Implantable drug depots have been applied for decades across a range of sites in the body, including the brain. For tubular structures in the body, coated stents have been applied to provide local high concentrations of a therapeutic, as found in drug eluting stents. In the gastrointestinal (“GI”) tract, coated stents have been explored, though suffer from a significant rate of complications including stent migration and tissue perforation. Moreover, the delivery of therapeutics from drug eluting stents is governed by diffusion limitations through tissue, potentially limiting delivery to therapeutics of lower molecular weight and particular physico-chemical characteristics which support partitioning of the drug into the mucosa.
In the GI tract, endoscopic injection, initially pioneered through the development of the Carr-Locke Needle, transformed the capacity to locally deliver therapeutics for a range of applications including hemostasis with epinephrine, sclerosant injection for variceal ablation, submucosal lifts with normal saline and other materials, as well as steroid injections for inflammation control, and injection of biologics for inflammatory stricture management. All of these applications apply a hypodermic needle, which can be deployed endoscopically supporting single site injection.
Recognizing that many GI pathologies, including inflammatory bowel disease, eosinophilic GI disorders, and Celiac disease, affect extended multi-centimeter segments of the GI tract, the present disclosure provides a solution for rapid circumferential submucosal deposition of controlled drug releasing systems.
Implantable drug depots have the capacity to locally meet therapeutic requirements by maximizing local drug efficacy and minimize potential systemic side effects. The GI tract represents a site with a broad range of pathology affecting its tubular structure. Its length and tubular structure though make the application and deposition of drug depots challenging as current injectable systems, as briefly described above, generally only facilitate single point administration.
According to aspects of the present disclosure, a kirigami-mediated injectable stent system is provided. The systems and methods described herein enable radial/circumferential and longitudinal intramucosal delivery for an extended release of therapeutics within tubular structures of the body. According to some aspects, a kirigami-based injectable stent system is provided that can enable ultra-long local drug release through deposition of drug-loaded polymeric particles in the tubular mucosa of the GI tract.
According to some aspects of the present disclosure, a stent for treating tissue within a gastrointestinal tract or trachea of a subject is provided. The stent includes a tubular body extending along a central axis and configured to move between a retracted position and an elongated position, and a plurality of projections formed into the tubular body, each projection configured to form a cutting edge to pierce a submucosal tissue within the gastrointestinal tract or trachea. Each projection among the plurality of projections is configured to undergo a change in orientation relative to the central axis when the tubular body moves between the retracted position and the elongated position.
According to some aspects of the present disclosure, a stent system for treating a tissue within a gastrointestinal tract or trachea of a subject is provided. The system includes a tubular body extending along a central axis to form a lumen within the tubular body an actuator received within the lumen and configured to move the tubular body between a retracted position and an elongated position, and a pattern of a plurality of cuts formed along the tubular body and extending through the tubular body to the lumen. The pattern of the plurality of cuts deploys into a plurality of interconnected projections that are configured to extend radially away from the tubular body relative to the central axis to engage a submucosal tissue within the gastrointestinal tract or trachea of a subject when the tubular body is moved towards the elongated position.
According to some aspects of the present disclosure, a method of inserting a stent into a gastrointestinal tract or trachea a subject is provided. The method includes positioning a stent to a target tissue site within a gastrointestinal tract or trachea, the stent having a tubular body extending along a central axis to form a lumen within the tubular body, and pressurizing an actuator received within the lumen to move the tubular body from a retracted position to an elongated position. A surface of the tubular body includes a pattern of a plurality of cuts configured to deploy into a plurality of interconnected projections as the tubular body is moved into the elongated position to engage the target tissue site of the subject.
The foregoing and other aspects and advantages of the disclosure will appear from the following description. In the description, reference is made to the accompanying drawings which form a part hereof, and in which there is shown by way of illustration a preferred configuration of the disclosure. Such configuration does not necessarily represent the full scope of the disclosure, however, and reference is made therefore to the claims and herein for interpreting the scope of the disclosure.
The invention will be better understood and features, aspects and advantages other than those set forth above will become apparent when consideration is given to the following detailed description thereof. Such detailed description makes reference to the following drawings.
Before any aspects of the invention are explained in detail, it is to be understood that the invention is not limited in its application to the details of construction and the arrangement of components set forth in the following description or illustrated in the following drawings. The invention is capable of other aspects and of being practiced or of being carried out in various ways. Also, it is to be understood that the phraseology and terminology used herein is for the purpose of description and should not be regarded as limiting. The use of “including,” “comprising,” or “having” and variations thereof herein is meant to encompass the items listed thereafter and equivalents thereof as well as additional items. Unless specified or limited otherwise, the terms “mounted,” “connected,” “supported,” and “coupled” and variations thereof are used broadly and encompass both direct and indirect mountings, connections, supports, and couplings. Further, “connected” and “coupled” are not restricted to physical or mechanical connections or couplings.
The use herein of the term “axial” and variations thereof refers to a direction that extends generally along an axis of symmetry, a central axis, an axis of rotation, or an elongate direction of a particular component or system. For example, axially extending features of a component may be features that extend generally along a direction that is parallel to an axis of symmetry or an elongate direction of that component. Further, for example, axially aligned components may be configured so that their axes of rotation are aligned. Similarly, the use herein of the term “radial” and variations thereof refers to directions that are generally perpendicular to a corresponding axial direction. For example, a radially extending structure of a component may generally extend at least partly along a direction that is perpendicular to a longitudinal or central axis of that component. The use herein of the term “circumferential” and variations thereof refers to a direction that extends generally around a circumference of an object or around an axis of symmetry, an axis of rotation, a central axis, or an elongate direction of a particular component or system.
As also used herein, unless specified or limited otherwise, the terms “approximately” and “substantially” and variations thereof, when used relative to a numerical value, define a range of values within 20% of the numerical value (e.g., within 15%, 10%, or within 5%).
