to the stent. On the other hand, the accumulation of nodes of the elastomer helps to enhance the stiffness of the stent. However, it has been found that once a certain amount of material has been introduced onto the nodes, further increasing the amount of the coating material only slightly improves the pressure resistance strength of the stent.

[0314] Example 10

[0315] In this example, the stent was further coated with an additional conformal coating comprising a mixture of PLCL and mometasone furoate (MF) as an active agent. The PLCL in the MF-containing coating comprised about 70% (mol%) lactic acid, with the balance being caprolactone (PLCL 70:30). A homogeneous solution of MF and PLCL in ethyl formate and anisole (50:50 v / v) was prepared. Then, the solution was sprayed onto a d = 10 mm stent having 16 strands or a d = 22 mm stent having 32 strands. The amount of MF carried by each stent was controlled by the thickness and loading rate of the MF-containing coating. In the case of the 10 mm stent, drug layers containing 20 wt% MF (80 wt% PLCL) and 40 wt% MF (60 wt% PLCL) were coated onto the stent, yielding 240 μg and 590 μg MF / stent, respectively. In another case, 800 μg MF was coated onto the 22 mm stent with 20 wt% MF (80 wt% PLCL) in the drug layer. The drug layers of these 22 mm and 10 mm stents had comparable thicknesses.

[0316] The in vitro release of MF from these MF-coated stents was determined. At 37 °C, each stent was incubated in a predetermined amount of pH 7.4 PBS buffer containing 2% SDS under gentle shaking. At each specified time point (see Figure 30A and 30B ), the buffer was completely removed for quantification of MF by HPLC, and fresh buffer was added. Figure 30A and 30BThe cumulative absolute mass and mass percentage of MF released from these three groups of scaffolds were separately illustrated. As expected, the amount of MF released daily depends on the total MF loaded into the scaffold. On the other hand, the 10 mm scaffold with a 40 wt% MF loading rate showed a significantly slower release percentage than its analogue with a 20 wt% MF loading rate. This result is different from what the present inventors observed for scaffolds with a relatively low MF loading rate (e.g., 5 wt% to 20 wt%). It was hypothesized that a high loading rate of MF in the drug coating might cause MF crystallization, resulting in slower drug release. In this regard, modulating the drug crystal size is an alternative method to control the drug release profile. Interestingly, the 22 mm and 10 mm scaffolds with a 20 wt% MF loading rate showed substantially the same percentage release profile, indicating that when the drug layer has a similar thickness, the release profile is hardly affected by the scaffold size.

[0317] Example 11

[0318] PLGA (10:90) scaffolds loaded with 590 μg MF and PLGA (75:25) scaffolds loaded with 530 μg MF with a diameter of 10 mm and a length of 6.5 mm were prepared. These scaffolds were sterilized using ethylene oxide and implanted into the left and right maxillary sinus cavities of healthy young New Zealand white rabbits aged 4 - 6 months. The scaffolds were transplanted on days 3, 7, 14, and 28, and the residual drug content was analyzed using HPLC - UV. The kinetic drug release (KDR) curve was determined by gravimetric analysis subtracting the residual drug from the initially loaded drug. The tissues around the scaffolds were collected and analyzed when unfolded to obtain the tissue drug concentration. Figure 31 Examples illustrate the in vivo KDR curves of MF - coated PLGA (10:90) and PLGA (75:25) scaffolds. Figure 32 Shows the concentration of MF in the sinus mucosa of rabbits sacrificed at a given time point. Figure 33 Shows the total amount of MF on the scaffold plus the drug content in the sinus mucosa of rabbits sacrificed at a given time point.

[0319] Example 12

[0320] As described above, a braided PLGA stent with a diameter of 17.5 mm (PLGA 10:90, 32 strands) was coated with a carrier coating, which was made of poly(L-lactide-co-ε-caprolactone) (PLCL) (especially L-PLCL (40:60)) cured with hexamethylene diisocyanate (HDI) in the presence of 1-dodecanol (DD) as a chain terminator and optionally using a Zn(Oct)2 catalyst. Then, a further therapeutically active agent-containing layer comprising 30 wt% MF and 70 wt% PLCL in a homogeneous solution of MF and PLCL prepared as described above in ethyl formate and anisole (70:30 v / v) was further coated on the stent, except that D,L-PLCL (80:20) or D,L-PLCL (90:10) was used as the carrier polymer instead of L-PLCL (70:30) as described in Example 10 above.

