Stent localized by low radial force

Abstract

A low radial force stent with good resistance to migration is described, comprising a microstructured surface that generates an inward radially oriented grip into the lumen. In particular, a stent for deployment into a biological lumen is described, in which a novel combination of low outward radial force in the lumen and resistance to shear slip is achieved by a hierarchical microstructured surface that provides a non-frictional grip on the lumen surface. Combinations of microstructured surfaces combining frictional and non-frictional grips for low radial forces are described, independent of axially dependent changes in the stent diameter or stent oversizing. These combinations of microstructured surfaces, when placed on the outer surface of the stent, provide a non-movable stent. The disclosed hierarchical levels of microstructure may be composed of a complex of microstructures themselves, and may or may not be self-similar to other hierarchical levels.

Description

Technical Field

[0001] Minimally invasive percutaneous implantation techniques have been found to be widely accepted as an alternative to conventional surgery for cardiac patients at high risk of morbidity and mortality. Minimally invasive procedures for the treatment of non-vascular lumen defects have been developed. In general, minimally invasive implantation techniques are likely to become a major means of treating many, if not all, lumen diseases.

[0002] Transcatheter pulmonary valve implantation can be safely performed, particularly in patients who have undergone surgery on the right ventricular outflow tract during repair of congenital heart disease. Nitinol stents are commonly used for the minimally invasive delivery and fixation of heart replacement valves. High elastic strain that is resistant to nitinol can reduce the risk of stent damage during implantation and can reduce the large diameter required for minimally invasive delivery.

[0003] Inadequate tissue conformity of the prosthesis can lead to stent migration from the target position, so accurate sizing of the stent for fixation in the host lumen should be considered. More specifically, with respect to heart valve repair, stent migration can result in a serious condition known as paravalvular regurgitation. On the other hand, excessive radial fixation force on the tissue can be associated with the risk of tissue wall degeneration and damage.

[0004] Stent migration and lumen wall damage are major concerns in all stent implantations. Unfortunately, improving these two failure modes requires conflicting design criteria. While stent migration can be reduced by increasing the radial force of the stent, reducing lumen damage corresponds to a decrease in the radial force of the stent.

[0005] In addition to wall degeneration and damage, matching the stent size to the vessel size is complicated by the continuous radially outward expansion force exerted by self-expanding stents after deployment, leading to negative chronic recoil and vessel dilation after follow-up. Therefore, it is necessary to reduce the radial force exerted by the stent, not only to reduce wall degeneration and damage, but also to reduce chronic recoil and vessel size dilation. Excessive radial force in vascular stents is required to produce a large positive radial force to fix the stent, and the negative effects of this positive radial force are contrary to the conditions for proper fixation and function.

[0006] Stent oversizing and excessive radial force can further lead to tissue remodeling. The presence of permanent stents and the amount of force they exert are known to affect host tissues, posing a risk of causing adverse biological processes such as thrombosis, in-stent restenosis, and neointimal proliferation. Therefore, stent localization features are needed to prevent stent oversizing to achieve the desired fixation outcome. The problem of tissue remodeling is not limited to cardiac stents, and therefore, it will be understood by those skilled in the art that the stents of this disclosure are useful in all medical stent placement applications.

[0007] In vivo functional evaluation related to stent localization and stent mating involves various mechanical parameters. Fortunately, the ratios of stent parameters fall within a specific range across stent types, provided that proper mating function is separated from the anti-displacement function. For the first time, the mechanical and structural functions of a novel class of low radial force stents are disclosed herein. [Background technology]

[0008] Embodiments disclosed herein include 40 streams connected by an inclined bridge The rat may consist of a repeating design of four rings. In some embodiments, the inner diameter of the stent may be about 30 mm, and the struts may have cross-sectional dimensions of about 0.3 mm × 0.4 mm on average.

[0009] Parallel plate compression (collapse test) involves placing a stent between parallel plates at room temperature on an Instron machine that applies pressure across the entire plate. The protocol involves obtaining a force in proportion to the plate separation. Then, the plate separation is increased to the initial plate separation, obtaining the opposite force in proportion to the plate separation. The two plots are usually different, indicating that the stent has undergone irreversible compression, and the difference can be referred to as hysteresis.

[0010] Radial compression involves applying a uniform, inwardly directed radial force. The crimp head resembles a multiple-fingered chuck that applies a uniform force around the outer circumference of the stent. As with parallel plate compression, the chuck motion is reversed and the force is measured according to the radius. In both tests, the test range depends on the size of the stent and generally extends to the point of discontinuity, i.e., buckling in the stent. Subsequent measurements are then taken immediately before the point of discontinuity.

[0011] When internal pressure is applied to an artery, stress is generated in both the longitudinal and circumferential directions. Longitudinal stress is a result of the internal pressure acting on the ends of the artery, causing strain along the axis. Circumferential (hoop) stress is a result of the radial action of the internal pressure acting on the artery wall, increasing its diameter. The deployment of a stent into an artery generates internal pressure within the artery. The effect of the crimping tool on the stent generates external pressure on the stent. As the stent expands, it exerts an outward radial force (RF) on the artery, which in turn generates hoop stress on the artery wall, associated with the expansion hoop force (HF). Similarly, RF applied by the crimping tool can generate hoop stress on the stent, from which HF can be derived.

[0012] The terms chronic outward force (COF) and radial resistance force (RRF) describe specific characteristics of stent material and geometric shape. COF refers to the stent's release force acting on the artery as it attempts to return to its nominal diameter. COF is modeled as a circumferential force and generates hoop stress. RRF refers to the force generated by the stent to resist compression. RRF is modeled as a compressive force acting on the stent. Compression is directed circumferentially and acts as a hoop force, despite the fact that it is called a radial force.

[0013] In an idealized model, RF and HF can be related through equations. In reality, RF and HF may differ in magnitude and direction, and may not be directly compared. However, in the tests described here, a standard thin-walled tube model is used when the force cannot be measured directly.

[0014] Therefore, in in vitro tests, direct measurement of RF is performed by the applied radial force, directed inward in the case of crimping and outward in the case of stent expansion. Mylar film tests compress the stent circumferentially, resulting in direct measurement of HF. This force can be expressed as RRF when reporting COF with values ​​for stent compression and stent expansion.

[0015] In a collapse test, RF or HF measurements are not performed because the applied load is vertical. On the other hand, it is easier to measure than hoop loading, and this loading mode is important in situations where the stent is exposed to external collapse, such as from the carotid artery. [Overview of the Initiative]

[0016] Referring to Figure 1A, which shows half the circumference of the stent 100, a portion of the stent cylinder 102 is shown. When the stent 100 is internally pressurized by the expansion radial force 104, a circumferential stress 106 may be generated. The strain on the cylinder wall 108 then generates a compressive radial force 110 and a hoop force 112. When the stent 100 is deployed in the artery, the expansion radial force 104 of the stent balances the compressive radial force 110.