In some implementations, devices or systems disclosed herein can be utilized, manufactured, or treated using methods embodying aspects of the invention. Correspondingly, any description herein of particular features, capabilities, or intended purposes of a device or system is generally intended to include disclosure of a method of using such devices for the intended purposes, of a method of otherwise implementing such capabilities, of a method of manufacturing relevant components of such a device or system (or the device or system as a whole), and of a method of installing or utilizing disclosed (or otherwise known) components to support such purposes or capabilities. Similarly, unless otherwise indicated or limited, discussion herein of any method of manufacturing or using for a particular device or system, including installing the device or system, is intended to inherently include disclosure, as embodiments of the invention, of the utilized features and implemented capabilities of such device or system.
The following discussion is presented to enable a person skilled in the art to make and use embodiments of the invention. Various modifications to the illustrated embodiments will be readily apparent to those skilled in the art, and the generic principles herein can be applied to other embodiments and applications without departing from embodiments of the invention. Thus, embodiments of the invention are not intended to be limited to embodiments shown, but are to be accorded the widest scope consistent with the principles and features disclosed herein. The following detailed description is to be read with reference to the figures, in which like elements in different figures have like reference numerals. The figures, which are not necessarily to scale, depict selected embodiments and are not intended to limit the scope of embodiments of the invention. Skilled artisans will recognize the examples provided herein have many useful alternatives and fall within the scope of embodiments of the invention.
Kirigami is a Japanese form of paper art similar to origami that includes cutting of the paper and can enable the design of a range of functional tools and programmable systems from macroscale soft actuators and robots to microelectronics and nanostructures. Buckling-induced kirigami structures are engineered to utilize local elastic instabilities for versatile shape transformation from flat, generally smooth surfaces to complex three-dimensional architectures. According to some applications, the buckling kirigami metasurfaces have been applied to footwear outsoles to generate higher friction forces and mitigate the risk of slips and falls in a range of environments.
As explained herein, inspired from the skin of scaly-skin animals like snakes and sharks, an injectable stent was developed which is composed of a periodic array of denticle-like needles (e.g., a kirigami cylindrical shell) integrated with a linear actuator (e.g., a pneumatic soft actuator). As detailed herein, a combination of finite element (“FE”) simulations and experiments, kirigami shells and linear actuators were identified to develop injectable stents in multiple length scales that can be easily deployed in the tubular lumen of the GI tract such as esophagus as well as arteries and airways. By pressurizing the soft actuator, the kirigami needles buckle out (e.g., extend) such that the resulting needles provide required stiffness and radial expansion (in some examples, up to 60% of the stent diameter) to enable injections of drug-loaded particles into the tissue of a subject (e.g., into submucosal tissues of the GI tract). These kirigami-based injectable stents serve as a class of drug-eluting stents, capable of releasing drug depots through multi-point deposition of drug particles, thereby enhancing sustained local delivery of therapeutics.
Referring to
The tubular body 12 can include a cylindrical outer shell 16 forming a lumen 17 (e.g., a hollow core) and an actuator 18 arranged within the lumen 17 of the outer shell 16. The outer shell 16 can include at least one cut 20. In the illustrated non-limiting example, the outer shell 16 can include a patterned array of a plurality of interconnected cuts 20 (e.g., openings). In the illustrated non-limiting example, the plurality of cuts 20 extend along at least a portion of the axial length of the tubular body 12. For example, the plurality of cuts 20 can extend along at least 50% of an entire length L0 of the tubular body 12. According to some non-limiting examples, the plurality of cuts 20 can extend along between about 50% and about 100% of the entire length L0 of the tubular body 12. According to the illustrated non-limiting example, the plurality of cuts 20 can extend along between about 80% and about 95% of the entire length L0 of the tubular body 12. In the illustrated non-limiting example, the plurality of cuts 20 extend along at least a portion of the circumference of the tubular body 12. For example, the plurality of cuts 20 can extend along at least 50% of the circumference of the tubular body 12. According to some non-limiting examples, the plurality of cuts 20 can extend along between about 50% and about 100% of the circumference of the tubular body 12. According to the illustrated non-limiting example, the plurality of cuts 20 can extend along between about 90% and about 100% of the circumference of the tubular body 12.
The length L0 of the tubular body 12 can be defined as an initial length between a first end 21 and an opposing send end 23 of the tubular body 12 when the tubular body 12 is in the retracted position (
The tubular body 12 can also define a nominal outer diameter D, defined as an initial diameter of the outer shell 16 when the tubular body 12 is in the retracted position (
When the tubular body 12 is elongated from the retracted position to the extended position, the tubular body 12 can define an elongated length LE (
The plurality of cuts 20 can be configured to form a kirigami-inspired pattern configured to undergo a shape change when stress is axially applied along the outer shell 16. via the actuator 18. The at least one cut 20 can form at least one projection element 22. In the illustrated non-limiting example, the series of patterned cuts 20 can form a plurality of projection elements 22 (e.g., needles). When the tubular body 12 of the stent 10 is in a retracted position (
For example, as will be described, the outer shell 16 can be configured to automatically respond to strain applied in a direction along the central axis 14. That is, the series of patterned cuts 20 form a surface on the outer shell 16 that buckles in response to applied axial strain to form a plurality of projection elements from that cut surface. In the illustrated non-limiting examples, the actuator 18 is configured to apply the axial strain, and that axial strain results in stress within the outer shell 16 that causes the projection elements 22 to extend outwards from an orientation in which the projection elements form a substantially uniform (e.g., flat) cylindrical surface, into an orientation in which the projection elements deploy radially outwards relative to the central axis 14. According to some non-limiting examples, the magnitude of applied axial strain to the outer shell 16 can correspond to a magnitude of radial extension of the projection elements 22. That is, owing to the pattern of cuts 20 formed in the outer shell 16, a surface is provided that transforms in a radial direction in response to strain applied in an axial direction.
Referring now to
With particular reference to
Referring now to
The patterned cuts 20 forming the projection elements 22 can be characterized by a needle length l, hinge length δ, and cut angle γ. The needle length l can be described as a characteristic length of the patterned cut 20 and can be considered as a length of the needle formed by the projection element 22. The needle length l can be defined by a distance between the needle tip 30 of the projection element 22 and either one of a first distal end 36 of the first edge 32 or a second distal end 38 of the second edge 34 (i.e., distal ends of the cut 20). According to some non-limiting examples, the projection elements 22 can define a needle length l between about 0.1 mm and about 60 mm. According to other non-limiting examples, the projection elements 22 can define a needle length l between about 1 mm and about 30 mm. According to yet further non-limiting examples, the projection elements 22 can define a needle length l between about 1 mm and about 15 mm. According to the illustrated non-limiting example, the projection elements 22 define a needle length l of about 10 mm.