[0321] As described in Example 10 above, the in vitro release of MF from these MF-coated stents was further determined. As Figure 34 shown, the MF release rate associated with D,L-PLCL (80:20) (Tg = 20 °C) was much faster than that of D,L-PLCL (90:10) (Tg = 35 °C). Without wishing to be bound by theory, it is believed that the glass transition temperature (Tg) of the carrier polymer plays an important role in determining the drug release profile. In this regard, in the absence of the Tg effect, a copolymer with a minimum amount of the more hydrophobic monomer (caprolactone), i.e., D,L-PLCL (90:10), would generally be expected to exhibit a faster release.

[0322] Example 13

[0323] As described above, a braided PLGA stent with a diameter of 17.5 mm (PLGA 10:90, 32 strands) was coated with a carrier coating, which was made of poly(L-lactide-co-ε-caprolactone) (PLCL), especially L-PLCL (40:60), cured with hexamethylene diisocyanate (HDI) in the presence of 1-dodecanol (DD) as a chain terminator and an optionally used catalyst. Then, another therapeutic agent-containing layer containing 30 wt% MF and 70 wt% polymer material in a homogeneous solution of MF and the polymer material prepared as described above in ethyl formate and anisole (70:30 v / v) was further coated on the stent, except that in addition to L-PLCL (70:30) described in Example 10 above, the tested polymer materials further included an admixture of PLCL (70:30) and PLGA (75:25) in a 75:25 wt / wt ratio, an admixture of PLCL (70:30) and PLGA (85:15) in a 75:25 wt / wt ratio, and an admixture of PLCL (70:30) and PLA in a 75:25 wt / wt ratio.

[0324] As described in Example 10 above, the in vitro release of MF from these MF-coated stents was further determined. As Figure 35 shown, when copolymers with a higher lactide content were mixed into PLCL (70:30), the MF release rate decreased. Without wishing to be bound by theory, it is believed that the MF release rate decreases as the Tg of the polymer mixed with PLCL (70:30) increases. In this regard, among the three polymers used (PLGA (75:25) Tg ~ 50 °C, PLGA (85:15) Tg ~ 55 °C, and PLA Tg ~ 60 °C), PLA has the highest glass transition temperature.

[0325] Example 14

[0326] As described above, a uniformly braided PLGA (10:90) stent (diameter = 17.5 mm, length = 10 mm, 32 strands, with a braiding angle of 90° or 128°) was coated with a carrier coating, which was made of poly(L-lactide-co-ε-caprolactone) (PLCL) cured with hexamethylene diisocyanate (HDI) in the presence of 1-dodecanol (DD) as a chain terminator as described above, and as described in Example 10 before, a conformal coating containing a mixture of PLCL and mometasone furoate was further coated.

[0327] To evaluate their mechanical properties, as Figure 38As schematically shown, the stents 100 are laid on their sides and compressed between two parallel plates 210a, 210b of a compression device 200 in a chamber maintained at 34°C and 80% relative humidity. In other words, the stents 100 are placed in a compressed state between the two parallel plates 210a, 210b such that the axis A of the tubular stent is parallel to the parallel plates and such that the tubular stent 100 is compressed between the parallel plates until a point where the distance d between the parallel plates is a percentage of the initial unconstrained diameter such that when the stent 100 is compressed in the compression device 200, the tubular stent has a first minimum width D measured perpendicular to the axis A equal to the distance d. The stents are compressed to 1.5 mm (8.6% of their initial diameter) or 3 mm (17.1% of their initial diameter). The stents are removed from the compression device on a weekly basis and the recovery minimum width D (also referred to as the second minimum width D) of each stent is measured immediately after removal from the compression device and six hours after removal of the stent from the compression device. The recovery % is calculated by dividing the second minimum width D immediately after or six hours after removal from the compression plate by the first minimum width D (which is equal to the gap distance d between the parallel plates, i.e., D(mm) / 1.5 mm × 100 or D(mm) / 3 mm × 100).