[0017] This disclosure describes how the Wenzel-Cassie microstructure on the outer surface of the stent can be modified to alter the usual relationship between radial and hoop forces, thus allowing for design considerations regarding how radial and hoop forces are distributed. For example, in some stent applications, radial forces may be more tolerable than hoop forces, and vice versa in other stent applications. In most cases, an overall reduction in both hoop and radial forces is desirable.

[0018] The clinical objectives of stent patency and stent localization are not directly related to or are isolated from forces when stent localization by the microstructure of this disclosure is utilized. Therefore, one or more embodiments of this disclosure may provide microstructure stents that are independent of these forces in order to achieve clinical objectives.

[0019] Referring to Figure 1B, a microstructured surface 150 is described, comprising a substrate 152, a first microstructure 154, a second microstructure 156, and a third microstructure 158. The first microstructure 154 may be a two-dimensional sinusoidal microstructure. The second microstructure 156 may be a pillar microstructure. The third microstructure 158 may be a fluted microstructure. Various microstructures are arranged hierarchically. In certain embodiments, the dimensions of the first microstructure 154 are 1.1 to 10 times the dimensions of the second microstructure 156, and the dimensions of the second microstructure 156 are 1.1 to 10 times the dimensions of the third microstructure 158, and the diameter, pitch, and height dimensions can be defined such that the pillars may have any elliptical or polygonal cross-sections.

Brief Description of the Drawings

[0020] [Figure 1A] It is a diagram of hoop force and radial force within the lumen. [Figure 1B] It is an exemplary embodiment having a hierarchical microstructure including three levels. [Figure 2] It is a diagram showing the relationship between hoop force and radial force. [Figure 3] It is an exemplary embodiment of a stent having a microstructure surface. [Figure 4] It is an exemplary embodiment of a three-component stent having a microstructure surface. [Figure 5] It is an exemplary embodiment of a stent in which the microstructure conforms to the contraction / expansion of the stent. [Figure 6] It is an exemplary embodiment of a microstructure stent whose radius can be adjusted in situ to minimize the radial force. [Figure 7] It is an exemplary embodiment of a microstructure stent in which the microstructure is localized on the stent and designed to match a specific geometric function of the target lumen.

Best Mode for Carrying Out the Invention

[0021] Generally, the term "appearance phenomenon" as used herein is understood to represent the interfacial structure formed between the microstructure surface of the stent and the lumen of the target tissue. In some embodiments, separating the compression load requirements created by the target lumen from the radial force requirements may be related to reducing the movement of the stent. Thus, the design-to-target methodology utilized in some embodiments of the present disclosure can balance the axial tensile test that measures movement against the radial force test that supports stent viability.

[0022] Generally, the term "composite hierarchical microstructure" as used herein is understood to describe a hierarchical structure where each layer can be at a level with a composite microstructure. The composite microstructure can be understood as various geometric microstructures arranged on a single layer or substrate, or arranged in layers on that substrate. Some composite hierarchical microstructures may be self-similar, i.e., the composite microstructure of one layer may be an enlarged or reduced version of the microstructure on one or more other layers.

[0023] As described above, all terms, methodologies, and test scenarios are described in detail and function as definitions of the terms appearing in this disclosure. For example, shear force, hoop force, radial compressive force, radial expansion force, and hysteresis derive their meaning from the protocols and test settings described herein.

[0024] In the evaluation of an intravascular stent, the force is recorded by a load cell placed at the end of a film loop arranged around the stent and can be measured in response to vertical displacement. Clinically relevant forces are, in certain embodiments, chronic outward force (COF) and radial resistance force (RRF). COF is the unloading force measured at a nominal diameter minus 1 mm. RRF is the force required when crimping the stent embodiment to its nominal diameter minus 2 mm.

[0025] In the evaluation of an intravascular stent embodiment as described herein, a crush test may be utilized to measure the force recorded by a force gauge connected to a plate applying vertical compression. Clinically relevant forces can be the radial stiffness, which is the radial force per unit length depending on displacement.

[0026] In the evaluation of an intravascular stent embodiment disclosed herein, a U-shaped measurement system measures the force applied by the stent to a tubular holder connected to an electronic balance. Clinically relevant forces can be the radial force (RF), which is typically recorded depending on the stent diameter.

[0027] In particular, in the evaluation of the biliary stent embodiments disclosed herein, radial force may be measured using a force gauge deployed inside the contracting cylinder. Clinically relevant forces may be radial force, typically recorded as RF relative to diameter, and RF at a diameter of 4 mm.

[0028] In the evaluation of the embodiments of giant costents disclosed herein, radial forces may be measured using the Mylar method. The resulting plot may be given as radial force against reduction in cross-sectional area.

[0029] In evaluating embodiments of esophageal stents disclosed herein, radial force may be determined by lateral compression, circumferential compression, and / or a three-point bending test. A clinically important parameter may be COF, measured as RF against diameter.

[0030] One aspect of this disclosure and embodiments described herein involve connecting a Wenzel-Cassie microstructure surface to various stent applications / designs and evaluating relevant clinical parameters, where oversizing may be replaced by a Wenzel-Cassie junction. This approach can result in stents with reduced radial forces and reduced chronic complexity. By separating the stent force requirements from the stent localization requirements, it is possible to create stent designs that are less prone to the complexity of traditional stent designs in the prior art and can be easily delivered through minimally invasive means.

[0031] The traditional frictional force used for localization in conventional stents requires a large outward-directed radial stent force that creates a mechanical bond between the stent and the lumen surface. It is understood that the frictional force is proportional to the RF (radius force). Therefore, for conventional stents, if the RF is 0, the frictional localization force is also 0. This is not true for Wenzel-Cassie non-frictional localization, where the localization force is non-zero and the outward radial force can be zero.

[0032] A hierarchical microstructure surface can be understood as a surface microstructure comprising a high surface energy region juxtaposed with a lower surface energy region. These high and low surface energy surfaces do not need to be stacked, but rather need to be interlocked or juxtaposed. In some embodiments, high Wenzel-Cassie localization forces can be achieved when the microstructure spans three-dimensional space and is stacked.

[0033] High surface energy microstructures can be understood as wet. In the case of water wetting, high surface energy microstructures can be understood as hydrophilic in some embodiments. Low surface energy microstructures can be understood as non-wetting. In the case of water wetting, low surface energy microstructures can be understood as hydrophobic in some embodiments. The combination of juxtaposed wet and non-wetting zones creates what is known as the Wenzel-Cassie interface, as disclosed herein.

[0034] It will be understood that the Wenzel-Cassie interface disclosed herein may not require outward radial forces to create localization forces compared to conventional prior art devices. Applying energy to move the stent parallel to the lumen surface may disrupt the low-energy states of the Wenzel-Cassie interface. In some embodiments, hierarchical microstructured surfaces may reduce contact with the lumen surface (e.g., surface area, pressure, friction, etc.) compared to conventional prior art devices. In certain embodiments, the stents disclosed herein may not contact the lumen surface at all, yet still create localization forces to fix the stent to the target site.