The hinge length δ can be described as the width of ligaments forming an interstitial spacing separating adjacent cuts 20. The hinge length δ can be defined by a distance between the needle tip 30 of a first projection element 22a and either one of the first distal end 36 or the second distal end 38 of a second, adjacent projection element 22b. According to some non-limiting examples, the cuts 20 can define a hinge length δ between about 0.1 mm and about 10 mm. According to other non-limiting examples, the cuts 20 can define a hinge length δ between about 0.1 mm and about 5 mm. According to yet further non-limiting examples, the cuts 20 can define a hinge length δ between about 0.1 mm and about 2 mm.
The cut angle γ can be described as the angle of the cut 20 forming either one of the first and second edges 32, 34 of the projection element 22 relative to a plane 25 intersecting and orthogonal to the central axis 14. According to some non-limiting examples, the cuts 20 can define a cut angle γ between about 0 degrees and about 90 degrees. According to other non-limiting examples, the cuts 20 can define a cut angle γ between about 5 degrees and about 45 degrees. According to yet further non-limiting examples, the cuts 20 can define a cut angle γ between about 10 degrees and about 45 degrees. According to the illustrated non-limiting example, the cuts 20 define a cut angle γ of about 30 degrees.
Referring still to
The cuts 20 forming the projection elements 22 can be evenly (e.g., periodically) circumferentially spaced around the outer shell 16 (see, e.g.,
The outer shell 16 of the tubular body 12 of the stent 10 can be formed from a thin sheet of material. According to some non-limiting examples, the outer shell 16 is formed of an elastomeric material (e.g., plastic, a polyester plastic, etc.). According to other non-limiting examples, the outer shell 16 can be formed of a metal, a polymer, or a composite. In some non-limiting examples, the outer shell 16 can be formed of rigid, thin sheets of steel, nitinol, or plastic and the “elasticity” of the material can be provided by the pattern of cuts 20. In other non-limiting examples, the outer shell 16 can be formed of soft flexible materials such as rubbers. In yet further non-limiting examples, the outer shell 16 can be formed of soluble polymers. The material of the outer shell 16 can have a shape memory, thereby allowing the projection elements 22 of the outer shell 16 to repeatedly transition between the deformed and undeformed states. According to some non-limiting examples, the outer shell 16 can define a wall thickness between about 0.01 mm and about 2 mm. According to other non-limiting examples, the wall thickness can be between about 0.05 mm and about 1 mm. According to yet further non-limiting examples, the wall thickness can be between about 0.05 mm and about 0.5 mm. According to the illustrated non-limiting example, wall thickness is about 0.13 mm.
As previously described herein, the outer shell 16 of the tubular body 12 can define a lumen (e.g., a hollow core) configured to receive an actuator 18.
The body 50 of the actuator 18 can define a hollow tube including an interior cavity 56. According to some non-limiting examples, the body 50 can define a wall thickness between about 0.01 mm and about 5 mm. According to other non-limiting examples, the wall thickness can be between about 0.05 mm and about 3 mm. According to yet further non-limiting examples, the wall thickness can be between about 0.05 mm and about 2 mm. According to the illustrated non-limiting example, wall thickness is about 1.5 mm.
The interior cavity 56 can extend through the body 50 between the first actuator end 52 and the second actuator end 54. In the illustrated non-limiting example, the interior cavity 56 forms a first opening 58 at the first actuator end 52 and a second opening 60 at the second actuator end 54. The actuator can also include a plug 62 and a cap 64. The plug 62 can be coupled at the second actuator end 54 of the actuator 18 to enclose the second opening 60. The plug 62 includes a plug boss 66 and a plug flange 68 at a distal end thereof extending radially outward from the plug boss 66. The plug boss can be configured to be received within the interior cavity 56 of the body 50. The plug flange 68 can be configured to abut the second actuator end 54 of the body 50, when the actuator 18 is in an assembled state (see, e.g.,
The cap 64 can be coupled at the first actuator end 52 of the actuator 18 to enclose the first opening 58. The body 50, plug 62, and cap 64 together define and enclose the interior cavity 56. The cap 64 can include a cap boss 70 and a cap flange 72 at a distal end thereof and extending radially outward from the cap boss 70. The cap boss 70 can be configured to be received within the first opening 58. The cap flange 72 can be configured to abut the first actuator end 52 of the body 50, when the actuator 18 is in the assembled state, to form a fluid impervious seal with the body 50. According to the illustrated non-limiting example, the cap 64 can be formed of an elastomeric material or a hard material (e.g., a plastic). According to some non-limiting examples, the cap 64 can include a nylon plastic quick-turn plug.
The cap 64 can include an inlet port 74 and a fluid passage 76 in fluid communication with the inlet port 74. The fluid passage 76 is configured to provide fluid communication between the inlet port 74 and the interior cavity of the actuator 18. The inlet port 74 can extend axially outward from the first end 21 of the outer shell 16 of the stent 10 (see
Referring now to
As best illustrated in
The projection elements 122 illustrated in
As illustrated in
As best illustrated in
Referring now to
According to one non-limiting example, the therapeutic agent can include an anti-inflammatory drug (e.g., budesonide, prednisone, colchicine, resveratol, etc.), and anti-proliferative drugs (e.g., paclitaxel, everolimus, sirolimus, among other -limus agents, etc.), for delivery to walls of the GI tract or trachea. Budesonide, for example, is an anti-inflammatory drug commonly used to treat inflammatory bowel disease and eosinophilic GI disorders. In the illustrated embodiment, budesonide can be encapsulated into poly lactic-co-glycolic acid (“PLGA”) microparticles using a continuous microfluidic droplet generation method (generally illustrated in
In the above description, reference is made to various dimensions, parameters, and characteristics of the stent 10 and its components. It is to be understood that these components can be sized based on the intended application. For example, within the GI tract, stents 10 can be configured for placement within the stomach, esophagus, colon, small intestine, or large intestine. Dimensions and parameters of the stents 10 can be chosen based on the application or dimensions of the tubular structures of the GI tract or trachea for a given subject. For example, depending on the target position of deployment of the stent, a desired diameter and length of the stent may be determined (i.e., based on a diameter and length of the target position). Based on a determined diameter and length of the stent, the pattern of cuts 20 (e.g., needle, length, cut angle, hinge length, etc.) can be determined such that the resulting kirigami stent 10 expands to reach a desired penetration depth. For example, hinge length can be determined or calculated based on needle length, cut angle, thickness, and/or material of the outer shell 16 to provide the pop-up deployment motion of the projection elements 22.