[0328] In certain embodiments, after maintaining the compressed state for 10 weeks at a distance d of 8.5% of the stent's fabricated diameter (e.g., a 17.5 mm stent compressed to 1.5 mm), and after a period of six hours after removing the tubular stent 100 from the compressed state, the first minimum width D (distance d) of the tubular stent can recover to a second minimum width D measured perpendicular to the axis, which is at least 450% (e.g., 450% to 1000%) of the first minimum width D (theoretical maximum 1166%). In certain embodiments, after maintaining the compressed state for 10 weeks at a distance d of 17% of the stent's fabricated diameter (e.g., a 17.5 mm stent compressed to 3.0 mm), and after a period of six hours after removing the tubular stent 100 from the compressed state, the first minimum width D of the tubular stent can recover to a second minimum width D measured perpendicular to the axis, which is at least 250% (e.g., 250% to 500%) of the first minimum width D (theoretical maximum 583%).

[0329] The results immediately after removal of the 90° and 128° braided angle stents compressed to 1.5 mm are presented in Figure 36A and the results immediately after removal of the 90° and 128° braided angle stents compressed to 3.0 mm are presented in Figure 36B In. As can be seen from Figure 36AAs can be seen, after 1 week, the 90° braided angle stent immediately recovered to approximately 230% from the 1.5 mm compressed state and to approximately 250% after 10 weeks. After 1 week, the 128° braided angle stent immediately recovered to approximately 175% from the 1.5 mm compressed state and to approximately 190% after 10 weeks. As can be seen from Figure 36B As can be seen, after 1 week, the 90° braided angle stent immediately recovered to approximately 165% from the 3.0 mm compressed state and to approximately 170% after 10 weeks. After 1 week, the 128° braided angle stent immediately recovered to approximately 140% from the 3.0 mm compressed state and to approximately 175% after 10 weeks.

[0330] The results of the 90° and 128° braided angle stents compressed to 1.5 mm after 6 hours of removal are presented in Figure 37A and Table 6, and the results of the 90° and 128° braided angle stents compressed to 3.0 mm after 6 hours of removal are presented in Figure 37B and Table 7. After 1 week, the 90° braided angle stent recovered to approximately 540% from the 6-hour 1.5 mm compressed state and to approximately 510% after 10 weeks. After 1 week, the 128° braided angle stent recovered to approximately 550% from the 6-hour 1.5 mm compressed state and to approximately 480% after 10 weeks. After 1 week, the 90° braided angle stent recovered to approximately 280% from the 6-hour 3.0 mm compressed state and to approximately 270% after 10 weeks. After 1 week, the 128° braided angle stent recovered to approximately 300% from the 6-hour 3.0 mm compressed state and to approximately 190% after 10 weeks.

[0331] Table 6

[0332]

[0333] Table 7

[0334]

[0335] As can be seen from Figure 37A As can be seen, at all time points, the recovery of the 90° braided angle stent from the 1.5 mm compressed state was similar to that of the 128° braided angle stent from the 1.5 mm compressed state, each showing a significant decline in recovery between 1 and 2 weeks. Representative recovery data are provided in Table 6.

[0336] Similarly, as can be seen from Figure 37B As can be seen, until week 7, the recovery of the 90° braided angle stent from the 3.0 mm compressed state was similar to that of the 128° braided angle stent from the 3.0 mm compressed state, each showing a significant decline in recovery between 1 and 2 weeks. Representative recovery data are provided in Table 7.

[0337] Although various embodiments have been specifically illustrated and described herein, it should be understood that modifications and variations of the present invention are encompassed by the above teachings and within the scope of the appended claims without departing from the spirit and intended scope of the present invention. For example, although a stent for sinus applications has been described herein, such a stent can also be used for other applications, such as eustachian tube stent implantation.