[0035] The Wenzel-Cassie interface, in particular, can bond a stent to the lumen surface via van der Waals interactions, as an example that does not necessarily involve mechanical interaction. In some embodiments, the stents of the present disclosure may include some frictional mechanisms, but at least a portion of the localized forces may be supplied by non-frictional Wenzel-Cassie joints, and thus RF may be reduced and separated from localized forces in the stents of the present disclosure compared to conventional stents of the prior art.

[0036] The Wenzel-Cassie interface can be characterized by various nomenclas in the literature. A Wenzel-Cassie interface may exist when there is a "suck-down" effect between the microstructure surface of a device and the target surface. The suck-down effect can create non-destructive inward-directed RF, as opposed to destructive outward-directed RF. The suck-down effect can reduce the energy of the Wenzel-Cassie interface, creating a "jointed" state where energy input is required for fracture.

[0037] For this reason, the Wenzel-Cassie junction phenomenon is also called "fluid thinning." Here, the term "Wenzel-Cassie interface" can be understood as a phenomenon that generates a fluid thinning state at the interface between the hierarchical microstructure surface and the target surface. For example, in some embodiments, the hierarchical microstructure creates a capillary effect, which, although not technically Wenzel-Cassie, still uses fluid interactions with changing surface energy configurations and gradients to generate inward RF. Thus, the hierarchical microstructure surface of this disclosure is characterized as a surface in which the contact area between the microstructure and the inner surface is reduced. In some embodiments, the contact area can be reduced by at least 25%, at least 30%, at least 35%, at least 40%, at least 45%, and most preferably at least 50% or more. In some embodiments, the contact area can be reduced by 75% or more. In further embodiments, the contact area can be reduced by 90% or more. Even further, the contact area can be reduced by 95%, 98%, and 99% or more.

[0038] The microstructured surfaces of this disclosure may be periodic in some embodiments. This periodicity can be used to create a secondary Wenzel-Cassie interface, which is hereby considered a Wenzel-Cassie interface. The secondary Wenzel-Cassie interface can be characterized as any type of mechanical deformation of the target surface and / or the microstructured surface that creates an interlock state between the microstructured surface and the target surface. It will be understood that the surface interlock and deformation resulting from the Wenzel-Cassie interface can be caused mostly by non-frictional forces.

[0039] For example, in some embodiments, the target surface may have an inherent wrinkle mode of deformation that causes wrinkles on the target surface without causing mechanical damage to the target surface. One or more periodicities of the microstructure surface may be adapted to the inherent wrinkle mode of the target surface so that the target surface and the microstructure surface interlock with minimal mechanical contact. In some alternative embodiments, the microstructure surface may have an inherent wrinkle mode, which can achieve the same secondary Wenzel-Cassie effect. Essentially, shear forces can be converted into wrinkles on either the stent surface or the lumen surface without stent displacement. This secondary Wenzel-Cassie effect may be particularly useful in lumens with peristaltic movement, such as the esophagus.

[0040] In certain embodiments of this disclosure, related but distinct secondary Wenzel-Cassie effects may be associated with Schallamach waves. These waves are generated on either a microstructured surface or a target surface and contain multiple wrinkle components, some of which may not be eigenmodes. The periodicity of the microstructured surface can "trap" some components of the Schallamach waves and allow others to "pass through." This phenomenon may be useful in localized stents where heterogeneous aspects are present in the lumen of the region where the stent is implanted, i.e., where food or feces are present. This type of secondary Wenzel-Cassie interface may be understood as the "slip-grip" interface as used herein.

[0041] In some embodiments, other secondary and tertiary Wenzel-Cassie interfaces may be incorporated, and the Wenzel-Cassie interfaces and the periodicity of the microstructure surface cooperate to localize the stent placed in a dynamically changing lumen. The periodicity of the microstructure may further cooperate and be coupled with other mechanical properties of the microstructure surface, such as the Young's modulus of the material containing the microstructure surface. Furthermore, the geometric shape of the microstructure surface and the method of attachment to the stent can also enhance or reduce these secondary Wenzel-Cassie effects.

[0042] Certain embodiments of the stents in this disclosure may be characterized by mechanical properties different from any stents previously described in the prior art. The microstructured surfaces of the disclosure, which utilize hydrophobic and hydrophilic regions that result in stents having increased displacement forces when the target contact surface / lumen is wet, or when oil or surfactant is applied, are not disclosed in the prior art. It should be understood that the functional modes and resulting effects as disclosed herein are the opposite of those of stents with mechanical friction surface textures where the in vivo environment tends to be slippery and therefore may not promote or may hinder frictional resistance to movement.

[0043] In friction treatments involving barbs, raised struts, surface cylinders, pyramids, etc., supplying large radial forces is necessary to effectively facilitate the interpenetration between the structure and the surface friction geometry. It is necessary to do so. Conversely, in this disclosure, the Wenzel-Cassie interface created by the microstructured surface when wet can prevent movement through various non-contact effects, broadly known as van der Waals forces. These localization forces can be understood to be mostly electronic and quantum mechanical, rather than mechanical and classical.

[0044] In some embodiments of this disclosure, the Wenzel-Cassie effect can create a “suck-down” effect between the lumen wall and the stent. Thus, an inward radial force can be applied rather than an outward radial force applied to the lumen for fixation. This embodiment is particularly useful in the case of stent placement for aortic aneurysms, where the strong outward radial force typical of prior art devices can worsen the aneurysm.

[0045] Those skilled in the art should understand that the effects disclosed in this application do not preclude a combination of the Wenzel-Cassie effect and the mechanical friction effect. In some embodiments, depending on location and physiological requirements, one effect, namely the Wenzel-Cassie effect or mechanical friction, may be preferred over the other in different parts of the same stent.

[0046] In addition, further embodiments may include adhesive and slippery surface textures. Such surface textures may be used alone or in combination. Fluid thinning and fluid thickening are common effects resulting from the disordering or ordering of water molecules caused by the influence that the spatially distributed surface energy pattern of the microstructure surface may have on electric dipoles in the interfacial fluid. In certain embodiments, regions of high surface energy can be juxtaposed with regions of low surface energy, thereby allowing peristaltic waves to pass through the stent without displacing it from its target position.

[0047] To enable those skilled in the art to utilize this disclosure, a brief description of the methods of use is provided below by describing various embodiments of this disclosure and by providing definitions of force terms used herein. Combinations of in vitro and ex vivo settings are useful for characterizing various embodiments of this disclosure. All tests are performed on laser-cut nitinol stents to illustrate the methodology, but it will be understood that various stent materials or stent geometric shapes that satisfy the parameter constraints of this disclosure may be used in this disclosure.

[0048] Experimental Protocol 1

[0049] Stent collapse

[0050] The stent section was positioned between two parallel plates mounted on the Instron machine. Two intersections on the stent were fixed to the bottom plate to prevent rotation. During compression, only the top plate was moved. The surface area of ​​the contact between the top plate and the stent was measured according to the plate displacement. Various contact surfaces (plates), including exvivo structures, were studied. The stent was compressed to a gap of 5 mm between the plates. The displacement rate was 1 cm / min. The plot of plate separation against force is of primary importance.