Referring now to
The method can begin at 302 by inserting the stent 10 into a tubular tissue structure of a subject in a first, insertion direction (e.g., relative to the central axis 14). For example, the stent 10 can be inserted into the GI tract (
Once the stent 10 is positioned at the tissue site of interest, the actuator 18 can be actuated 304 from the retracted position towards the extended position, thereby deploying the projection elements 22 radially outward into the deformed state. For example, the actuator 18 can be pressurized by the pressurized fluid source 75 coupled to the inlet port 74 and the actuator 18 can begin to elongate to engage the enclosed first and second ends 21, 23 of the outer shell 16 of the tubular body 12, thereby elongating the outer shell 16 and deforming the projection elements 22 to deploy radially outwards.
With the stent 10 in the extended position, the projection elements 22 can engage 306 the tissue of the subject to form a pattern of circumferential injection sites into the tissue. According to some non-limiting examples, the stent 10 can be moved in a second, removal direction by applying a pulling force to the first end of the tubular body 12 of the stent 10. By moving the stent 10 in the second direction with the projection elements 22 deployed, the projection elements can be further driven into the tissue of the subject to increase the insertion depth of the projection elements 22. For example, the projection elements 22, when deployed, generally extend from the second end 23 towards the first end 21 of the tubular body 12, owing to the needle angle θ (see, e.g.,
As previously described, the projection elements 22 can be loaded with a therapeutic agent (see, e.g.,
For removal of the stent 10, the stent 10 can be moved in the first direction (towards the second end 23) to remove the projection elements 22 from the tissue of the subject. With the projection elements 22 removed, the stent 10 can be actuated from the extended position towards the retracted position to stow the projection elements into the undeformed state. Once the stent 10 is in the retracted position, the stent 10 can be removed from the subject by moving the stent 10 in the second, removal direction, for example, by again applying a pulling force to the first end of the tubular body 12 of the stent 10.
Referring now to
According to some non-limiting examples, the mold 410 can be sprayed with a releasing agent for easy demolding. Then, the elastomeric actuator body 50 and plug 62 can be cast separately using an elastomeric material (e.g., a silicone-base rubber, vinylpolysiloxane, a-silicone). According to some non-limiting examples, the elastomeric material can be a duplicating elastomer (e.g., Elite Double 8). The casted mixture can be mixed for a predetermined period of time (e.g., two minutes), placed in a vacuum for degassing, and then allowed to set at a predetermined temperature (e.g., room temperature) for a predetermined period of time (e.g., thirty minutes) to cure.
With the body 50 formed, strands of fiber reinforcement material can be wrapped 404, 406 within the helical recesses 420 along the body 50 (
Referring now to
According to some non-limiting examples, the outer shell 216 can include small apertures 292 perforated along lateral edges of the outer shell 216, which can be used to facilitate alignment when formed into a cylindrical shape. According to the illustrated non-limiting example, circular cutouts 294 can be coupled to the first and second ends 221, 223 of the outer shell 216. The circular cutouts 294 can be configured as end caps for the outer shell 216 when formed into a cylindrical shape. In the illustrated non-limiting example, the circular cutouts 294 can include one or more tabs 296 extending outward from the circular cutouts 294. The tabs 296 can be configured to be coupled to the first and second ends 221, 223 of the outer shell 216 (e.g., via an adhesive) to secure the circular cutouts 294 to the outer shell 216. The circular cutout 294 arranged at the first end 221 of the outer shell 216 can include a central aperture 298. The central aperture 298 can be configured to receive the inlet port 274 (see
As illustrated in
According to some non-limiting example, a surface coating can include a radiopaque coating. For example, at least a portion of the outer shell 216 can be coated in a radiopaque coating. The radiopaque coating can make the outer shell 216 of the stent 200 radiopaque. According to some non-limiting examples, the entire outer shell 216 can be coated with the radiopaque coating. According to other non-limiting examples, at least the projection elements 222 can be coated with the radiopaque coating. According to some non-limiting examples, the outer shell 216 can be coated with a thin layer of tungsten filled conductive ink (e.g., RO-948 Radio Opaque Ink, MICROCHEM).
Finally, as illustrated in
The following description includes particular non-limiting examples of stents that utilize the systems and methods previously described herein. The following examples are not intended to limit the disclosure. In the following description, a systematic study is described, in which the properties of the stents described herein are characterized using various parameters and tests.
As described herein, the stents (e.g., stents 10, 100, 200) can include a cylindrical kirigami skin that includes a periodic array of snake denticle-like cuts, which can be embedded in thin plastic sheets. According to some non-limiting examples, color-coded polyester plastic shim stocks can be used to fabricate the kirigami surfaces with snake skin-like needles. To measure the material properties of the shim stocks, uniaxial tensile tests were carried out on the 80 mm×43 mm plastic specimens with a range of thicknesses, t=0.05, 0.08, 0.10, 0.13, 0.19 mm, according to ASTM D882-18 (Standard Test for Tensile Properties of Thin Plastic Sheeting). A uniaxial testing machine (e.g., an Instron 5942 series Universal Testing System) with a 500 N load cell was used to test specimens. All the tests were conducted under uniaxial tensile loading by applying a constant displacement rate of 0.5 mm/s quasi-statically until the 500 N load cell threshold. The response is characterized by linear elastic region followed by a plateau. Nominal stress-strain curves can be seen in
The stents can also include a pneumatic fiber-reinforced soft actuator made of a 1.5 mm thick silicone-based rubber. The silicone-based rubber can be Vinylpolysiloxane (a-silicone) duplicating elastomer (e.g., “Elite Double 8”) was used to cast the soft actuator. To measure the material properties, three dog-bone specimens (
The pneumatic fiber-reinforced soft actuator can provide a linear motion to induce tensile strain in the kirigami skin and trigger the needles to pop out. The radial expansion (εr) and axial extension (εa) of the stent and popping angle of the needles (θ) can be tuned by controlling the actuator pressure (P/P0), where P0=1 atm.