Claims

1. A method of loading an implantable stent, comprising: A) Providing a) an implantable stent having a diameter in a prepared state of 10 mm to 15 mm, comprising i) a plurality of polymeric chains comprising a first polymeric material; ii) a coating on the polymeric chains comprising an elastomer; iii) a drug-containing layer comprising mometasone furoate and a second polymeric material, the second polymeric material may be the same as or different from the first polymeric material, and iv) a surface layer on top of the drug-containing layer, wherein the surface layer causes a more linear release of the mometasone furoate, and wherein the implantable stent is adapted to conform to the shape of the middle nasal meatus space; and b) a delivery catheter configured to maintain the stent in a radially constrained shape; and B) Loading the implantable stent into the delivery catheter such that the stent is maintained in a radially constrained shape.

2. The method according to claim 1, wherein the first polymeric material comprises poly(lactide-co-glycolide).

3. The method according to any one of claims 1-2, wherein the stent comprises between 8 and 64 polymeric chains.

4. The method according to claim 1, wherein the plurality of polymeric chains are braided.

5. The method according to claim 4, wherein the braided polymeric chains cross each other at a braiding angle of 80° to 140°.

6. The method according to claim 1, wherein the elastomer comprises poly(lactide-co-caprolactone).

7. The method according to claim 6, wherein the elastomer comprises poly(lactide-co-caprolactone) having a molar percentage range of 30 to 50 of lactide and a molar percentage range of 50 to 70 of caprolactone.

8. The method according to claim 1, wherein the elastomer is a crosslinked elastomer.

9. The method according to claim 8, wherein the crosslinked elastomer comprises an ethyl carbamate and / or urea crosslinking agent.

10. The method according to claim 1, wherein the first polymeric material is biodegradable.

11. The method according to claim 1, wherein the stent further comprises a surface layer.

12. The method according to claim 1, wherein the drug-containing layer has a thickness of 5 to 20 μm.

13. The method according to claim 1, wherein the surface layer has a thickness of about 1 μm.

14. A delivery system comprising: (a) an implantable stent having a diameter in a prepared state of 10 mm to 15 mm, comprising i) a plurality of polymeric chains comprising a first polymeric material; ii) a coating on the polymeric chains comprising an elastomer; iii) a drug-containing layer comprising mometasone furoate and a second polymeric material, the second polymeric material may be the same as or different from the first polymeric material, and iv) a surface layer on top of the drug-containing layer, wherein the surface layer causes a more linear release of the mometasone furoate, and wherein the implantable stent is adapted to conform to the shape of the middle nasal meatus space; and b) a delivery catheter, wherein the stent is placed in the delivery catheter in a radially constrained shape.

15. The delivery system according to claim 14, wherein the stent comprises between 8 and 64 polymeric chains.

16. The delivery system according to claim 14, wherein the plurality of polymeric chains are braided.

17. The delivery system according to claim 16, wherein the braided polymeric chains cross each other at a braiding angle of 80° to 140°.

18. The delivery system according to claim 14, wherein the elastomer comprises poly(lactide-co-caprolactone).

19. The delivery system according to claim 18, wherein the elastomer comprises poly(lactide-co-caprolactone) having a lactide molar percentage range of 30 to 50 and a caprolactone molar percentage range of 50 to 70.

20. The delivery system according to claim 14, wherein the elastomer is a crosslinked elastomer.

21. The delivery system according to claim 20, wherein the crosslinked elastomer comprises an urethane and / or urea crosslinker.

22. The delivery system according to claim 14, wherein the first polymeric material is biodegradable.

23. The delivery system according to claim 14, wherein the stent further comprises a surface layer.

24. The delivery system according to claim 14, wherein the first polymeric material comprises poly(lactide-co-glycolide).

25. The delivery system according to claim 14, wherein the drug-containing layer has a thickness of 5 to 20 μm.

26. The delivery system according to claim 14, wherein the surface layer has a thickness of about 1 μm.

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