[0051] Experimental Protocol 2

[0052] Stent crimping and expansion

[0053] The modified four-finger milling chuck has curved aluminum pieces attached to each finger so that a circular cross-section is obtained in the closed position, and between the curved aluminum pieces is a stretchable elastomer membrane to prevent stent "extrusion" between the fingers. Small pressure transducers were embedded in the inner surface of each finger to form a smooth surface upon which various test surfaces could be placed.

[0054] After crimping, the chuck was operated in the opposite direction to allow for self-expansion. The stent was crimped in vitro to 15 mm. Furthermore, the stent was further compressed to a diameter of 6 mm to copy the crimping to a delivery system with a diameter of 18 F. All measurements were performed at near body temperature (37°C).

[0055] To understand the relationship between axially symmetric and circumferentially symmetric boundary conditions and to study the contact forces between struts, forces were measured in the horizontal plane and in a plane inclined at a 9-degree angle from the horizontal plane. The RF of a two-strut stent is determined by summing the radial reactions of all nodes (RFsum). To obtain the total RF of the stent, the result for the two-strut model stent was multiplied by 80, which reflects the axial and circumferential symmetry in typical stent designs.

[0056] The HF of an isolated stent portion was determined by summing the circumferential forces (HFsum1 and HFsum2) at the nodes on the lateral free surface of the stent. In some calculations, the contribution of the contact force (CF) was included in the HF. To obtain the total HF of the stent, the result of the two-strut model was multiplied by 2 (axisymmetric). Curves for RF and HF with respect to diameter were obtained.

[0057] Experimental Protocol 3

[0058] Stent Exvivo Lumen Interaction

[0059] In all tests, sheep pulmonary arteries were used. Symmetrical boundary conditions were applied in the circumferential direction. The pressure between the stent and the arterial lumen was measured. The ends of the arteries were fixed longitudinally to achieve inflation without change in length. In some cases, balloons were used to supply uniform pressure between the stent and the artery. To simulate stent deployment, contact between the stent and the lumen was minimized during insertion and then allowed to be made available by releasing the crimping mechanism.

[0060] To study the degree of oversizing, the diameter of the ex vivo artery was selected. Size adjustment was performed by applying a pressure of 2.2 kPa to the lumen, and the system was equilibrated to measure the size or diameter of the artery.

[0061] The arterial HF was determined by summing the circumferential reactions along all nodes of the tangential free surface (HFsum) and multiplying by 2. The increase in HF due to stent placement was calculated by subtracting the arterial HF from the HF value after pressurization.

[0062] During stent deployment, the vessel wall has a tensile HF that typically has a characteristic direction different from the stent's HF direction. Therefore, measurements were delayed until these two directions became one in the equilibrium deployment state. Note that after stent-arterial contact, the arterial HF increases and the stent HF decreases. The equilibrium point was determined at the intersection of the curves.

[0063] Referring to Figure 2, which shows half the circumference of the stent 200, it shows a cross-section of the stent cylinder 202 showing the radial force RFsum. The stent strut element 204 shows the hoop forces HFsum1 206 and HFsum2 208. The lumen element 210 shows the hoop force HFsum212, and the total HF = HF lumen = 2HFsum. The contact force CF214 is generated by the contact surface 216, and the total HF = HF stent = 2(HFsum1 + HFsum2 - CF).

[0064] The subject matter of this disclosure is further illustrated and illustrated by the following specific but non-limiting examples and experimental results. These examples and experimental results may include compilations of data representing data collected at various points in time during the development and experimentation process related to the subject matter of this disclosure.

[0065] Experiment 1

[0066] Radial size reduction without lumen contact

[0067] For the following experiments, the RF and HF were determined for stents coated with a smooth coating and stents coated with a composite pillar microstructure. All stents tested had a cylindrical geometric shape without end flares. Referring here to Figure 3, stent 300 is shown. Stent 300 may include a lumen contact surface 302, an inner surface 304, and a microstructure layer 306. In some embodiments, the stent wall may include nitinol struts 308. The microstructure layer 306 may include a composite pillar microstructure comprising a first layer 312 and a second layer 314.

[0068] The RF and HF values ​​for various stent diameters are shown in Table 1 below. Each stent was coated with nitinol, exhibiting a textured hysteresis curve (no contact with artery).

[0069] Table 1. Radial and hoop hysteresis of standard stents Stent diameter (radial direction, hoop) Force compression (radial direction, hoop) Force extension [Table 1]

[0070] It should be noted that the nitinol coating thickness was adjusted so that there was no detectable difference in the hysteresis curve between the coated smoothness and the coated microstructure. Table 1 shows the smooth and microstructured stents.

[0071] A characteristic of traditional conventional stents is that the force during stent delivery is higher than the force after stent delivery for the same diameter. This hysteresis is one of the main reasons why smooth stents must be oversized to leave sufficient radial force after delivery to secure the stent in the lumen. It should be noted that radial force is far more hysteretic than hoop force, meaning that increasing the radial force has little effect on stent patency. Consequently, there is no reason for a large radial force other than to overcome hysteresis and provide appropriate frictional localization force. Hysteresis can be beneficial if patency is separated from the localized frictional force.

[0072] Furthermore, in esophageal stent placement, patient comfort depends not on the hoop force, but on the length of the stent. It should be noted that this depends on the force and radial force. In esophageal stent placement, discomfort is the main adverse event of stent placement when the final lumen diameter is larger than the minimum diameter required to achieve lumen patency. In traditional stent placement applications, unnecessarily increasing the lumen diameter beyond the clinically required diameter can lead to lumen deterioration. For example, in aortic aneurysm stent placement, stent placement can cause significant dilation of the artery. In coronary artery stent placement, abrupt changes in lumen diameter can cause thrombus-forming turbulence.

[0073] Experiment 2

[0074] Reduction of radial force due to lumen contact

[0075] The hysteresis problem of traditional conventional stents can be explained by the fact that if post-delivery RF is to be increased, the delivery RF must be substantially increased, which is a limiting factor in developing stent designs. To meet these requirements, designs are likely to increase the minimum delivery diameter of the stent. Therefore, to make stent delivery practical in a wide range of applications, it is necessary to reduce the post-delivery RF specifications.

[0076] It should be recognized that simply pulling the stent against the lumen is not possible in order to test the effect of the microstructure surface on post-delivery RF requirements regarding stent slip or migration. Most stents are designed to reduce their radius when axial force is applied. As a result, the direct application of axial stress will narrow the stent so as not to reflect the phenomenon of migration in vivo.

[0077] Axial forces are negligible in vivo, and this fact is one reason why delivery strategies involving stent diameter reduction due to the application of axial forces can have practical applications. Nevertheless, there are localized axial displacements of the stent relative to the lumen that accumulate over time. The magnitude of these small displacements can be quantified by fixing the stent diameter and measuring the shear force for a specific diameter. To do this, an incompressible cylinder must be introduced into the stent lumen to prevent changes in the stent diameter under axial stress. This is a realistic measure of the stent's properties, as the axial displacement is the sum of small displacements where the radius does not change, and not the effect of large macroscopic axial forces that narrow the stent.