For example,
Numerical models of the kirigami stents can be constructed with different combinations of t and l, and non-linear finite elements (FE) analyses can be employed to capture the deformation of the stents subjected to the applied actuator pressure using a FE package such as ABAQUS/Explicit. All the simulations were carried out using the commercial Finite Element (FE) package ABAQUS 2017. The Abaqus/Explicit solver was employed for the simulations. FE models were constructed of the elastomer actuator, Kevlar fiber, nylon plastic plug, and kirigami plastic shell to investigate the deformation response of the kirigami stent.
A linear elastic material model was used for Kevlar fiber, polyester plastic, and nylon plastic. Kevlar fiber has a density of 1.13E3 kg/m3, Young's modulus of 31067 MPa, and Poisson's ratio of 0.36 with a circular beam section of 0.0889 mm radius. Polyester plastic sheet has a density of 1.13E3 kg/m3, Young's modulus of 3655 MPa, Poisson's ratio of 0.4 with shell section of 0.127 mm thickness. The nylon Plastic has a density of 1.15E3 kg/m3, Young's modulus of 4000 MPa, and Poisson's ratio of 0.36. The constitutive behavior of the elastomer was captured using a nearly-incompressible Neo-Hookean hyperelastic model (Poisson's ratio of v_0=0.499 and density of 1.0E3 kg/m3) with directly imported uniaxial test data described in “Material characterization” for silicone-based rubber.
Different element types were used to construct the three-dimensional (3D) FE models of the kirigami stent. Linear beam element (Abaqus element type B31, seed size=1) for the Kevlar fibers, 3D shell element with reduced integration (Abaqus element type S4R, seed size=1) for the plastic kirigami, and 3D brick element (Abaqus element type C3D8, seed size=1.5) for the elastomer actuator and the plastic plug. The Dynamic Explicit solver with a time period of 1000 and a mass scaling factor of 1000 (to facilitate convergence) was used. TIE constraint (surface to surface) was applied between the fibers and the elastomeric body. General Contact type interaction with penalty friction coefficient 0.2 for tangential behavior and “hard” contact for normal behavior were applied. Finally, the pressure load applied to the inner surface of the linear actuator using SMOOTH step amplitude curve, and the deformation of the kirigami stent model was monitored as a function of the applied pressure.
Using the numerical methods, as will be described below, an optimal stent design was identified that exhibits larger radial expansion (εr) and higher out-of-plane stiffness of the needles (K33) for better engagement with the surrounding tissue and injection while actuated.
A systematic study was carried out to predict the effect of t (the thickness of the outer shell of the stent) and l (the needle length) on the evolution of εa, εr, and θ as a function of the applied actuator pressure (P/P0) for an esophageal-sized stent with L0=8 cm and D0=12.5 mm. The deformation response of the kirigami stents can be controlled by varying the thickness of the kirigami shell (t) and the needle length (l) as a function of applied pressure. For example,
In the color maps shown in
To identify a preferred t for a given stent geometry, kirigami surfaces were fabricated with various thicknesses, and experimentally investigated the effect of t on the stiffness of the kirigami needles in the normal direction, denoted by K33. To measure the stiffness, a normal stiffness test was carried out (e.g., using an Instron 5942 series Universal Testing System). First, the surfaces were uniaxially stretched to different levels of strains, ε{circumflex over ( )}22=0, 0.5, 0.10, 0.15, and 0.20, which result in buckling out the needles. At each level of applied strain, the surfaces were immobilized to an acrylic plate and then compressed in the vertical direction, as illustrated in
As previously described, the kirigami shell (or stent) is capable of reversible shape transformation from flat configuration (for device delivery and removal) to 3D surfaces with popped-up needles (for injections) that enables facile delivery, robust deployment, and safe removal of the drug releasing system.
Referring now to
The experimental images were compared to the numerical snapshots obtained from non-linear FE simulations, showing that the kirigami shell is initially flat and then transform into 3D configurations with buckled out or protruding needles upon pressurizing the actuator. By releasing the pressure, the needles are popped in or retract and recovered their original undeformed shape. The deformation of the prototype was quantified and the experimental data (blue markers) was compared to the FE results (red dashed lines), showing a close agreement. The evolution of axial extension (εa=L/L0), radial expansion (εr=D/D0), and popping angle (θ) as a function of the actuator pressure (P/P0) illustrated in
Finally, to ensure the ability to fabricate the stents in multiple sizes and consistency with this numerical prediction, further esophageal stents were fabricated with multiple combinations of t and l, and their deformation was characterized using both FE simulations and experiments, showing an excellent qualitative agreement.
Evaluation of Controlled Penetration of Stent Needles to the GI Mucosa.
Micro-computed tomography (micro-CT) imaging and histology from ex vivo and in vivo experiments have been employed to demonstrate that the stent needles can be inserted by more than 1 mm into the submucosa of swine esophageal tissue without causing perforation. The penetration depth of the needles (d) can be controlled by incorporating the arc-shaped features (i.e., dimples with R=1.5 mm) on the two sides of the projection elements (i.e., needles) as illustrated in
To make the external surface of kirigami stent (especially the needles) radiopaque, the flat kirigami surfaces were coated with a thin layer of tungsten filled conductive ink (RO-948 Radio Opaque Ink, MICROCHEM) using a roller. The coated kirigami surface was left overnight to dry. The radiopaque stent prototypes with different needle's lengths (H) were deployed in the esophagus harvested from a Yorkshire pig. The esophagus was rinsed for approximately 10 sec under running tap water to wash away contaminants such as gastric fluid. To deploy the stent, a custom 3D printed fixture was used. The fixture consisted of a 20 mm diameter tube 3D printed out of VeroClear plastic. The 20 mm tube was placed inside the ex vivo esophagus to hold it open for deployment, and the stent with a given needle's length inserted into the esophagus via the tube. Once it reached the proximal esophagus, the pneumatic linear actuator inside the stent was inflated by pumping air using a plastic syringe connected to the stent (e.g., via the inlet port) via a Tygon PVC clear tubing results in popping up the needles. Syringe stopcock was used to maintain the pressure inside the stent's pneumatic actuator and keep all the needles popped up at the maximum angle (˜22°) against the surrounding esophageal tissue. The kirigami needles were inserted into the tissue by gently pulling the Tygon tubing backward via application of ˜8N force.