[0078] The experimental setup involves using an artificial lumen, available from manufacturers such as Syndaver, which fits into the target after stent deployment. By varying the diameter of the artificial lumen, various post-deployment RFs can be precisely obtained. Under axial force, the cylindrical shape preventing stent narrowing must fit to the post-delivery radius of the stent, no radial or hoop forces should be applied, and furthermore, it must be incompressible. Various PVC cylinders with diameter increments of 0.5 mm within the expected range of the post-delivery stent diameter may be used.

[0079] As shown in Table 2, the addition of a microstructured surface reduces the need for high post-delivery RF for microstructured stent surfaces. Table 2 compares smooth surfaces and microstructured surfaces with respect to axial shear (slip) for various post-delivery RFs.

[0080] Table 2 Shear force of microstructured stents and smooth stents [Table 2]

[0081] The influence of RF on stent design can be determined by the axial and / or circumferential stent symmetry. In some embodiments, an increase in RF can be achieved by adding struts. On the other hand, the effect of HF may depend solely on axial symmetry in some embodiments. It should be noted that for individual strut elements, RF may be considerably lower than HF. However, in embodiments with cylindrical symmetry, this relationship is reversed, and RF can far exceed HF. This result can be interpreted as the tangential HF vector picking up the radial component of the force when strut elements are joined in a cylindrical geometric shape.

[0082] It should also be noted that in certain embodiments, end flare geometry is not considered. In regions where end flare geometry is utilized, the RF can be substantially increased because the flare induces a radial moment caused by axial bending, in addition to the HF conversion to RF caused by cylindrical geometry. Therefore, in the case of flare stent designs, the ability to reduce delivered RF is constrained, and the reduction in delivered RF obtained by reducing the number of struts and strut diameter is kept to a minimum. Thus, in some embodiments, it may be necessary to eliminate the geometry of the flare stent. The traditional rationale for the geometry of the flare stent is to reduce stent migration after delivery.

[0083] It should also be noted that in some embodiments, greater compression (higher RF) leads to greater hysteresis. In these embodiments, increasing RF by adding struts or increasing the strut diameter can worsen the delivery RF problem by reducing the return of post-delivery RF relative to the delivery RF.

[0084] Surprisingly, in some embodiments, the stent with the highest ratio of post-delivery RF to delivered RF was found to be the stent with the lowest post-delivery RF. This relationship makes the present disclosure particularly useful, as characterized in Table 2. Thus, microstructured stents may be lightweight, easy to deliver, and present less foreign body surface when post-delivery RF requirements are reduced. Post-delivery RF requirements can be largely dependent on migration issues. On the other hand, maintaining an open lumen (the purpose of the stent) can be improved with high HF. The present disclosure provides embodiments of stents that can convert RF to HF while increasing shear force (reducing migration).

[0085] Experiment 3

[0086] Crimping or delivery diameter

[0087] To evaluate the effect of crimping diameter on RF and HF, the hysteresis of embodiments with various radial displacements was studied. RF on the crimping tool and HF within the stent correspond to changes in the slope of the loading curve below an arbitrary threshold radius (8 mm in diameter in these tests). Above the radial threshold, RF can increase dramatically from 8 mm to 18 F (6 mm in diameter). In some embodiments, the emergence of this general nonlinear relationship in the stent between delivery RF and crimp diameter is relatively accompanied by an increase in post-delivery RF. It is possible to increase the delivered RF without significantly increasing it.

[0088] In some embodiments, the cause of nonlinearity may include self-contact between struts. Modifying the usual Hooke's Law relationship between force and displacement can lead to being overwhelmed by a volume relationship involving powers of displacement. In other words, a reaction force in response to compression can be overwhelmed by a contact force without increasing the reaction force itself.

[0089] The larger the percentage difference (as a percentage of the relaxed stent radius) between the radial displacement of the stent when in contact with the lumen and the radial displacement during delivery in some embodiments, the greater the separation of delivered RF from post-delivery RF. Surprisingly, across a wide variety of stent types, the percentage difference is within a relatively narrow range and may therefore be a limiting feature in some embodiments of the stents disclosed herein.

[0090] Table 3 below shows the typical changes in hysteresis according to the delivery diameter.

[0091] Table 3. Delivery diameter for hysteresis [Table 3]

[0092] Experiment 4

[0093] Stent-Lumen Interaction

[0094] In Experiment 4, RF and HF were studied to compare smoothly coated stents with stents coated with microstructures. The microstructured stents were of the type disclosed in Experiment 1.

[0095] In this series of experiments, the study aimed to compare the HF (fracture force) with respect to the stent diameter after a maximum crimping force (CF) of 7.3 N. The deployment force (DF) and equilibrium force (EF) were recorded. It should be noted that arterial stiffness may affect the equilibrium force and equilibrium diameter.

[0096] In the case of a smoothly coated stent, the stent surface may be in contact with the lumen of the ductus arteriosus. As the lumen expands, the stent surface may lose contact with the lumen surface, and therefore, there may be relatively little friction between the lumen surface and the stent surface. This can reduce the shear force.

[0097] Table 3. 30mm stent deployment in sheep aorta (CF=7.3, DF=2.2, EF=0.4) Aortic diameter; Hoop force (smooth); Hoop force (microstructure) [Table 4]

[0098] Experiment 5

[0099] Oversized stent-arterial interaction

[0100] To evaluate the effect of oversized stents on lumen integrity, sheep arteries were used. In Tables 4-7 below, the hoop stress of pressurized (10 kPa) sheep arteries before and after stent placement is given for varying degrees of stent oversizing.

[0101] Table 4. Hoop stress of pressurized (10 kPa) arteries in sheep before and after stent placement in relation to stent oversizing (smoothing). 10kPa = 1.5psi, 1psi = 52mmHg Pressurized diameter, stent oversize, maximum hoop stress (baseline over 10kPa) 27mm 10% 25kPa 25mm 20% 72kPA 23mm 30% 86kPa

[0102] Table 5. Hoop stress of pressurized (10 kPa) arteries in sheep before and after stent placement in relation to stent oversizing (microstructure). 10kPa = 1.5psi, 1psi = 52mmHg Pressurized diameter, stent oversize, maximum hoop stress (baseline over 10kPa) 27mm 10% 37kPa 25mm 20% 102kPA 23mm 30% 115kPa

[0103] Table 6. Deployment and equilibrium of HF and RF for various oversizes (smooth). [Table 5]

[0104] Table 7. Development and equilibrium of HF and RF for various oversizes (microstructures). [Table 6]

[0105] The HF and RF values ​​are recorded in development and equilibrium for various oversizing sizes. In some embodiments, with a 10% oversizing, the maximum physiological stress is exceeded only on the luminal side, and the hoop stress of the artery where the stent is placed doubles the value of the physiological state. In some embodiments, with a 20% oversizing (RF=16N), the maximum physiological stress triples over most of the arterial wall on the luminal side. In some embodiments, oversizing values ​​exceeding 30% (RF=17.5N) significantly exceed the maximum physiological stress along the entire region where the stent is placed, reaching the outer surface of the artery.

[0106] Table 8. Maximum radial stress in the sheep aorta (smooth stent) (N=50 sections) [Table 7]

[0107] Table 9. Hoop stress supported by microstructure in the sheep aorta (N=50 sections) [Table 8]

[0108] A defect is defined as a change in circumference exceeding 1 mm / min over a 2-minute period after the equilibrium time (more than 1 minute after deployment).

[0109] Experiment 6

[0110] Bending test / Bending flexibility

[0111] Bending flexibility is defined by a three-point bending test in which the stent is constrained at its ends six times the stent diameter apart. Flexibility is measured by displacing the center of the stent by a distance equal to the stent diameter. It is the power required to do so.

[0112] Table 10. Bending force (N) [Table 9]

[0113] In some embodiments, stent comfort may depend on minimizing the tendency of the stent to straighten the anatomical structure of the lumen.

[0114] Experiment 7

[0115] Flexibility of peristalsis

[0116] The following experimental setup consists of a swallowing robot with a soft body made of ECOFLEX00-30, having a conduit diameter of 20 mm and a length of 200 mm. The robot was developed to mimic the human swallowing process under various rheological and medical conditions. The robot can generate peristaltic waves with various characteristics similar to those produced when a food bolus moves from the upper to the lower esophagus in humans. Successful passage of the bolus is called effective peristalsis. Ineffective peristalsis is often associated with the residue of food particles remaining in the esophagus.

[0117] The robot consists of 12 equally spaced layers, each containing four pneumatic chambers surrounding the robot's conduit. The purpose of these chambers is to allow each layer to deform sequentially under supplied air pressure, so that the deformation shape can take the form of a peristaltic wave progressing from the top layer to the bottom layer. The swallowing robot (SR) is a realistic example of the biological process of swallowing. Testing such a robot allows for a broader range of measurement and evaluation methods compared to human studies.

[0118] To determine the displacement of the stent from its initial position to its displaced position due to the influence of the movement of peristaltic contraction waves of various characteristics passing through the robot's layers, generated by a pneumatic valve, stent movement experiments were conducted using SR (Simulation Simulation).

[0119] Two different stent dimensions were used: a 23mm x 28mm x 100mm stent and a second 18mm x 21mm x 150mm stent. Displacement data are the average of the two different stent dimensions used in the experiment. The experiment was conducted at near room temperature (36°C). The robot conduit and stent surfaces were kept moist by applying a dry mouth spray before each experiment. A baseline pressure of 25 psi was maintained in all chambers throughout the experiment.

[0120] A velocity of 20 mm / s was used to simulate peristaltic waves. The wave trajectory used to simulate peristalsis was inspired by the pressures typically observed in humans during the same process. The wave profile remained the same throughout all experimental runs, and the amplitude was scaled with coefficients of 1, 1.2, 1.5, and 2. By varying the amplitude, the pressure within the working chamber was varied, and therefore the displacement of adjacent chamber heads with a width of 10 mm was varied. By varying the scaling coefficient, various radial compressive forces of conduit occlusion were achieved and evaluated in comparison to movement under the influence of peristaltic motion.

[0121] Table 11. Stent movement in response to peristaltic cycle at a scaling factor of 1.2. [Table 10]

[0122] Experiment 9

[0123] Shear movement force

[0124] In some embodiments, the radius of a stent may decrease when an axial force is applied while it is in contact with the lumen. This decrease in radius can significantly reduce the radial force, and in some cases, the radial force may approach zero. Therefore, in shear tests in which the stent is axially translated relative to the lumen, the stent must be constrained so as not to decrease in radius without applying further radial force. For stents supplying less than 1 kPa, the stent is oversized to apply an outward radial pressure of 1 kPa (the lumen is reduced by the sutures). Often, this last step is not performed, and the failure to perform this step in such a test should be considered a defective test.

[0125] The experimental setup was as follows, and various stents were deployed and measured using artificial lumens or those recovered from animals. The lumens were fixed to one of the Instron heads (tensile test) at at least four points on one end. The internal volume of the stent was filled with an incompressible lightweight mandrel. The distal end of the stent was connected to the movable head of the Instron, and the lumens were attached to the other head of the Instron in a similar manner.

[0126] The stent was deployed with a target radial force or an oversized radius relative to the relaxed inner diameter of the lumen. The artificial lumen is typically not elastic, but possesses the slippery characteristics of most biological lumens. The elastic characteristics of the natural lumen are usually not important, as sufficient axial displacement to elongate the lumen provides sufficient resistance to migration. Typically, the displacement rate is 0.5–1 N / sec, usually 1 N / sec. As long as the time to migration exceeds 1 second, the results are generally independent of this rate.

[0127] Generally, the force of motion is measured as a 10% time differential discontinuity, but generally, the release that brings about motion is sudden.

[0128] Table 12. Mobility force (low radial force compared to no texture) [Table 11]

[0129] Summary of the experiment

[0130] The experiments described above help establish safe level limits, performance limits, and localization effectiveness using stents utilizing microtextured surfaces. The results demonstrate that in several embodiments having microstructured surfaces, stent localization can be achieved without relying on radial forces to generate the friction required by conventional stents of the prior art. The following portions of this disclosure disclose device specifications that may be useful for embodiments having low radial force stents. Low radial forces can be achieved by the active adhesive function of microstructured surfaces used in applications involving the placement of stents inside hollow body cavities.

[0131] The device specifications given in this section are applicable to all stent placement surgeries currently performed on humans, from coronary stents to esophageal stents. Therefore, specifications are given in percentages (relative parameters) rather than in absolute units. Where an embodiment clearly differs from the range of specifications provided herein, the specifications for that particular embodiment are given. All specifications are given in ranges because tissue characteristics, stent materials, stent geometry, and stent dimensions vary. However, as a result of this study, we have found that many of the disclosed embodiments tend to fall within the given specifications. Furthermore, specifications enabled by microstructured surface stents (low radial force) may differ significantly from those of non-textured stents. These specification differences are clinically significant.

[0132] Radial force Compressive radial force is a device specification that affects usability. Traditional stents of the prior art generally require high radial compressive forces to achieve radial dimensions suitable for deployment in their intended applications, thus presenting several technical challenges. Deployment devices may generally be large to ensure the delivery of the required radial compressive force to properly deliver the stent device to the target location. The compression profile is also important. The compression profile can be obtained by measuring the compressive radial force of at least two radial reductions. Typically, the stent radius can be reduced by about 10% to 20% in some embodiments. However, if the radial compression is too large, in some embodiments there is a risk of permanently deforming the stent, reducing its effectiveness or rendering it ineffective.

[0133] Stent specification 1, compression profile

[0134] In some embodiments, assuming that the radial pressure at ΔR=0 is 0 Newtons (N), if the radial pressure at ΔR=10% is X, then the radial force at ΔR=20% may be in the range of 1.10X to 1.5X, preferably 1.10X to 1.25X.