The deployed stent in the esophagus was then transferred into the micro-CT scanner and scanned following the protocol for soft tissue. Using the image viewer software, the penetration of the needles into the tissue was monitored by taking tomographic images at multiple views. The penetration depths were measured using both the cross-section and top views, where we were able to see the needle tips penetrated to the esophageal submucosa. The precise depths were obtained through measuring the distance between the inner surface of the tissue and the tip of the needles, d, as shown in
Note that all the needles were penetrated into tissues with an average tilting angle θ˜22°, which is the maximum popping angle (considering the surrounding tissue wall) achieved by pressurizing the soft actuator up to Pmax/P0=6.5 and gently pulling it backward via application of ˜8 N force. Therefore, the maximum penetration depths (dmax) were predicted as dmax=H sin θ=0.57, 0.96, and 1.34 mm for the needles 1 to 3 (dashed lines in
The kirigami stent prototypes were deployed for in vivo evaluations in a large animal model (50 to 80 kg female Yorkshire pigs ranging between 4-6 months of age). The pig was chosen as a model because its gastric anatomy is similar to that of humans and has been widely used in the evaluation of biomedical GI devices. An overtube (with D_in=⅝″ and D_out=¾″—US Endoscopy), with endoscopic guidance, was placed into the proximal esophagus to assist the placement of the stent. The stent with 8 cm length and 12.5 mm diameter was inserted into the esophagus via the tube pushed by the end of a scope. Once it reached the proximal esophagus, the overtube was removed, results in exposure of the stent to the esophageal mucosa. Similar to the ex vivo deployment, the pneumatic linear actuator inside the stent was actuated by pumping air using a plastic syringe connected to the stent via a Tygon PVC clear tubing caused buckling up the needles. Syringe stopcock was used to maintain the pressure inside the pneumatic actuator and keep all the needles popped up against the mucosa. The kirigami needles were then inserted into the submucosa by gently pulling the Tygon tubing backward via application of ˜8 N force. After deployment, the stent was left in place for 2 minutes before retrieval. The stent was then retracted by releasing the actuator pressure that makes the needles to buckle in and recover its original shape for easy removal.
As illustrated in
In Vivo Delivery of Fluorescent Polystyrene Microparticle
To enable loading and delivery of polymeric particles with the injectable stent system, the external surface of the stent (i.e., kirigami shell) was coated with a solution of fluorescent magnetic polystyrene microparticles. Fluorescent magnetic polystyrene microparticles (Fluorescent Nile Red Magnetic Particles, 1.0% w/v, 4.0-4.9 μm nominal size) and 25% w/v of Dextran sulfate sodium salt in double-distilled H2O Water were mixed with a ratio of 5:2. 10% w/w of glycerol as a plasticizer was added to the mixture. The final mixture was vortexed for 10 minutes before coating.
A custom-built benchtop spray coating set-up with programmable stent movement and rotation was used to achieve a uniform thin film coating of the solution onto the kirigami stent shell, shown in
The airbrush 560—used to spray the coating solution through its nozzle—was connected with a silicone tubing to a 30 ml pressurized coating solution vessel 564 and placed on a magnetic stirrer for continuous mixing, feeding and spraying the solution. The vessel 564 was equipped with a pressure pump 566 controlled by software (e.g., on the PC controlling unit 568). Two nitrogen gas tanks 552 were used to supply pressure for the pressure pump 566 (400 KPa) and airbrush 560 (50 KPa) during the coating process. The feeding pressure was optimized (5-60 KPa) and set to 40 KPa (equal to 40 μl/min) to reach a constant solution flow and uniform spraying pattern. The whole coating process consisted of eight coating steps (four infuse and four withdraw).
The fluorescent magnetic microparticles were delivered in vivo in three porcine esophagi using the coated prototypes (
Similar kirigami-based stents with proper size and needle stiffness were prototyped to demonstrate the ability of safe delivery and circumferential injection of fluorescent microparticles in other tubular parts of the body including femoral arteries and trachea. For example, as illustrated in
In vivo sustained drug release through deposition of polymeric particles loaded with therapeutics
To evaluate the performance of the kirigami stents for extended drug release, in vivo studies were conducted in swine, and using budesonide as a model drug, demonstrating that the injectable stent delivers drug depots for up to a week through multipoint submucosal deposition of drug-loaded polymeric particles. Budesonide, an anti-inflammatory drug commonly used to treat inflammatory bowel disease and eosinophilic GI disorders, was encapsulated into poly lactic-co-glycolic acid (PLGA) microparticles using continuous microfluidic droplet generation method. Three formulations of budesonide loaded PLGA particles with 75, 100, and 125 mg/ml concentration of budesonide, were formulated and are denoted by BUD 75, BUD 100, and BUD 125, respectively. Additionally, 100 mg/ml concentration of fluorescent budesonide-PLGA particles (BUD 100F) was synthesized via the addition of a fluorescent agent.
Budesonide-PLGA [Poly(D,L-lactide-co-glycolide) ester terminated, lactide:glycolide 75:25, Mw 76,000-115,000, Sigma Aldrich] microparticles were synthesized using a continuous microfluidic drug-PLGA droplet generation method, shown in
The one reagent glass 3D flow-focusing microfluidic chip 608 with hydrophilic surface and 100 μm deep channels was used, followed by a solvent extraction step. Two partially miscible solvents including dichloromethane and water were used as drug solvent/carrier and droplets carrier phases, respectively. Budesonide (75, 100, and 125 mg/ml) and 1% w/v PLGA were dissolved in DCM as an organic fluid. 2% w/v PVA in double-distilled water was used as an aqueous/carrier phase for droplet generation. All fluids passed through a 0.2 μm pore microfilter before droplet production. To generate fluorescence-sensitive budesonide-PLGA particles, 0.3% w/v of PLGA-SH (LG 50:50, PolySciTech) and 20 μl of Alexa Flour 647 C2 Maleimide dye (Invitrogen) was also added to the budesonide-PLGA solution.