[0135] In this technical field, it will be understood that radial compression is often reported in units of force. Using units of force may require certain assumptions about the dimensions of the stent. The conversion of force from Newtons to Pascals is 1 g / cm². 2 = 100 Pa, and the area is calculated from A = dlπ, where "d" is the diameter of the stent and "l" is the length of contact between the stent and the body lumen.

[0136] Some modern stent compression profiles are 2.0 or higher. This may be acceptable if ΔR=10% generates a pressure of less than 1.8 kPa. However, a compression profile of 2.0 means that ΔR=20% is unnecessarily high. While this does not affect the patient, it may impose high demands on the design of the deployment device. It also significantly increases stent hysteresis, which can lead to a chain reaction of adverse effects on both the design and the patient. An ideal stent should have a compression profile of approximately 1.

[0137] Table 13. Based on the capabilities of the deployment device. [Table 12]

[0138] Stent specification 2, maximum radial compression force

[0139] In some embodiments, the maximum radial compressive force of a low radial force stent may be 2.7 kPa to 1.5 kPa, preferably 2.26 kPa to 1.50 kPa, and more preferably 2.0 kPa to 1.5 kPa.

[0140] Stent specification 3, radial expansion profile

[0141] In some embodiments, assuming that the radial pressure at ΔR=0 is 0 Newtons (N), if the radial pressure at ΔR=10% is X, then the radial force at ΔR=20% is in the range of 1.09X to 1.42X, preferably 0.91X to 1.18X, and more preferably 0.91X to 1.10X.

[0142] In certain embodiments, the stent needs to exhibit no change in radial force between compression and expansion. As mentioned earlier, this difference is called hysteresis.

[0143] Table 14. Based on tissue necrosis and comfort. [Table 13]

[0144] Stent specification 4, maximum radial expansion

[0145] In some embodiments, the maximum radial expansion force of a low radial force stent may be 1.82 kPa to 1.00 kPa, preferably 1.62 kPa to 1.00 kPa, and more preferably 1.38 kPa to 1.00 kPa.

[0146] Axial force Stent specification 5, minimum peristaltic shear force (comfort)

[0147] In some embodiments, it will be understood that desired specifications of this specification may not be met by conventional stents of the prior art. Embodiments of low radial force stents of this disclosure can resist movement by finding an equilibrium position of about 20 to 32 psi (1.2 times), preferably between 25 and 32 psi, and more preferably between 29 and 32 psi.

[0148] Stent specification 6, minimum peristaltic shear force (maximum force)

[0149] In some embodiments, it will be understood that desired specifications of this specification may not be met by conventional stents of the prior art. Embodiments of low radial force stents of this disclosure can resist movement by finding an equilibrium position of about 20 to 35 psi, preferably 25 to 32 psi, and more preferably 29 to 32 psi.

[0150] This test will be understood to apply both axial (shear) and radial forces in a natural combination. The easily measurable force is the radial force applied to the peristaltic actuator. The shear stress can be derived from this pressure and many of the characteristics of the peristaltic motion defined above. Therefore, the minimum peristaltic shear force can be given by the radial pressure applied to the 10 mm actuator.

[0151] Table 15. Minimum peristaltic shear force [Table 14]

[0152] Stent specification 7, minimum shear force

[0153] Certain embodiments of the stent of this disclosure may include a shear force greater than 3.0 N with a radial expansion pressure of less than 1 kPa, preferably greater than 4.75 N, and more preferably greater than 5.0 N. (Experiment 9)

[0154] Stent specification 8, maximum bending force

[0155] Certain embodiments of the stents of this disclosure may include bending forces of 105 to 55 N, preferably 96 to 55 N, and more preferably 67 to 55 N.

[0156] Bending force may also be a combination of axial and radial forces, or a measure of flexibility. See Experiment 6 (above). In some embodiments, stent comfort may depend on minimizing the tendency of the stent to straighten the anatomical structure of the lumen.

[0157] Hoop power Stent specification 9, maximum hoop stress (compression)

[0158] In some embodiments of this disclosure, the compressive hoop force may be less than 8.5%, preferably less than 7.8%, and more preferably less than 6.5%, of the radial force in compression at a 50% reduction in stent diameter. (Experiment 1)

[0159] Stent specification 10, maximum hoop stress (extension)

[0160] In some embodiments of this disclosure, the extension hoop force may be less than 8.1%, preferably less than 7.3%, and more preferably less than 6.1%, of the radial force during expansion, after a 50% reduction in stent diameter following a 55% compression. (Experiment 1)

[0161] Hysteresis Stent specification 11, maximum hoop hysteresis

[0162] In some embodiments of this disclosure, the hysteresis ([compressive force - tensile force] / compressive force) may be less than 35%, preferably less than 25%, and more preferably less than 20%, at a 20% reduction in stent radius after a 25% reduction. (Experiment 1)

[0163] Stent specification 12, maximum radial hysteresis

[0164] In some embodiments of this disclosure, the hysteresis ([compressive force - tensile force] / compressive force) may be less than 37%, preferably less than 31%, and more preferably less than 25%, at a 20% reduction in stent radius after a 25% reduction. (Experiment 1)

[0165] Stent specification 13, post-delivery hysteresis shear force

[0166] Comparing specific embodiments of the untextured and textured stents of this disclosure with each other (identical stent designs except for texture), the textured stent may have at least 10 times, preferably at least 15 times, and more preferably at least 20 times, the shear force of the untextured design when the stent is compressed and then unfolded (Experiment 2).

[0167] Stent specification 14, post-delivery hysteresis shear force

[0168] Comparing specific embodiments of the untextured and textured stents of this disclosure (with identical stent designs except for texture), the textured stent may have a shear force at least 20 times greater than that of the untextured stent when deployed with a post-delivery radial force of 2 kPa. (Experiment 2)

[0169] Stent specification 15, post-delivery hysteresis shear force

[0170] Comparing specific embodiments of the untextured and textured stents of this disclosure (with identical stent designs except for texture), the textured stent will have at least 10 times the shear force of the untextured stent when deployed with a post-delivery radial force of 20 kPa. (Experiment 2)

[0171] Stent specification 16, post-delivery hysteresis shear force

[0172] Comparing specific embodiments of the untextured and textured stents of this disclosure (with identical stent designs except for texture), the textured stent may have a higher relative shear force ratio (shear force of this disclosure / untextured shear force) at 2 kPa compared to 20 kPa. (Experiment 2)

[0173] Stent specification 17, anterior-posterior hysteresis

[0174] In some embodiments having the stent of the present disclosure, the stent may undergo hysteresis of less than 5%, preferably less than 3%, and more preferably less than 1%, when compressed by 50% and returned to its pre-compressed stent size.

[0175] It should be understood that a useful baseline for hysteresis is to measure the compressive force from the starting diameter to the minimum diameter (metal-to-metal contact), the compressive force returning to the original diameter, and the change in compressive force (pre-force) from the expansion force (post-force). In conventional stents of the prior art, the post-force is often smaller than the pre-force, which causes serious problems when the stent size is to fit the lumen size. Clinicians typically assume some hysteresis and adjust the stent size accordingly. Therefore, there is a need for a stent that returns to its designed size with minimal hysteresis loss, thereby enabling a precise fit of the stent to the lumen size. This advantage in some embodiments of the stents disclosed herein can occur when the ends of the stent remain open due to a ring of textured coating. Unlike metal struts, plastic coatings may not undergo the same permanent deformation as metals. Furthermore, the textured coating may supply its own radial force by pulling radially. When these effects are combined, the results of Experiment 3 are obtained.

[0176] Table 16. Anterior-posterior hysteresis [Table 15]

[0177] The following examples and embodiments are for illustrative purposes only and are not intended to limit the scope of this work.

[0178] Example 1

[0179] An embodiment of a stent with improved contact area.

[0180] One of the main causes of chronic esophageal irritation resulting from esophageal stent placement is the flared structure at the end of the stent, which can be found in the prior art. The flare can have a radius of up to twice that of the stent body. Flares can be a major design limitation because flared stents are more difficult to compress to delivery size than stents without flares. Flares on prior art esophageal stents are provided solely to increase the radial force of the stent and reduce stent migration. Embodiments of the low radial force stent of this disclosure increase shear force without increasing radial force and discomfort. Higher shear stress can reduce the area of ​​the stent that needs to be in contact with the esophagus.

[0181] Referring to Figure 4, a stent 400 for keeping a lumen 402 having a constriction 404 open comprises a constriction-expansion element 406, an elastomer lumen 408, and two fixing rings 410. The elastomer lumen 408 has a ring-shaped microstructure 409 of the present disclosure around each end of the soft elastomer lumen 408. The constriction-expansion element 406 comprises a wire mesh having a short axial length but sufficient radial force to expand the constriction 404. The elastomer lumen 408 is a continuously smooth surface tube made of silicone or a similar material. The fixing rings 410 are slightly inclined so that they can be compressed along their long axis. As depicted in Figure 4, the constriction-expansion element 406 is initially positioned in the region of the constriction 404, and then the elastomer lumen 408 is positioned such that the expansion element is located midway along the axial length of the soft elastomer lumen 408. The two ends of the elastomer lumen 408 are fixed using the fixing ring 410 by compressing each end to a smaller width along its long axis and fixing the lumen ends.

[0182] Example 2

[0183] Embodiment of a low-compression profile stent

[0184] Referring to Figure 5, a low-compression profile stent is a standard stent with the same intended placement radius. The stent 500 has a maximum radius 502 that is less than 50% of the radius of the stent 504. The radius of the end 506 is the same as the radius of the intermediate 508. The stent is coated with an elastomer 510 having a spiral ribbon of microstructure 511.

[0185] Example 3

[0186] Embodiment of a stent having a low radial expansion profile

[0187] Referring to Figure 6, the low radial expansion force profile stent 600 comprises a helix 602 coated with a soft polymer 604 having a microstructured texture 605. The helix 602 is typically elongated with a radius smaller than the intended implantation radius. The helix 602 has a band 606 having a distally oriented ratchet texture 608. The band 606 is attached distally to the helix 602 at its end 610. The band 606 passes through a hub 612 having a ratchet in the proximal direction. When the stent 600 is implanted, the band 606 is pulled proximal, and the ratchet texture 608 on the band 606 engages with the ratchet in the hub 614, shortening the axial length of the stent 600. The helix 602 expands and fits snugly into the target lumen. The band 606 is made of durable plastic. During the procedure, the stent can be relaxed by cutting band 606. The radial expansion profile is in the range of 1.09X to 1.42X.

[0188] Example 4

[0189] Embodiment of a stent with reduced contact area

[0190] Referring to Figure 7, a metal wire woven stent 700 having a central bulge 702 is implanted in a lumen 704 having a constriction 706. The bulge 702 is positioned proximal to the constriction 706. The stent is coated with a soft polymer 708. A microstructured texture 710 is placed in the bulge 702 region of the coating. The stent 700 has a maximum shear force exceeding 3.0 N with radial expansion of less than 1 kPa.

Claims

1. An implantable device for placement within a body lumen, A portable device comprising a tubular member having an outer surface and an inner surface, wherein the outer surface comprises a substrate, the substrate further comprising at least one first hierarchical microstructure pattern and at least one second hierarchical microstructure pattern, the at least one first hierarchical microstructure pattern and the at least one second hierarchical microstructure pattern are arranged to generate a Wenzel-Cassie state.

2. The portable device according to claim 1, wherein the at least one first hierarchical microstructure pattern and the at least one second hierarchical microstructure pattern comprises a two-dimensional sinusoidal microstructure, a pillar microstructure, or a fluted microstructure.

3. The portable device according to claim 1, wherein the substrate further comprises at least one third hierarchical microstructure pattern.

4. The implantable device according to claim 3, wherein the at least one third hierarchical microstructure pattern is a grooved microstructure.

5. The portable device according to claim 1, wherein the at least one first hierarchical microstructure pattern includes dimensions 1.1 to 10 times the dimensions of the at least one second hierarchical microstructure pattern.

6. The portable device according to claim 3, wherein the at least one second hierarchical microstructure pattern includes dimensions 1.1 to 10 times the dimensions of the at least one third hierarchical microstructure pattern.

7. The implantable device according to claim 2, wherein the pillar microstructure has an elliptical or polygonal cross-section.

8. The implantable device according to claim 1, wherein the tubular member includes a maximum radial compression pressure in the range of 1.82 kPa to 1.00 kPa.

9. The implantable device according to claim 1, wherein the substrate has an adhesive surface texture.

10. The portable device according to claim 1, wherein the substrate has a slippery surface texture.

11. The implantable device according to claim 1, wherein the substrate comprises a nitinol strut.

12. The implantable device according to claim 11, wherein the tubular member includes a nitinol coating having a texture hysteresis curve.

13. The implantable device according to claim 1, wherein the tubular member further comprises a constriction and expansion element, a soft elastomer lumen, and at least one fixing ring.

14. The implantable device according to claim 13, wherein the constriction and expansion element comprises a wire mesh.

15. The tubular member includes an elastomer coating, as described in claim 1. vinegar.

16. The implantable device according to claim 15, wherein at least one microstructure pattern is arranged in a helical ribbon pattern.

17. The implantable device according to claim 1, wherein the tubular member comprises a metal woven wire, and the metal woven wire further comprises a bulging portion.

18. The implantable device according to claim 17, wherein the tubular member includes a soft polymer coating.

19. The implantable device according to claim 18, wherein the at least one first hierarchical microstructure pattern and / or the at least one second hierarchical microstructure pattern are arranged on the soft polymer coating.

20. An implantable device for placement within a body lumen, A portable device comprising a tubular member having an outer surface and an inner surface, wherein the outer surface comprises a substrate, the substrate further comprising at least one first hierarchical microstructure pattern, at least one second hierarchical microstructure pattern, and at least one third hierarchical microstructure pattern, the at least one first hierarchical microstructure pattern, the at least one second hierarchical microstructure pattern, and the at least one third hierarchical microstructure pattern are arranged to generate a Wenzel-Cassie state.

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