The microfluidic system set-up 600 includes two pressure pumps 606 equipped with in-line flow rate sensors to monitor and control the streams flow rates. Two flow rate sensors, 30-1000 μl/min and 1-50 μl/min, were employed in the organic line and aqueous line, respectively. An air compressor (not shown) provided the supply pressure for the pressure pumps 606 at 400 KPa working pressure. The pumps 606 were connected to 30/400 ml and 30 ml volume remote pressure chambers 602, 604 placed on magnetic stirrer for continuous mixing and delivering of PVA in water and DCM-PLGA-Budesonide solution to the chip 608 with 10 μl/min aqueous/carrier rate and 1.35 μl/min organic/drug-PLGA solutions rate, respectively. The particle synthesis process was continuously continued to reach 500 mg of particles while the DCM solvent was evaporating/by connecting the particle's collection siliconized stirred vessel to very mild vacuum pressure (about 650 Torr). Three formulations of budesonide-PLGA particles was synthesized with 75, 100, and 125 mg/ml concentration of budesonide, denoted by BUD75, BUD100, and BUD125, respectively. Additionally, 100 mg/ml concentration of fluorescent budesonide-PLGA particles (BUD 100F) was synthesized via addition of Alexa Flour 647 C2 Maleimide as described.
The size of the prepared formulations for the drug-loaded particles (BUD 75, BUD 100, BUD 125, and BUD 100F) was measured for an average of 80-100 particles. A digital camera equipped with an optical microscope used to visualize the particles, and counted by advanced image analysis software. About 9-11 mg of microparticles (MPs) in 3 replicates were suspended and dissolved in 0.5 ml of acetonitrile by vortexing for 5 min. Then, 500 μl of the solution with 5-fold dilution were prepared and drug concentration in the replicates was measured using HPLC analysis (High Performance Liquid Chromatography) described below.
The obtained HPLC data were used to calculate drug loading and encapsulation efficacy parameters reported in
Budesonide kinetic release studies were analyzed using High-Performance Liquid Chromatography (HPLC). A 1260 Infinity II HPLC system equipped with a 1260 quaternary pump, 1260 Hip ALS autosampler, 1290 thermostat, 1260 TCC control module, and 1260 diode array detector. Data processing and analysis was performed using software. Budesonide chromatographic isocratic separation was carried out on an Agilent 4.6×150 mm Zorbax Eclipse XDB C-18 analytical column with 5 μm particles, maintained at 30° C. The optimized mobile phase consisted of 20 mM dipotassium phosphate buffer (pH 3.00 adjusted with phosphoric acid) and acetonitrile [30:70 (v/v)] at a flow rate of 1.00 mL/min over a 5 min run time. The injection volume was 5 and the selected ultraviolet (UV) detection wavelength was 244 nm at a bandwidth of 4.0, no reference wavelength, and an acquisition rate of 40 Hz.
Drug release occurs through polymeric membrane erosion, allowing the drug to diffuse out from the dialysis membrane. The in vitro release kinetics of budesonide from the PLGA particles in a biorelevant fluid, phosphate-buffered saline (PBS), were analyzed using High-Performance Liquid Chromatography. The in vitro release of budesonide from microparticles was performed using a horizontal shaker with 200 rpm speed at 37° C. Three to 5 milligrams of budesonide loaded microparticles were added to 1 ml phosphate buffered saline (PBS pH 7.4 (1×)) with 0.1% Tween 20. Experiments were performed at 37° and samples were taken at 2, 4, 6 h, and then daily up to 7 days of release. Buffer were refreshed at different time intervals and the drug content was analyzed using HPLC analysis of 500 μl of supernatant solution as described via High-Performance Liquid Chromatography analysis. Notably, the release profile of encapsulated budesonide demonstrated an initial burst release followed by linear drug release up to approximately 40% across all the formulations over 7 days of incubation in PBS at 37° C., as illustrated in
The needle surfaces/tips of kirigami stents were loaded by pipetting 20 μl of the BUD 100F particle solution two times per needle with a 5 h interval for drying at room temperature.
To evaluate the ability of kirigami stents, loaded with the budesonide-PLGA microparticles, to achieve long-term delivery, we administered them to a large animal model (three Yorkshire pigs). The details of delivery, deployment and removal of the stents were the same as fluorescent particle-loaded stents previously described. The three animals were euthanized at three different time points: 1 day, 3 days and 7 days after deployment/removal in compliance with the AVMA Guidelines on Euthanasia. Endoscopic evaluation of the esophagus over the course of the study was performed to further explore the esophagus and ensure the absence of any ulceration or injury. The esophagi of three pigs were harvested and 8 mm diameter biopsies were used to take biopsies at least seven needle penetration sites per retrieved esophagus. The penetration sites were recognized by using an IVIS Spectrum in vivo imaging system. Note that the drug-loaded particles (BUD 100F) were fluorescence-sensitive due to incorporation of Alexa Flour 647 C2 Maleimide. The biopsies were then frozen until extraction. Budesonide was extracted from esophageal tissue by placing each biopsy in 500 μl of 5% BSA in PBS and homogenizing two times by 6500 rpm for 30 seconds. A 100 μl fraction of the homogenate was collected. 50 μl of 5 μg/ml hydrocortisone in acetonitrile and 1 mL ethyl acetate was added for budesonide extraction. These samples were vortex and centrifuged for ten minutes at 13000 rpm. Following centrifugation, the supernatant was evaporated to dryness. Samples were reconstituted in 300 μl acetonitrile and 200 μl of the reconstitute were pipetted into a 96-well plate containing 200 μl of Nanopure water and used for ultraperformance liquid chromatography—tandem mass spectrometry (UPLC-MS/MS) analysis.
The esophagi was analyzed using ultraperformance liquid chromatography-tandem mass spectrometry (UPLC-MS/MS). The analysis was performed on a Waters ACQUITY UPLC- I-Class System aligned with a Waters Xevo-TQ-S mass spectrometer. Liquid chromatographic separation was performed on an ACQUITY UPLC Charged Surface Hybrid C18 (50 mm×2.1 mm, 1.7-μm particle size) column at 50° C. The mobile phase consisted of aqueous 0.1% formic acid and 10 mM ammonium formate solution (mobile phase A) and an acetonitrile: 10 mM ammonium formate and 0.1% formic acid solution [95:5 (v/v)] (mobile phase B). The mobile phase had a continuous flow rate of 0.6 ml/min using a time and solvent gradient composition. The initial composition (100% mobile phase A) was held for 1 min, after which the composition was changed linearly to 50% mobile phase A over the next 0.25 min. At 1.5 min, the composition was 20% mobile phase A, and at 2.5 min, the composition was 0% mobile phase A, which was held constant until 3 min. The composition returned to 100% mobile phase A at 3.25 min and was held at this composition until completion of the run, ending at 4 min, where it remained for column equilibration. The total run time was 4 min, and sample injection volume was 2.5 μl. The mass spectrometer was operated in the multiple reaction monitoring (MRM) mode. Sample introduction and ionization was by electrospray ionization (ESI) in the positive ionization mode. MassLynx 4.1 software was used for data acquisition and analysis. Stock solutions of budesonide and internal standard hydrocortisone were prepared in methanol at a concentration of 500 μg/ml. A twelve-point calibration curve was prepared in methanol ranging from 1 to 5000 ng/ml.
The concentrations of budesonide delivered using the injectable stents are reported in
Discussion.
In summary, given the importance of innovative device development, a class of drug releasing systems has been developed which are capable of multipoint injecting drug depots in the tubular mucosa of the GI tract such as the esophagus, enables sustained local drug delivery. Implementations of such a system were developed by: (i) design, FE modeling, and prototyping a kirigami-based stent platform and characterize the mechanics for robust deployment, multi-point injection, and safe removal in the tubular mucosa of the GI tract, and (ii) in vivo evaluation of the capacity to deposit drug-loaded polymeric particles for extended release using a large animal model. To develop the kirigami stent, first, buckling-induced kirigami surfaces were engineered to undergo a shape transformation from flat surfaces to 3D textured surfaces with popped-up needles. By turning kirigami surfaces to cylindrical kirigami skins, a systematic study was presented through combining FE simulations and experiments to investigate the effect of kirigami mesostructure (needle length and thickness) on the mechanical response of kirigami shells. Next, a fluid-powered elastomeric actuator was employed to generate linear output motion using a simple control input (i.e., pressurization of a working fluid) to trigger the kirigami shell for injection. By combining kirigami design principles and the pneumatic soft actuator a new way to deliver drug depots locally is provided and can be used to administer other APIs. Altogether, this design of injectable kirigami stent offers a unique mechanism with a range of advantages: (i) can be applied to various length-scales to be matched with the size of the target tubular compartments of the GI tract and airways; (ii) be able to rapidly deploy by more than 50% radial expansion and release therapeutics into submucosa through circumferential injections, and (iii) shape recovery to the original flat configuration by releasing the actuator pressure for safe removal.
Plasma surface treatment that activates the plastic kirigami surfaces and results in the creation of hydrophilic surfaces, and laser engraving the needle surfaces to increase surface area were used as two post-treatment techniques to improve adhesion bond between the coating layer (drug-particle solution) and kirigami stents needles that consequently enhance drug loading capacity. However, some drug particles may be lost by washing off the stent during delivery. Further studies on various polymeric or plastic surfaces to make the kirigami shell with enhanced drug loading capacity as well as polymeric sacrificial layers to protect the drug-coated particles, can be performed to further boost drug loading capacity and protected delivery without losing drug particles that finally leads to improved local drug delivery. Introducing a sheath will also protect the stent during delivery and eliminate the use of a separate introducer for facile delivery. Other designs including bi-material kirigami surfaces that includes plastic hinges to provide out of plane popping motion and insoluble drug-loaded needle tips for injection and deposit drug depots by dissolution into the mucosa could improve delivery performance, and reduce the risk of tissue inflammation, and safe removal. These technologies can be further improved through further in vivo testing of injectable kirigami stents with the aforementioned improved designs and evaluation sustained release within a target therapeutic range across different drugs depending on the target application sites. Another design of kirigami-based stents includes a design in which the kirigami spikes act as actuators to pop out and expose the attached small hypodermic needles for insertion. The hypodermic needles are connected via microchannels to the space inside the actuator as a drug reservoir to transfer liquid therapeutics.
While various spatial and directional terms, such as top, bottom, lower, mid, lateral, horizontal, vertical, front, and the like may be used to describe examples of the present disclosure, it is understood that such terms are merely used with respect to the orientations shown in the drawings. The orientations may be inverted, rotated, or otherwise changed, such that an upper portion is a lower portion, and vice versa, horizontal becomes vertical, and the like.
Within this specification, embodiments have been described in a way which enables a clear and concise specification to be written, but it is intended and will be appreciated that embodiments may be variously combined or separated without parting from the invention. For example, it will be appreciated that all preferred features described herein are applicable to all aspects of the invention described herein.
Thus, while the invention has been described in connection with particular embodiments and examples, the invention is not necessarily so limited, and that numerous other embodiments, examples, uses, modifications and departures from the embodiments, examples and uses are intended to be encompassed by the claims attached hereto. The entire disclosure of each patent and publication cited herein is incorporated by reference, as if each such patent or publication were individually incorporated by reference herein.
Various features and advantages of the invention are set forth in the following claims.
The present application is based on, claims priority to, and incorporates herein by reference in its entirety, U.S. Provisional Patent Application No. 63/041,154 filed Jun. 19, 2020, entitled “Kirigami-inspired Stents for Sustained Local Delivery of Therapeutics.”
This invention was made with Government support under Grant No. R01 EB000244 awarded by the National Institutes of Health. The Government has certain rights in the invention.
Number | Date | Country | |
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63041154 | Jun 2020 | US |